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Nanoparticulates As Drug Carriers

Northeastern University, USA 

2 Nanoparticle Engineering 30 
2.1 Drug release mechanisms 32 
3 Site-specific Targeting with Nanoparticles: Importance of Size 
and Surface Properties 33 
4 Conclusions 37 
References 38 
4. Genetic Vaccines: A Role for Liposomes 43 
Gregory Gregoriadis, Andrew Bacon, Brenda McCormack and Peter Laing 
1 Introduction 43 
2 The DNA Vaccine 44 
3 DNA Vaccination via Liposomes 45 
3.1 Procedure for the entrapment of plasmid DNA into liposomes 46 
3.2 DNA immunization studies 47 
3.3 Induction of a cytotoxic T lymphocyte (CTL) response by 
liposome-entrapped plasmid DNA 50 
4 The Co-delivery Concept 51 
References 53 
5. Polymer Micelles as Drug Carriers 57 
Elena V. Batrakova, Tatiana K. Bronich, Joseph A. Vetro and 
Alexander V. Kabanov 
1 Introduction 57 
2 Polymer Micelle Structures 58 
2.1 Self-assembled micelles 58 
2.2 Unimolecular micelles 61 
2.3 Cross-linked micelles 62 
3 Drug Loading and Release 63 
3.1 Chemical conjugation 63 
3.2 Physical entrapment 64 
3.3 Polyionic complexation 66 
4 Pharmacokinetics and Biodistribution 68 
5 Drug Delivery Applications 72 
5.1 Chemotherapy of cancer 72 
5.2 Drug delivery to the brain 76 
5.3 Formulations of antifungal agents 77 
5.4 Delivery of imaging agents 77 
5.5 Delivery of polynucleotides 78 
Contents xv 
6 Clinical Trials 79 
7 Conclusions 79 
References 80 
6. Vesicles Prepared from Synthetic Amphiphiles  Polymeric Vesicles 
and Niosomes 95 
Ijeoma Florence Uchegbu and Andreas G. Schatzlein 
1 Introduction 95 
2 Polymeric Vesicles 96 
2.1 Polymer self assembly 97 
2.2 Polymers bearing hydrophobic pendant groups 98 
2.3 Block copolymers 101 
2.4 Preparing vesicles from self-assembling polymers 102 
2.5 Self assembling polymerizable monomers 103 
3 Polymeric Vesicle Drug Delivery Applications 104 
3.1 Drug targeting 104 
3.2 Gene delivery 105 
3.3 Responsive release 106 
3.3.1 pH 106 
3.3.2 Enzymatic 106 
3.3.3 Magnetic 107 
3.3.4 Oxygen 108 
4 Non-ionic Surfactant Vesicles (Niosomes) 108 
4.1 Self assembly 108 
4.2 Polyhedral vesicles and giant vesicles (Discomes) Ill 
4.3 Vesicle preparation 113 
5 Niosome Delivery Applications 113 
5.1 Drug targeting 113 
5.1.1 Anti cancer drugs 113 
5.1.2 Anti infectives 115 
5.1.3 Delivery to the brain 115 
5.2 Topical use of niosomes 116 
5.2.1 Transdermal 116 
5.2.2 Ocular 116 
5.3 Niosomal vaccines 116 
5.4 Niosomes as imaging agents 117 
6 Conclusions 117 
References 117 
xv i Contents 
7. Recent Advances in Microemulsions as Drug Delivery Vehicles 125 
M Jayne Lawrence and Warankanga Warisnoicharoen 
1 Definition 125 
1.1 Microemulsion versus an emulsion 125 
1.2 Microemulsion versus a nanoemulsion 126 
1.3 Microemulsions 128 
1.4 Microemulsions, swollen micelles, micelles 129 
1.5 Microemulsions and cosolvent systems 130 
2 Microemulsions as Drug Delivery Systems 130 
2.1 Self-emulsifying drug delivery systems (SEDDS) 131 
2.2 Related systems 133 
2.2.1 Microemulsion gels 133 
2.2.2 Double or multiple microemulsions 134 
2.3 Processed microemulsion formulations 134 
2.3.1 Solid state or dry emulsions 134 
3 Formulation 135 
3.1 Surfactants and cosurfactants 136 
3.2 Oils 138 
3.3 Characterization 139 
4 Routes of Administration 139 
4.1 Oral 139 
4.1.1 Proteins and peptides 140 
4.1.2 Other hydrophilic molecules 141 
4.1.3 Hydrophobic drugs 142 
4.2 Buccal 144 
4.3 Parenteral 144 
4.3.1 Long circulating microemulsions 147 
4.3.2 Targeted delivery 148 
4.4 Topical delivery 148 
4.4.1 Dermal and transdermal delivery 148 
4.5 Ophthalmic 154 
4.6 Vaginal 156 
4.7 Nasal 157 
4.8 Pulmonary 158 
4.8.1 Antibacterials 159 
5 Conclusion 160 
References 160 
Contents xvii 
8. Lipoproteins as Pharmaceutical Carriers 173 
Suwen Liu, Shining Wang and D. Robert Lu 
1 Introduction 173 
2 The Structure of Lipoproteins 174 
3 Chylomicron as Pharmaceutical Carrier 175 
4 VLDL as Pharmaceutical Carrier 176 
5 LDL as Pharmaceutical Carrier 177 
5.1 LDL as anticancer drug carriers 178 
5.2 LDL as carriers for other types of bioactive compounds . . . .179 
5.3 LDL for gene delivery 179 
6 HDL as Pharmaceutical Carriers 179 
7 Cholesterol-rich Emulsions (LDE) as Pharmaceutical Carriers . . . .180 
8 Concluding Remark 181 
References 182 
9. Solid Lipid Nanoparticles as Drug Carriers 187 
Karsten Mader 
1 Introduction: History and Concept of SLN 187 
2 Solid Lipid Nanoparticles (SLN) Ingredients and Production . . . .188 
2.1 General ingredients 188 
2.2 SLN preparation 189 
2.2.1 High shear homogenization and ultrasound 189 
2.3 High pressure homogenization (HPH) 189 
2.4 Hot homogenization 190 
2.5 Cold homogenization 190 
2.5.1 SLN prepared by solvent emulsification / 
evaporation 191 
2.5.2 SLN preparations by solvent injection 191 
2.5.3 SLN preparations by dilution of microemulsions or 
liquid crystalline phases 192 
2.6 Further processing 193 
2.6.1 Sterilization 193 
2.6.2 Drying by lyophilization, nitrogen purging and 
spray drying 194 
3 SLN Structure and Characterization 196 
4 The "Frozen Emulsion Model" and Alternative SLN Models . . . . 200 
5 Nanostructured Lipid Carriers (NLC) 201 
6 Drug Localization and Release 202 
xviii Contents 
7 Administration Routes and In Vivo Data 203 
8 Summary and Outlook 205 
References 205 
10. Lipidic Core Nanocapsules as New Drug Delivery Systems 213 
Patrick Saulnier and Jean-Pierre Benoit 
1 Introduction 213 
2 Lipidic Nanocapsule Formulation and Structure 215 
2.1 Process 215 
2.2 Influence of the medium composition 216 
2.3 Structure and purification of the LNC by dialysis 217 
2.4 Imagery techniques 218 
3 Electrical and Biological Properties 219 
3.1 Electro kinetic comportment 219 
3.2 Evaluation of complement system activation 220 
4 Pharmacokinetic Studies and Biodistribution 220 
5 Drug Encapsulation and Release 222 
5.1 Ibuprofene 222 
5.2 Amiodarone 223 
6 Conclusions 223 
References 224 
11. Lipid-Coated Submicron-Sized Particles as Drug Carriers 225 
Evan C. linger, Reena Zutshi, Terry O. Matsunaga and Rajan Ramaswami 
1 Technology 225 
2 Ultrasound Contrast Agents 228 
3 Sonothrombolysis ^r_._ 232 
4 Clinical Studies 237 
5 Blood Brain Barrier 239 
6 Drug Delivery 242 
6.1 Targeted bubbles 242 
6.2 Targeted submicron-sized droplets 244 
7 Gene Delivery 245 
8 Oxygen Delivery 247 
9 Pulmonary Delivery 248 
10 Conclusion 249 
References 250 
Contents xix 
Nanocapsules: Preparation, Characterization and Therapeutic 
Applications 255 
Ruxandra Grefand Patrick Couvreur 
1 Introduction 255 
2 Preparation 257 
2.1 Nanocapsules obtained by interfacial polymerization 257 
2.1.1 Oil-containing nanocapsules 257 
2.1.2 Nanocapsules containing an acqueous core 259 
2.2 Nanocapsules obtained from preformed polymers 261 
3 Characterization 263 
4 Drug Release 265 
5 Applications 266 
5.1 Oral route 266 
5.2 Parenteral route 267 
5.3 Ocular delivery 269 
6 Conclusion 270 
References 271 
Dendrimers as Nanoparticulate Drug Carriers 277 
Sbnke Svenson and Donald A. Tomalia 
1 Introduction 277 
2 Nanoscale Containers  Micelles, Dendritic Boxes, Dendrophanes, 
and Dendroclefts 279 
2.1 Dendritic micelles 279 
2.2 Dendritic box (Nano container) 280 
2.3 Dendrophanes and dendroclefts 282 
3 Dendrimers in Drug Delivery 282 
3.1 Cisplatin 283 
3.2 Silver salts 285 
3.3 Adriamycin, methotrexate, and 5-fluorouracil 285 
3.4 Etoposide, mefenamic acid, diclofenac, and venlafaxine . . . . 286 
3.5 Ibuprofen, indomethacin, nifedipine, naproxen, paclitaxel, 
and methylprednisolone 287 
3.6 Doxorubicin and camptothecin  self-immolative dendritic 
prodrugs 289 
3.7 Photodynamic therapy (PDT) and boron neutron capture 
therapy (BNCT) 291 
4 Nano-Scaffolds for Targeting Ligands 292 
4.1 Folic acid 292 
4.2 Carbohydrates 293 
4.3 Antibodies and biotin-avidin binding 294 
4.4 Penicillins 295 
5 Dendrimers as Nano-Drugs 295 
6 Routes of Application 296 
7 Biocompatibility of Dendrimers 297 
8 Conclusions 299 
References 299 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly 
Soluble Drugs 307 
Rainer H. Muller and Jens-Uive A. H. Junghanns 
1 Introduction 307 
2 Definitions 308 
3 Physicochemical Properties of Drug Nanocrystals 309 
3.1 Change of dissolution velocity 309 
3.2 Saturation solubility 309 
3.3 Does size really matter? 311 
3.4 Effect of amorphous particle state 312 
4 Production Methods 313 
4.1 Precipitation methods 313 
4.1.1 Hydrosols 313 
4.1.2 Amorphous drug nanoparticles (NanoMorph) . . . .313 
4.2 Homogenization methods 314 
4.2.1 Microfluidizer technology 314 
4.2.2 Piston-gap homogenization in water (Dissocubes) . . 314 
4.2.3 Nanopure technology 315 
4.3 Combination Technologies 315 
4.3.1 Microprecipitation and High Shear Forces 
4.3.2 Nanopure XP technology 316 
5 Application Routes and Final Formulations 317 
5.1 Oral administration 317 
5.2 Parenteral administration 319 
5.3 Miscellaneous administration routes 321 
6 Nanosuspensions as Intermediate Products 322 
Contents xxi 
7 Perspectives 324 
References 324 
Cells and Cell Ghosts as Drug Carriers 329 
Jose M. Lanao and M. Luisa Sayalero 
1 Introduction 329 
2 Bacterial Ghosts 329 
2.1 Application of bacterial ghosts as a delivery system 331 
3 Erythrocyte Ghosts 333 
3.1 Applications of erythrocyte ghosts as a delivery system . . . .335 
4 Stem Cells 338 
5 Polymorphonuclear Leucocytes 340 
6 Apoptopic Cells 340 
7 Tumor Cells 340 
8 Dendritic Cells 341 
9 Conclusions 341 
References 342 
Cochleates as Nanoparticular Drug Carriers 349 
Leila Zarif 
1 Introduction 349 
2 Cochleates Nanoparticles in Oral Delivery 350 
2.1 Cochleate structure 350 
2.2 Cochleate preparation 350 
2.2.1 Which phospholipid and which cation to use? 350 
2.2.2 Which molecules can be entrapped in cochleates 
nanoparticles 352 
2.2.3 Multiple ways of preparing cochleates 353 
2.3 Cochleates as oral delivery system for antifungal agent, 
amphotericin B 355 
2.3.1 In candidiasis animal model 355 
2.3.2 In aspergillosis animal model 355 
2.3.3 In cryptococcal meningitis animal model 357 
2.3.4 Toxicity of amphotericin B cochleates 357 
2.3.5 Pharmacokinetics of amphotericin B cochleates . . . . 357 
2.4 Other potential applications for cochleates 359 
2.4.1 Cochleate for the delivery of antibiotics 359 
2.4.2 Delivery of clofazimine 360 
xxii Contents 
2.4.3 Delivery of tobramycin 360 
2.4.4 Cochleate for the delivery of anti-inflammatory 
drugs 361 
2.5 Other uses of cochleates 361 
3 Conclusion 361 
References 362 
17. Aerosols as Drug Carriers 367 
N. Renee Labiris, Andrew P. Bosco and Myrna B. Dolovich 
1 Introduction 367 
2 Pulmonary Drug Delivery Devices 369 
2.1 Nebulizers 369 
2.2 Metered-dose inhalers 371 
2.3 Dry powder inhalers 373 
3 Aerosol Particle Size 373 
4 Targeting Drug Delivery in the Lung 376 
5 Clearance of Particles from the Lung 378 
5.1 Airway geometry and humidity 378 
5.2 Lung clearance mechanisms 379 
6 Nanoparticle Formulations for Inhalation 381 
6.1 Diagnostic imaging 382 
6.2 Vaccine delivery 383 
6.3 Anti Tuberculosis therapy 385 
6.4 Gene therapy 386 
7 Conclusion 388 
References 388 
18. Magnetic Nanoparticles as Drug Carriers 397 
Urs O. Hafeli and Mathieu Chastellain 
1 Introduction 397 
2 Definitions 398 
2.1 Properties of magnetic materials 398 
2.2 Nanoparticles 400 
3 Magnetic Nanoparticles 401 
3.1 Iron oxide based magnetic nanoparticles 401 
3.2 Cobalt based magnetic nanoparticles 402 
3.3 Iron based magnetic particles 402 
3.4 Encapsulated magnetic nanoparticles 403 
3.5 Biocompatibility issues of magnetic nanoparticles 403 
Contents xxiii 
4 Application of Magnetic Nanoparticles as Drug Carriers 404 
4.1 Magnetic hyperthermia 405 
4.2 Magnetic chemotherapy 406 
4.3 Other magnetic treatment approaches 408 
4.4 Magnetic gene transfer 409 
5 Conclusions 410 
References 411 
19. DQAsomes as Mitochondria-Specific Drug and DNA Carriers 419 
Volkmar Weissig 
1 Introduction 419 
2 The Self Assembly Behavior of Bis Quinolinium Derivatives 420 
2.1 Monte Carlo computer simulations 420 
2.2 Physico-chemical characterization 421 
2.3 Structure activity relationship studies 422 
3 DQAsomes as Mitochondrial Transfection Vector 424 
4 DQAsomes as Carriers of Pro-apoptotic Drugs 429 
5 Summary 432 
References 432 
20. Liposomal Drug Carriers in Cancer Therapy 437 
Alberto A. Gabizon 
1 Introduction 437 
2 The Challenge of Cancer Therapy 439 
3 The Rationale for the Use of Liposomal Drug Carriers in Cancer . . 442 
4 Liposome Formulation and Pharmacokinetics  Stealth 
Liposomes 445 
5 Preclinical Observations with Liposomal Drug Carriers 
in Tumor Models 448 
6 Liposomal Anthracyclines in the Clinic 449 
6.1 Doxil 450 
6.2 Myocet 454 
6.3 Daunoxome 454 
7 Clinical Development of Other Liposome-entrapped 
Cytotoxic Agents 455 
8 The Future of Liposomal Nanocarriers 456 
References 457 
xxiv Contents 
21. Nanoparticulate Drug Delivery to the Reticuloendothelial System 
and to Associated Disorders 463 
Mukul Kumar Basu and Sanchaita Lala 
1 Introduction 463 
2 Reticuloendothelial System and Associated Disorders 464 
3 Uptake of Nanoparticles by the Reticuloendothelial System 464 
3.1 Sites of uptake 464 
3.2 Mechanism of uptake 465 
3.3 Factors influencing uptake 468 
3.4 Role of surface modifications on uptake 469 
4 Active Targeting of Nanoparticles by Receptor Mediated 
Endocytosis 471 
5 Application in Chemotherapy 473 
6 Summary 475 
References 477 
22. Delivery of Nanoparticles to the Cardiovascular System 481 
Ban-An Khazv 
1 Introduction 481 
2 Targeting the Myocardium with Immunoliposomes 481 
3 Other Nanoparticle-Targeting of the Cardiovascular System 484 
4 Novel Application of Nano-Immunoliposomes 485 
5 CSIL as Targeted Gene or Drug Delivery 492 
6 Conclusion 495 
References 496 
23. Nanocarriers for the Vascular Delivery of Drugs to the Lungs 499 
Thomas Dziubla and Vladimir Muzykantov 
1 Introduction 500 
2 Biomedical Aspects of Drug Delivery to Pulmonary Vasculature . . 500 
2.1 Routes for pulmonary drug delivery: Intratracheal vs 
vascular 501 
2.2 Pulmonary vasculature as a target for drug delivery 501 
3 Pulmonary Targeting of Nanocarriers 503 
3.1 Effects of carrier size on circulation and tissue distribution . .503 
Contents xxv 
3.2 Passive targeting 505 
3.2.1 Mechanical retention 505 
3.2.2 Charge-mediated retention and non-viral gene 
delivery 506 
3.2.3 Pulmonary enhanced permeation-retention (EPR) 
effect 507 
3.3 Active targeting 507 
4 Carrier Design 509 
4.1 Biocompatibility 509 
4.2 Material selection (by application) 510 
4.2.1 Imaging 510 
4.2.2 Gene delivery 510 
4.2.3 Delivery of therapeutic enzymes 511 
4.2.4 Small molecule drugs 512 
4.3 Types of nanocarriers 512 
4.4 Mechanisms of drug loading 512 
4.5 Drug release mechanisms 515 
4.6 Nanocarriers for active targeting 516 
5 Conclusion: Safety Issues, Limitations and Perspectives 517 
References 518 
24. Nanoparticulate Carriers for Drug Delivery to the Brain 527 
Jorg Kreuter 
1 Introduction 527 
2 Nanoparticles 528 
3 Biodistribution 530 
3.1 Influence of surfactants on the biodistribution of 
nanoparticles 530 
3.2 Influence of PEGylation on the biodistribution of 
nanoparticles 532 
4 Pharmacology 534 
5 Brain Tumors 536 
6 Toxicology 538 
7 Mechanism of the Delivery of Drug Across the Blood-Brain 
Barrier with Nanoparticles 539 
8 Summary 541 
9 Conclusions 542 
References 542 
Nanoparticles for Targeting Lymphatics 549 
William Phillips 
1 Introduction 549 
1.1 The lymphatic vessels 550 
1.2 Lymph nodes 551 
2 Potential for Nanoparticles for Drug Delivery to Lymphatics . . . . 553 
3 Importance of Lymph Nodes for Disease Spread and 
Potential Applications of Lymph Node Drug Delivery 554 
3.1 Cancer 554 
3.2 HIV 555 
3.3 Filaria 555 
3.4 Anthrax 556 
3.5 Tuberculosis 556 
3.6 Importance of lymph node antigen delivery for development 
of an immune response 557 
4 Factors Influencing Nanoparticle Delivery to Lymph Nodes 559 
4.1 Nanoparticle size 559 
4.2 Nanoparticle surface 559 
4.3 Effect of massage on lymphatic clearance of subcutaneously 
injected liposomes 560 
4.4 Macrophage phagocytosis 561 
4.5 Fate of nanoparticles in lymph nodes 561 
5 Nanoparticle Diagnostic Imaging Agents for Determining Cancer 
Status of Lymph Nodes 561 
5.1 Subcutaneous injection of iodinated nanoparticles for 
computed tomography imaging 561 
5.2 Subcutaneous and intraorgan injection of magnetic 
resonance (MRI) contrast agents 563 
5.3 Intravenous injection of magnetic nanoparticles for 
MRI imaging 563 
5.4 Nanoparticle diagnostic agents for localizing the sentinel 
lymph node 565 
5.5 Radiolabeled nanoparticles for sentinel lymph node 
identification 566 
5.6 99mTc-Colloidal nanoparticles for sentinel node identification . 566 
5.7 Optical 568 
5.8 Ultrasound nanobubbles 569 
6 Recently Introduced Medical Imaging Devices for Monitoring 
Lymph Node Delivery and Therapeutic Response 569 
Contents xxvii 
7 Nanoparticle Lymph Node Drug Delivery 571 
7.1 Confusion in reporting lymph node delivery 571 
7.2 Calculation of lymph node retention efficiency 573 
8 Specific Types Nanoparticles for Lymph Node Targeting 573 
8.1 PLGA nanoparticles 573 
8.2 Micelles 574 
8.3 Liposomes 574 
9 Avidin Biotin-Liposome Lymph Node Targeting Method 577 
10 Massage and the Avidin-Biotin Liposome Targeting Method 578 
11 Nanoparticles for Lymph Node Anti-Infectious Agent Delivery . . . 580 
12 Liposomes for Intraperitoneal Lymph Node Drug Delivery 581 
12.1 Intraperitoneal liposome encapsulated drugs 582 
12.2 Effect of liposome size on intraperitoneal clearance 583 
12.3 Avidin/Biotin-liposome system for intraperitoneal and 
lymph node drug delivery 584 
12.4 Mediastinal lymph node drug delivery with avidin-biotin 
system by intrapleural injection 585 
12.5 Avidin biotin for diaphragm and mediastinal lymph node 
targeting 586 
13 Nanoparticles for Cancer Therapy 587 
13.1 Intralymphatic drug delivery to lymph nodes 587 
13.2 Nanoparticles for treatment of metastatic lymph nodes of 
upper GI malignacies 589 
13.3 Lessons from endolymphatic radioisotope therapy 591 
14 Advantages of Nanoparticles for Lymphatic Radiotherapy 592 
15 Intraoperative Radiotherapy for Positive Tumor Margins 
and for Treatment of Lymph Nodes 593 
16 Potential of Using Radiolabeled Nanoparticles for Intra tumoral 
Radionuclide Therapy 593 
17 Liposome Pharmacokinetics after Intra tumoral Administration . . .595 
18 Rhenium-Labeled Liposomes for Tumor Therapy 595 
19 Nanoparticles for Immune Modulation 597 
20 Conclusions 598 
References 598 
26. Polymeric Nanoparticles for Delivery in the Gastro-Intestinal Tract 609 
Mayank D. Bhavsar, Dinesh B. Shenoy and Mansoor M. Amiji 
1 Oral Drug Delivery 609 
2 Anatomical and Physiological Considerations of Gastro-intestinal 
Tract (GIT) for Delivery 610 
3 Introduction to Polymeric Nanoparticles as Carriers 614 
4 Preparation of Polymeric Nanoparticles 615 
5 Design Consideration for Nanoparticle-based Delivery Systems . . 619 
5.1 Polymer characteristics 619 
5.2 Drug characteristics 620 
5.3 Application characteristics 621 
6 Nanoparticles in Experimental and Clinical Medicine 621 
6.1 Drug delivery in the oral cavity 621 
6.2 Gastric mucosa as a target for oral nanoparticle-mediated 
therapy 625 
6.3 Nanoparticles for delivery of drugs and vaccines in the small 
intestine 626 
6.4 Nanoparticles for colon-specific delivery 632 
7 Integrating Polymeric Nanoparticles and Dosage Forms 634 
8 Toxicology and Regulatory Aspects 636 
8.1 Safety 637 
8.2 Quality of material/characterization 638 
8.3 Environmental considerations 638 
9 Conclusion and Outlook 639 
References 640 
Nanoparticular Carriers for Ocular Drug Delivery 649 
Alejandro Sanchez and Maria J. Alonso 
1 Biopharmaceutical Barriers in Ocular Drug Delivery. Classification 
of Nanoparticulate Carriers for Ocular Drug Delivery 650 
2 Nanoparticulate Polymer Compositions as Topical Ocular Drug 
Delivery Systems 651 
2.1 First generation: Polymer nanoparticles and nanocapsules 
for topical ocular drug delivery 652 
2.1.1 Acrylic polymers-based nanoparticles 654 
2.1.2 Polyester-based nanoparticles and nanocapsules . . .655 
2.1.3 Polysaccharide-based nanoparticles 657 
2.2 Second nanoparticles generation: The coating approach . . . . 659 
2.2.1 Polyacrylic coating 659 
2.2.2 Polysaccharide coating 660 
2.2.3 Polyethyleneglycol (PEG) coating 662 
Contents xxix 
2.3 Third nanoparticles generation: Towards functionalized 
nanocarriers 663 
3 Nanoparticulate Polymer Compositions as Subconjuctival Drug 
Delivery Systems 665 
4 Nanoparticulate Polymer Compositions as Intravitreal Drug 
Delivery Systems 665 
5 Conclusions and Outlook 667 
References 668 
Nanoparticles and Microparticles as Vaccine Adjuvants 675 
Janet R. Wendorf, Manmohan Singh and Derek T. O'Hagan 
1 Introduction 675 
2 Nanoparticle and Microparticle Preparation Methods 678 
2.1 Nanoparticles and microparticles made from polyesters . . . . 678 
2.2 Nanoparticles and microparticles made with chitosan 681 
2.3 Other nanoparticles and microparticles 681 
3 Adjuvant Effect of Nanoparticles and Microparticles 681 
3.1 Nanoparticles and microparticles as mucosal adjuvants . . . . 682 
3.2 Nanoparticles and microparticles as systemic adjuvants . . . . 686 
4 Delivery of DNA Using Nanoparticles and Microparticles 688 
5 Conclusions 690 
References 691 
Pharmaceutical Nanocarriers in Treatment and Imaging of Infection 697 
Raymond M. Schijfelers, Gert Storm and Irma A. J. M. Bakker-Woudenberg 
1 Introduction 697 
2 Carriers that are Easily Recognized as Foreign Materials 698 
3 Carriers that Avoid Recognition as Foreign Materials 701 
4 Local Application of Carriers 705 
5 Concluding Remarks 706 
References 707 
Introduction. Nanocarriers for Drug 
Delivery: Needs and Requirements 
Vladimir Torchilin 
Fast developing nanotechnology, among other areas, is expected to have a dramatic 
impact on medicine. The application of nanotechnology for treatment, diagnosis, 
monitoring, and control of biological systems has recently been determined by the 
NIH as nanomedicine. Among the approaches for exploiting nanotechnology developments 
in medicine, various nanoparticulates offer some unique advantages as 
pharmaceutical delivery systems and image enhancement agents.1,2 Several varieties 
of nanoparticles are available3: different polymeric and metal nanoparticles, 
liposomes, micelles, quantum dots, dendrimers, microcapsules, cells, cell ghosts, 
lipoproteins, and many different nanoassemblies. All of these nanoparticles can 
play a major role in diagnosis and therapy. This book is attempting to present the 
broad overview of different nanoparticulate drug delivery systems with all their 
advantages and limitations, as well as potential areas of their clinical applications. 
The paradigm of using nanoparticulate pharmaceutical carriers to enhance the 
in vivo efficiency of many drugs, anti-cancer drugs, first of all, well established 
itself over the past decade both in pharmaceutical research and clinical setting, and 
does not need any additional proofs. Numerous nanoparticle-based drug delivery 
and drug targeting systems are currently developed or under development.4,5 
Their use aims to minimize drug degradation upon administration, prevent undesirable 
side effects, and increase drug bioavailability and the fraction of the drug 
accumulated in the pathological area. Pharmaceutical drug carriers, especially the 
2 Torchilin 
ones for parenteral administration, are expected to be easy and reasonably cheap 
to prepare, biodegradable, have small particle size, possess high loading capacity, 
demonstrate prolonged circulation, and, ideally, specifically or non-specifically 
accumulate in required pathological sites in the body.6 
High molecular weight (40 kDa or higher), long-circulating macromolecules, 
including proteins and peptides, conjugated with water-soluble polymers, are capable 
of spontaneous accumulations in various pathological sites such as solid tumors, 
infarcts, and inflammations via the enhanced permeability and retention effect 
(EPR).7'8 This effect is based on the fact that pathological (tumor, infarct) vasculature, 
unlike vasculature of healthy tissues, is "leaky", i.e. penetrable for macromolecules 
and nanoparticulates which allows for macromolecules to accumulate 
in the pathological tissue (such as interstitial tumor space). In the case of tumors, 
such accumulation is also facilitated by the fact that lymphatic system, responsible 
for the drainage of macromolecules from normal tissues, is virtually not working 
in case of many tumors as the result of the disease.8 It has been found that the 
effective pore size of most peripheral human tumors range from 200 nm to 600 nm 
in diameter, with a mean of about 400 nm. The EPR effect allows for passive targeting 
to tumors and other pathological sites based on the cut-off size of the leaky 
Among particulate drug carriers, liposomes, micelles and polymeric nanoparticles 
are the most extensively studied and possess the most suitable characteristics 
for encapsulation of many drugs and diagnostic (imaging) agents. Many other systems 
meeting certain more specific requirements (and reviewed in this book) are 
also suggested and currently under development. Making these nanocarriers multifunctional 
and stimuli-responsive can dramatically enhance the efficiency of various 
drugs carried by these carriers. These functionalities are expected to provide: 
(a) prolonged circulation in the blood10'11 and the ability to accumulate in various 
pathological areas (such as solid tumors) via the EPR effect (protective polymeric 
coating with PEG is used for this purpose)12,13; (b) ability to specifically recognize 
and bind target tissues or cells via the surface-attached specific ligand (monoclonal 
antibodies as well as their Fab fragments and some other molecules are used for this 
purpose)14; (c) ability to respond local stimuli characteristic of the pathological site 
by, for example, releasing an entrapped drug or specifically acting on cellular membranes 
under the abnormal pH or temperature in disease sites (this property could 
be provided by surface-attached pH- or temperature-sensitive coatings); (d) ability 
to penetrate inside cells bypassing the lysosomal degradation for efficient targeting 
of intracellular drug targets (for this purpose, the surface of nanocarriers may be 
decorated by cell-penetrating peptides). Those are just the most evident examples. 
Some other specific properties can also be listed, such as an attachment of diagnostic 
moieties. Even the use of individual functionalities is already associated with highly 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 3 
positive clinical outcome  the success of Doxil, doxorubicin in long-circulating 
PEG-coated liposome, represents a good example.15 
In addition, there are numerous engineered constructs, assemblies, architectures, 
and particulate systems, whose unifying feature is the nanometer scale size 
range (from a few to 250 nm). Together with already listed systems, these include 
cyclodextrins, niosomes, emulsion particles, solid lipid particles, drug nanocrystals, 
metal and ceramic nanoparticles, protein cage architectures, viral-derived capsid 
nanoparticles, polyplexes, cochleates, and microbubbles.4,5,16-19 Therapeutic and 
diagnostic agents can be encapsulated, covalently attached, or adsorbed on to such 
nanocarriers. These approaches can easily overcome drug solubility issues, particularly 
with the view that large proportions of new drug candidates emerging from 
high-throughput drug screening initiatives are water insoluble. Yet, some carriers 
have a low capacity to incorporate active compounds (e.g. dendrimers, whose size 
is in the order of 5-10 nm). There are alternative nanoscale approaches for solubilization 
of water insoluble drugs too.20-23 One approach is to mill the substance 
and then stabilize smaller particles with a coating; this forms nanocrystals in size 
ranges suitable for oral delivery, as well as for intravenous injection.24,25 Pharmacokinetic 
profiles of injectable nanocrystals may vary from rapidly soluble to slowly 
dissolving in the blood. 
In general, the development of drug nanocarriers for poorly soluble pharmaceuticals 
represents a special task and still faces some unresolved issues. The therapeutic 
application of hydrophobic, poorly water-soluble agents is associated with 
some serious problems, since low water-solubility results in poor absorption and 
low bioavailability.26 In addition, drug aggregation upon intravenous administration 
of poorly soluble drugs might lead to such complications as embolism27 and 
local toxicity.28 On the other hand, the hydrophobicity and low solubility in water 
appear to be intrinsic properties of many drugs,29 since it helps a drug molecule to 
penetrate a cell membrane and reach important intracellular targets.30,31 To overcome 
the poor solubility of certain drugs, clinically acceptable organic solvents are 
used in their formulations,28 as well as liposomes32 and cyclodextrins.16 Another 
alternative is associated with the use of various micelle-forming surfactants in formulations 
of insoluble drugs. 
By virtue of their small size and by functionalizing their surface with synthetic 
polymers and appropriate ligands, nanoparticulate carriers can be targeted 
to specific cells and locations within the body after intravenous and subcutaneous 
routes of injection. Such approaches may enhance detection sensitivity in medical 
imaging, improve therapeutic effectiveness, and decrease side effects. Some of the 
carriers can be engineered in such a way that they can be activated by changes in the 
environmental pH, chemical stimuli, by the application of a rapidly oscillating magnetic 
field, or by application of an external heat source.19,33-35 Such modifications 
4 Torchilin 
offer control over particle integrity, drug delivery rates, and the location of drug 
release, for example, within specific organelles. Some are being designed with the 
focus on multifunctionality; these carriers target cell receptors and delivers drugs 
and biological sensors simultaneously. Some include the incorporation of one or 
more nanosystems within other carriers, as in the micellar encapsulation of quantum 
dots; this delineates their inherent nonspecific adsorption and aggregation in 
biological environments.36 
The use of nanoparticulate drug carriers seems to be especially important 
for developing efficient anticancer therapies. Although significant advances have 
occurred in our understanding of tumor origin, growth, metastasis, and many different 
types of pharmacological agents have been developed over the years to treat 
tumors, the problem of optimum delivery remains a formidable challenge. For any 
of the drug therapy strategies to be effective, the agent must be able to reach the 
tumor mass in sufficient concentration, traverse through the tumor microcirculation, 
diffuse into the interstitium, and remain at the site for the duration to induce 
tumoricidal effect. As was already mentioned, due to the porosity of the tumor 
vasculature and the lack of lymphatic drainage, blood-borne macromolecules and 
nanoparticles are preferentially distributed in the tumor via the EPR effect. However, 
nanoparticles can also be actively targeted to tumors by modifying their 
surface with certain cell-specific ligands for receptor-mediated uptake. The use 
of specific "vector" molecules can further enhance tumor targeting of nanocarries 
or make them the EPR-effect independent. The latter is especially important 
for the cases of tumors with immature vasculature, such as tumors in the early 
stages of their development, and for delocalized tumors. Vector molecules (those 
having affinity toward ligands characteristic for target tissues), capable of recognizing 
tumors were found among antibodies, peptides, lectins, saccharides, hormones, 
transferrin and some low molecular weight compounds (riboflavin, folate). 
From this list, antibodies and their fragments provide the most universal opportunity 
to target various for targeting and have the highest potential specificity. 
Vector molecules can be used for the targeting of nanoreservoir delivery systems 
as well. PEG-modified long-circulating doxorubicin-containing immunoliposomes 
targeted with anti-HER-2/neu monoclonal antibody fragments represent a recent 
example of increased efficiency of targeted delivery systems.37 In all studied HER2- 
overexpressing models, immunoliposomes showed potent anticancer activity superior 
to that of control non-targeted liposomes. In part, this superior activity was 
attributed to the ability of the immunoliposomes to deliver their load inside the 
target cells via the receptor-mediated endocytosis, which is obviously important if 
the drug's site of action sites locates inside the cell. 
An important problem is associated with the clearance of drug carriers from the 
circulation. Nanoparticular pharmaceutical carriers administered into the systemic 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 5 
circulation will be essentially removed within an hour of administration by the 
macrophages of the reticulo-endothelial system. To prolong the circulation of 
nanoparticles by evading the macrophages, their surface is modified with watersoluble 
polymers. Poly(ethylene glycol) (PEG) is very popular for surface modification 
of nanoparticulate drug delivery systems, since it has a long history of 
safe usage in biological and pharmaceutical products. Surface-bound PEG chains 
extend into the aqueous physiological environment and repel proteins, decrease 
antibody formation, and increase the circulation of the formulation in the plasma 
for extended period of time by the steric repulsion mechanism.38 
With rapid advances in molecular biology and genetic engineering, there is 
an unprecedented opportunity for delivery of drugs and genes to intracellular 
targets.39 In the case of cancer, for instance, the effectiveness of many anticancer 
drugs is limited due to its inability to reach the target site in sufficient concentrations 
and to exert the pharmacological effect. Current gene delivery systems are 
classified as being either viral or non-viral in origin. Viruses are efficient in delivery 
of genes; however, they suffer from poor safety profile. Non-viral gene delivery systems, 
albeit not as efficient as viruses, have promise of safety and reproducibility in 
manufacturing. To enhance delivery of drugs to intracellular targets and gene transfection 
efficiency using non-viral delivery systems, it is necessary to identify ways 
of overcoming the cellular barriers, for example, by using various cell-penetrating 
proteins and peptides.40,41 
Self-assembled nanosystems (nanoassemblies) for targeting subcellular 
organelles, such as the mitochondria, are also developed.42 It has become increasingly 
evident that mitochondrial dysfunction contributes to a variety of human 
disorders. Moreover, since the middle of 1990s, mitochondria, the "power houses" 
of the cell, have also become accepted as the cell's "arsenals", which reflects their 
increasingly acknowledged key role during apoptosis. Based on these recent developments 
in mitochondrial research, increased pharmacological and pharmaceutical 
efforts have led to the emergence of "Mitochondrial Medicine" as a whole new field 
of biomedical research. 
Nanoparticulate drug delivery systems are very important for the delivery of 
peptide and protein drugs and may represent a valid alternative to soluble polymeric 
carriers used earlier. The use of this type of carriers allows achieving much 
higher active moiety/carrier material ratio compared with "direct" molecular conjugates. 
They also provide better protection of protein and peptide drugs against 
enzymatic degradation and other destructive factors upon parenteral administration, 
because the carrier wall completely isolates drug molecules from the environment. 
All nanoparticulate carriers have the size, which excludes a possibility of renal 
filtration. Among particulate drug carriers, liposomes are the most extensively studied 
and poses the most suitable characteristics for protein (peptide) encapsulation. 
6 Torch i I in 
Similar to macromolecules, protein and peptide drug-bearing liposomes are capable 
of accumulating in tumors of various origins via the EPR effect.6-8 In some 
cases, however, the liposome size is too large to provide an efficient accumulation 
via the EPR effect presumably due to relatively small tumor vasculature cut 
off size.43,44 In such cases, alternative delivery systems with smaller sizes, such as 
micelles (prepared, for example, from PEG-phospholipid conjugates) can be used. 
These particles lack the internal aqueous space and are smaller than liposomes. 
Protein or peptide pharmaceutical agent can be covalently attached to the surface 
of these particles or incorporated into them via chemically attached hydrophobic 
group ("anchor"). 
In conclusion, even a brief listing of some key problems of nanocarrier-mediated 
drug delivery shows how broad and intense this area is. In addition to this, 
nanoscale-based delivery strategies are beginning to make a significant impact on 
global pharmaceutical planning and marketing. The leading experts in the area of 
nanparticulate-mediated drug delivery attempted to address these and many other 
topics in this book. We strongly believe that every reader will find the book useful 
and stimulating. 
1. West JL and Halas NJ (2000) Applications of nanotechnology to biotechnology commentary. 
Curr Opin Biotechnol 11:215. 
2. La Van DA, Lynn DM and Langer R (2002) Moving smaller in drug discovery and 
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3. Sahoo SK and Labhasetwar V (2003) Nanotech approaches to drug delivery and imaging. 
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4. Miiller, RH (1991) Colloidal Carriers for Controlled Drug Delivery and Targeting. 
Wissenschaftliche Verlagsgesellschaft: Stuttgart, Germany and CRC Press: Boca Raton. 
5. Cohen S and Bernstein H (eds.) (1996). Microparticulate Systems for the Delivery of Proteins 
and Vaccines. Marcel Dekker, New York. 
6. Gref R, Minamitake Y, Peracchia MT, Trubetskoy V, Torchilin VP and Langer R (1994) 
Biodegradable long-circulating polymeric nanospheres. Science 263:1600. 
7. Maeda H (2001) SMANCS and polymer-conjugated macromolecular drugs: Advantages 
in cancer chemotherapy. Adv Drug Deliv Rev 46:169. 
8. Maeda H, Sawa T and Konno T (2001) Mechanism of tumor-targeted delivery of macromolecular 
drugs, including the EPR effect in solid tumor and clinical overview of the 
prototype polymeric drug SMANCS. / Control Rel 74:47. 
9. Yuan F, Dellian M, Fukumura M, Leunig M, Berk BD, Torchilin VP and Jain RK (1995) 
Vascular permeability in a human tumor xenograft, Molecular size dependence and 
cutoff size. Cancer Res 55:3752. 
10. Lasic DD and Martin F (eds.) (1995) Stealth Liposomes. CRC Press: Boca Raton. 
Introduction. Nanocarriers for Drug Delivery: Needs and Requirements 7 
11. Torchilin VP and Trubetskoy VS (1995) Which polymers can make nanoparticulate drug 
carriers long-circulating? Adv Drug Del Rev 16:141. 
12. Lukyanov AN, Hartner WC and Torchilin VP (2004) Increased accumulation of PEGPE 
micelles in the area of experimental myocardial infarction in rabbits. / Control Rel 8, 
13. Maeda H, Wu J, Sawa T, Matsumura Y and Hori K (2001) Tumor vascular permeability 
and the EPR effect in macromolecular therapeutics: A review. / Control Rel 65:271. 
14. Torchilin VP (1998) Polymer-coated long-circulating microparticular pharmaceuticals. 
] Microencapsulation 15:1. 
15. O'Shaughnessy JA (2003) Pegylated liposomal doxorubicin in the treatment of breast 
cancer. Clin Breast Cancer 4,318. 
16. Thompson D and Chaubal MV (2000) Cyclodextrins (CDS)  excipients by definition, 
drug delivery systems by function (Part I: Injectable applications). Drug Del Technol 2:34. 
17. Zhang L and Eisenberg A (1995) Multiple morphologies of "crew-cut" aggregates of 
polystyrene-b-poly(acrylic acid) block copolymers. Science 268:1728. 
18. Gref R, Domb A, Quellec P, Blunk T, Muller RH, Verbavatz JM and Langer R (1995) 
The controlled intravenous delivery of drugs using PEG-coated sterically stabilized 
nanospheres. Adv Drug Del Rev 16:215. 
19. Cammas S, Suzuki K, Sone C, Sakurai Y, Kataoka K and Okano T (1997) Thermorespensive 
polymer nanoparticles with a core-shell micelle structure as site specific drug 
carriers. / Control Rel 48:157. 
20. Kabanov AV, Batrakova EV and Alakhov VY (2002) Pluronic block copolymers as novel 
polymer therapeutics for drug and gene delivery. / Control Rel 82:189. 
21. Kwon GS (2003) Polymeric micelles for delivery of poorly water-soluble compounds. 
Crit Rev Ther Drug Can Syst 20:357. 
22. Jones M and Leroux J (1999) Polymeric micelles  a new generation of colloidal drug 
carriers. Eur J Pharm Biopharm 48:101. 
23. Torchilin VP (2001) Structure and design of polymeric surfactant-based drug delivery 
systems. / Control Rel 73:137. 
24. Muller RH and Keck CM (2004) Challenges and solutions for the delivery of biotech 
drugs  a review of drug nanocrystal technology and lipid nanoparticles. / Biotechnol 
25. Kraft WK, Steiger B, Beussink D, Quiring JN, Fitzgerald N, Greenberg HE and 
Waldman SA (2004) The pharmacokinetics of nebulized nanocrystal budesonide suspension 
in healthy volunteers. / Clin Pharmacol 44:67. 
26. Lipinski CA, Lombardo F, Dominy BW and Feeney PJ (2001) Experimental and computational 
approaches to estimate solubility and permeability in drug discovery and 
development settings. Adv Drug Del Rev 46:3. 
27. Fernandez AM, Van Derpoorten K, Dasnois L, Lebtahi K, Dubois V, Lobl TJ, Gangwar S, 
Oliyai C, Lewis ER, Shochat D and Trouet A (2001) N-Succinyl-(beta-alanyl-L-leucyl- 
L-alanyl-L-leucyl) doxorubicin: An extracellularly tumor-activated prodrug devoid of 
intravenous acute toxicity. / Med Chem 44:3750. 
8 Torchilin 
28. Yalkowsky SH (ed.) (1981) Techniques of Solubilization of Drugs. Marcel Dekker: New York 
and Basel. 
29. Shabner BA and Collings JM (eds.) (1990) Cancer Chemotherapy: Principles and Practice. 
J. B. Lippincott Co: Philadelphia. 
30. Yokogawa K, Nakashima E, Ishizaki J, Maeda H, Nagano T and Ichimura F (1990) Relationships 
in the structure-tissue distribution of basic drugs in the rabbit. Pharm Res 
31. Hageluken A, Grunbaum L, Nurnberg B, Harhammer R, Schunack W and Seifert R 
(1994) Lipophilic beta-adrenoceptor antagonists and local anesthetics are effective direct 
activators of G-proteins. Biochem Pharmacol 47:1789. 
32. Lasic DD and Papahadjopoulos (eds.) (1998) Medical Applications of Liposomes. Elsevier: 
New York. 
33. Le Garrec D, Taillefer J, VanLier JE, Lenaerts V and Leroux JC (2002) Optimizing 
pH-responsive polymeric micelles for drug delivery in a cancer photodynamic therapy 
model. / Drug Targ 10:429. 
34. Meyer O, Papahadjopoulos D and Leroux JC (1998) Copolymers of N-isopropylacrylamide 
can trigger pH sensitivity to stable liposomes. FEBS Lett 41:61. 
35. Chung JE, Yokoyama M, Yamato M, Aoyagi T, Sakurai Y and Okano T (1999) Thermoresponsive 
drug delivery from polymeric micelles constructed using block copolymers 
of poly(N-isopropylacrylamide) and poly(butylmethacrylate). / Control Rel 62:115. 
36. Stroh M, Zimmer JP, Duda DG, Levchenko TS, Cohen KS, Brown EB, Scadden DT, 
Torchilin VP, Bawendi MG, Fukumura D and Jain RK (2005) Quantum dots spectrally 
distinguish multiple species within the tumor milieu in vivo. Nat Med 11:678. 
37. Park JW, Kirpotin DB, Hong K, Shalaby R, Shao Y, Nielsen UB, Marks JD, 
Papahadjopoulos D and Benz CC (2001) Tumor targeting using anti-her2 immunoliposomes. 
/ Control Rel 74:95. 
38. Veronese FM and Harris JM (2002) Introduction and overview of peptide and protein 
pegylation. Adv Drug Del Rev 54:453. 
39. Torchilin VP and Lukyanov AN (2003) Peptide and protein drug delivery to and into 
tumors: Challenges and solutions. Drug Discov Today 8:259. 
40. Schwarze SR, Ho A, Vocero-Akbani A and Dowdy SF (1999) In vivo protein transduction: 
Delivery of a biologically active protein into the mouse. Science 285:1569. 
41. Gupta B, Levchenko TS and Torchilin VP (2005) VP: Intracellular delivery of large 
molecules and small particles by cell-penetrating proteins and peptides. Adv Drug Del 
Rev 57:637. 
42. Weissig V (2003) Mitochondrial-targeted drug and DNA delivery. Crit Rev Ther Drug 
CarrSyst 20:1. 
43. Weissig V, Whiteman KR and Torchilin VP (1998) Accumulation of protein-loaded longcirculating 
micelles and liposomes in subcutaneous Lewis lung carcinoma in mice. Pharm 
Res 15:1552. 
44. Hobbs SK, Monsky WL, Yuan F, Roberts WG, Griffith L, Torchilin VP and Jain RK (1998) 
Regulation of transport pathways in tumor vessels: Role of tumor type and microenvironment. 
Proc Natl Acad Sci USA 95:4607. 
Nanoparticle Flow: Implications 
for Drug Delivery 
Alexander T. Florence 
1. Introduction 
While the experimental study of nanoparticle flow in vivo proves to be difficult, a 
variety of theoretical and practical techniques are becoming available to allow some 
understanding of the phenomena involved. These processes include (a) convective 
flow induced by the flow of blood, lymph or interstitial fluid, (b) the influence of the 
interaction of nanoparticles with themselves or with biological components and the 
effect of this on their transport, and (c) the effect of fluid flow and hence shear forces 
on particle access to, interaction with and removal from receptors. Diffusion and 
movement of particle suspensions in complex media such as interstitial tissue must 
also be considered. Much of the theoretical work which is relevant to this exploration 
of nanoparticle flow has not been directed towards biological endpoints, but 
this body of knowledge, and the analogous literature on the dynamic behavior of 
bacteria, erythrocytes and platelets provides the basis of a more rigorous analysis 
of the factors involved in drug carrier nanoparticle flow. 
As discussed in this book, nanoparticles are of value in drug, vaccine and gene 
delivery because their small dimension compared with microparticles allows them 
to interact more effectively with cells, be safely injected, and amongst other characteristics, 
diffuse further into tissues, and into and through individual cells. The 
flow of nanoparticles in capillaries, lymphatics, tumor vessels, their extravasation1 
and movement in the cytoplasm of cells are all aspects of the topic covered in 
10 Florence 
this overview, albeit from a phenomenological viewpoint. It is clear that particle 
diameter is a key parameter in the characterization and behavior of nanoparticle 
suspensions. In several of the analyses here, it becomes clear that another advantage 
of nanoparticles may be the relative lack of effect of shear stress, once particles 
adhere to surfaces as a prelude to uptake; this is contrasted with targeted microspheres 
whose residence on receptors and surfaces is size dependent, the larger 
particles being more vulnerable to detachment by shear forces. 
This chapter considers questions relating to the flow of nanoparticles in vivo, 
but which has often been simulated in vitro by chemical engineers and physicists 
interested in particle behavior in flow conditions. Spherical particles are the norm, 
but not all nanosystems are spherical. The influence of asymmetry on the transport 
of nanoparticles in vivo is largely unknown, although the rheological characteristics 
of asymmetric particle suspensions have been understood for a long time. Flow 
behavior of nanoparticles in complex networks of narrow capillaries has relevance 
in the design and operation of microfluidic devices, as well as in drug delivery 
and targeting, and in toxicology2; the extent to which it is relevant for delivery and 
targeting is explored here. Figure 1 illustrates diagrammatically some of the areas 
of interest. 
In physical terms, the following situations could be considered: (i) particle flow 
in rapidly flowing blood, including segregation and deposition of particles, and the 
behavior of particles at bifurcations in the capillary supply; (ii) the effect of shear 
on adhesion of particles of different size and shape; (iii) particle flow in more static 
conditions, for example, in the tumor interstitium or the lymphatics; (iv) flow of 
particles in tissues, including the flow of particles into narrow pores and (v) flow or 
diffusion within the anisotropic interior of cells. Added complications arise where 
bioadhesive or ligand-decorated particles are involved. In the latter case, binding 
and flow are linked. 
The enquiry can be divided into a discussion of the flow and movement of 
nanoparticles as a function of their size and route of administration, the influence 
of convection and blood or lymph flow, the influence of flow dynamics on the 
interaction of nanoparticles with target tissues and receptors, and the movement 
of nanoparticles once they have been absorbed and are making their way through 
individual cells and tissues towards a target. 
The routes of administration where flow is important potentially include all, 
even the oral route where particle flow and dynamics in the gut lumen and in the 
vicinity of both villi and microvilli is important. Particles below a critical size are 
taken up by the M-cells of Peyer's patches and by normal enterocytes, albeit in 
small quantities and find themselves in the lymph vessels, lymph nodes, blood, 
liver and spleen.3'4 If the flow of nanoparticles away from their site of absorption 
is restricted due to the flow in lymph or blood being slow, this will reduce 
Nanoparticle Flow: Implications for Drug Delivery 11 
* II 
Fig. 1. A diagrammatic representation (not to scale) of some of the areas where flow and 
transport of nanoparticles is key. I: flow in the GI tract after oral administration; II: access to and 
adhesion to M-cells of Peyer 's patches or to enterocy tes; III: passage into the mesenteric lymph; 
IV: flow in the lymph vessels and entrapment in the lymph nodes (not shown); V: transport 
between lymph and blood. A: blood flow; B: adhesion to capillary walls; C: extravasation and 
flow in tissue; D: flow and deposition at vessel bifurcations; and E: movement into tumours. 
Each route (the subcutaneous route is also indicated) will involve a complex sequence of 
nanoparticle pathways, most involving lymph, blood and intestinal fluid. 
bioavailability and distribution. Rapid flow provides superior sink conditions and 
hence the use of everted gut sacs and cell monolayers in vitro can give unrealistic 
results for nanoparticle transit. The influence of flow dynamics on extravasation and 
perhaps on the enhanced permeability and retention (EPR) effect for nanocarriers 
has perhaps not been fully addressed. Both involve consideration of particle size, the 
diffusion and flow of nanoparticles through narrow channels, as well as navigation 
of tortuous environments. The availability of "extreme" nanoparticles in the form 
of dendrimers5 and quantum dots6 makes this topic a vital one in understanding 
the fate, toxicity7 or accumulation of what is metrically a wide range of systems. 
2. Background 
Our own interest in this field has resulted in part from studies on the size dependency 
of nanoparticle uptake after oral administration, where mesenteric lymphatic 
transport of 500 nm nanoparticles post absorption is determined by the flow of particles 
in a single file in the smallest mesenteric lymph vessels (Fig. 2). In addition, 
12 Florence 
Fig. 2. 500 nm polystyrene latex particles flowing in the mesenteric lymph vessels mostly 
in single-file mode, from Jani et al. 
studies on the flow of liposomes and niosomes, and the extrusion of flexible vesicles 
from glass capillaries under pressure, converting polyhedral vesicles into multilayer 
tubules, led us to consider the influence of stress forces on carrier integrity. 
Clearly, the elasticity of vesicles is important in their negotiation of capillaries when 
their diameter exceeds that of the capillary. The flow of particles in fabricated capillaries 
which have a radius close to the particle radius is a challenge that has been 
tackled theoretically.8'1* We have suggested that multi-bilayer tubules (Fig. 3) can 
act as models for such flow experiments.11 
Rheological examination of nanoparticle-blood mixtures and nanoparticle suspensions 
of mixed radius has also illustrated the potential complexity of particle 
Fig. 3. A flexible non-ionic surfactant based multi-bilayer tube, around 1 /xm in diameter, 
extruded from a suspension of polyhedral niosomes, which might be adapted for use as a 
model for the study of capillary nanoparticle flow.11 
Nanoparticle Flow: Implications for Drug Delivery 13 
movement in blood (unpublished data). In addition, erythrocyte blockage at bifurcations, 
or narrowing of vessels can lead to slowing down of blood flow and a 
change in rheology as the haematocrit increases. More recently, investigation of 
the transport of nanoparticles across cell monolayers12 and intracellular transport 
of dendrimers13 has assisted in defining some of the issues involved in targeting 
within cells. 
3. Studies on Nanoparticle Flow 
The work of Fokin and colleagues14 on the transport of viral-sized colloids, following 
intravenous or intra-lymphatic injection, is relevant to drug delivery even if 
their objectives were different. 100-200 nm diameter sulphur colloid particles reach 
the lymph after IV injection in around 25 minutes; after intra-lymphatic injection 
particles appear in the venous blood only after 4 seconds. Following subcutaneous 
injection, similar particles14-16 reach the lymph after 2-9 min, although 95% of the 
particles remain at the injection site for at least 45 minutes. Here, the nanoparticles 
are being used as indicators of blood and lymph flow. What is also relevant to drug 
delivery is the influence of fluid flow on the movement and fate of nanoparticles. 
Ilium et al.17 observed uptake rates of 1.27 jxm and 15.8 /zm polystyrene particles 
in the lung and liver after IV injection. The sequestration in the lung was size 
dependent, but possibly affected because the smaller particles were taken up by 
the Kupffer cells of the liver, leaving the larger particles free to be taken up by the 
lung tissue. The rapidity of this suggests that, in effect, flow of the microparticles 
is solely determined by blood flow. 
4. Convection and Diffusion 
Blood flow drives the convective flow of suspended particles. Diffusional transport 
occurs in static conditions or conditions of low fluid velocity. In a tube of flowing 
liquid, convective dynamics propel the particles in the direction of flow, but at the 
walls of the tube, there is the possibility of particle diffusion resulting in deposition. 
Blood velocity (mm s_1) in arterioles and venules is a function of vessel diameter, 
as shown in Fig. 4. In venules, the maximum velocity according to Jain18 is approximately 
12 mm s_1, while in arterioles, it can reach about 30 mm s^1. 
Fluid velocity in tubes is not constant throughout the diameter of the tube as 
Fig. 5 indicates, a feature that is important when the interaction of nanoparticles 
with epithelial cells or capillary walls is considered. 
The radial variation of shear is a factor that must be considered in polydisperse 
nanoparticulate systems and where nanoparticles adhere to erythrocytes, causing 
two distinct size distributions. If nanoparticles adhere to erythrocytes20 or other 
14 Florence 
o 0 o 
I" I 
0 25 
- 30 
" 20 
5 g |5i d^^OD0 a 
) 25 50 
Vessel Diameter (nm) 
Fig. 4. Maximum blood velocity (mms 1) in arterioles and venules as a function of vessel 
diameter, redrawn from R. K. Jain.18 
Fig. 5. Diagram showing the velocity pattern in a tube of flowing liquid. Particles of different 
size separate according to their diameter. The large particles, being unable to approach 
close to the capillary wall, experience the faster fluid streamlines toward the centre; hence, 
they move more rapidly, as described by Silebi and DosRamos.18,19 This is the basis of the 
field flow fractionation. 
blood elements, the translocation of the particles is controlled by the particular 
element to which it adheres. The rheology of suspensions of mixed particles is 
complex: viscosity reduces first with an increase in the fraction of larger particles 
in a suspension, and as the volume fraction increases, so does the viscosity.21 
Ding et al.22 formulated a theoretical model examining particle migration 
in nanoparticle suspensions flowing through a pipe. "The model considers particle 
migration due to spatial gradients in viscosity and shear rate as well as 
Brownian motion. Particle migration due to these effects can result in significant 
non-uniformity in particle concentration over the cross section of the pipe" in 
particular for larger particles. Three mechanisms were proposed by the authors 
for migration in such non-uniform shear flow: (i) shear induced migration where 
Nanoparticle Flow: Implications for Drug Delivery 15 
particles move from regions of higher shear rate to regions of lower shear rate; 
(ii) viscosity gradient induced migration  particles move from regions of higher 
viscosity to regions of lower viscosity and (iii) self-diffusion due to Brownian 
Diffusion inside microtubules has been studied to understand taxol binding to 
tubulin structures.23 The dimension of the tubulin lumen is of the order of 17nm, 
approaching macromolecular dimensions, leading to friction between the inner 
walls and the moving macromolecules. This "hindrance" will also be an issue in 
the movement of nanoparticles in the smallest capillaries. With dendrimers whose 
diameters may be as small as 6 nm, the application of hindered theory to their movement 
could be relevant. No vessels are of this small radius, but the key parameter 
is the ratio of particle diameter to capillary diameter. The approach may well be 
important in cellular networks. It is not only capillary vessels that are the conduits 
of particle movement, but after extravasation, there is passage through cellular networks. 
The process could be considered to be akin to diffusion in porous networks. 
Binding of the moving particle (or macromolecule) to the luminal surface of the vessel 
will also hinder free flow or movement, a positive event in the case of specific 
ligand targeting of "decorated" systems. 
Polydisperse nanosystems can segregate during flow or migrate differentially 
leading to concentration differences.22 Particle-image velocimetry (PIV)24 has been 
used to track the flow characteristics of microparticles. The effects of flow on adhesion 
of monocytes to endothelial cells25 is relevant to the influence of flow and shear 
in particle interactions and uptake. 
The significance of flow can be demonstrated by the use of pharmacological 
agents which change normal vessel patency, so that by the concomitant use of 
noradrenalin26 or angiotensin26,27 which constrict only normal vessels, the ratio of 
tumor to normal tissue blood flow can be optimized. 
5. Bifurcations 
Many theoretical studies of nanoparticle flow deal with linear tubes, whereas 
in vivo movement occurs through complex vessel architectures with bifurcations28,29 
(Fig. 6). Behavior at bifurcations in a vascular or capillary system is dependent not 
only on particle diameter, but also on the rigidity or flexibility of the particle concerned. 
Colloid transport in a bifurcating structure has been the subject of one recent 
paper.30 It is a process which depends on the orientation of the bifurcations, especially 
with particles whose density is greater than that of the medium, as well as on 
the different flow rates in the individual branches which are likely to be of different 
If nanoparticles are trapped or associate at bifurcations or indeed other obstacles 
in capillaries, then it is likely that they might associate more permanently, thereby 
16 Florence 
A & - 
Q.30IAnin HB M ,.., 2...s. M ae'0.3 
Fig. 6. Three-dimensional distributions of nanoparticles in a bifurcation airway model of 
Zhang et ah, Aerosol Sci., 2005, 36, 211-233. DEF is the deposition enhancement factor, the 
representations shown here being for a steady inhalation. While these data are for air-flow, 
not dissimilar patterns of deposition might be estimated to occur in liquid flows. Deposition 
in these models occurs primarily by Brownian diffusion; deposition efficiencies increase with 
decreasing nanoparticle size and lower inlet Reynolds numbers. 
changing their intrinsic rheological behavior. Flexible particles do not of course 
suffer the same constraints in movement and progress, but their flexibility can lead 
to slow negotiation of movement around obstacles (Fig. 7). 
6. Interaction with Blood Constituents and 
Endogenous Molecules 
Nanoparticles may interact with blood constituents31: the adsorption of albumin, 
IgG and fibrinogen from blood onto hydrophobic particles is well known, but the 
Nanoparticle Flow: Implications for Drug Delivery 17 
Fig. 7. Two captured pictures from a video of a large vesicle moving in a flowing stream 
of smaller vesicles. The stills show a flexible vesicle approaching an obstacle, and rolling 
around the obstacle while adhering to it, a process encouraged by its elasticity. 
effect of nanoparticles on blood has been less well studied. Kim's31 data indicate 
that the interaction of nanoparticles with erythrocytes changes the dynamics of 
flow of both erythrocytes and particles. Chambers and Mitragotri20 found that 
nanoparticles as large as 450 nm adhered to erythrocytes, and thus remained in 
the circulation for several weeks. The percentage of latex nanospheres in the circulation 
over a period of 6 hrs was highly dependent on particle size, retention 
times decreasing with increasing diameter from 220 nm to 1100 nm. These data are 
difficult to interpret on the basis of flow, as the erythrocytes with attached nanoparticles 
are eliminated somewhat faster than the native erythrocytes. Gorodetsky and 
colleagues32 explored interactions of carboplatin (CPt) nanoparticles (formed by 
CPt interaction with fibrinogen) with the fibrin mesh caused by the induction of 
clot formation. 
18 Florence 
7. Nanoparticles with Surface Ligands 
There appear to be no rheological studies comparing surface protein decorated 
nanoparticles with the unadorned forms. Certainly, it is possible that aggregation 
may be caused by the change in surface properties and that this will in turn change 
flow patterns and perhaps masking of ligands33 as posited in Fig. 8. Nanoparticles 
are of course sensitive to the medium in which they are placed34 even in vitro when 
cell media can cause significant increases in diameter because of particle flocculation. 
We have suggested that the interaction with surface receptors of nanoparticles 
decorated with ligands is more complex than intimated in discussions of targeting 
generally.33 Figure 8 represents some of the factors: the aggregation of particles, 
the masking of ligands by this process, the detachment of ligands and the shearinduced 
removal of attached particles as discussed above. The instability of plant 
lectins, frequently used as surface proteins on nanosystems, is discussed by Gabor 
et al.35 The processes illustrated in Fig. 8 might explain some of the lack of complete 
success of targeted drug delivery. 
8. Deposition on Surfaces and Attachment to Receptors 
in Flow Conditions 
Nanoparticles in vivo flow in blood, lymph or tissue fluid at greater or lesser 
velocities, as discussed above. Deposition of particles which might occur in a 
static situation is itself a complex process, and will depend on the rugosity of the 
Aggregation and loss of 
ligand accessibility 
Repulsion Blocking by 
cleaved ligands 
n B n 
Fig. 8. Diagram illustrating variations from the ideal of a single ligand-decorated nanoparticle 
interacting with receptors spaced at an appropriate distance from the particles. The diagram 
shows the loss of ligand accessibility which would follow from the aggregation of the 
particles before interaction with the desired surface, repulsion between a particle attached 
to the receptor surface, and an approaching particle and blockage of the receptors due to 
interaction of cleaved ligands with the receptors. 
Nanoparticle Flow: Implications for Drug Delivery 19 
receiving surface.36 Particle deposition from flowing suspensions has been the subject 
of research37 which has considered not only diffusion, convection, geometrical 
interception and migration under gravity, but also the influence of tangential 
Patil et al.39 examined the rate of attachment of 5, 10, 15 and 20 /zm particles 
with a reconstituted P-selectin glycoprotein ligand-1 construct 19.ek.Fc. The rate of 
attachment was not affected by particle diameter. However, the shear stress required 
to set the adherent particles in motion (Sc) decreased with increasing particle diameter, 
and the rolling velocity of the 19.ek.Fc microspheres increased with increasing 
diameter. From their data, if we extrapolate the critical shear (a plot of 1 / S c is linear 
with diameter over the range 5-20 jxm), it suggests that particles below one micron 
in diameter will not be removed by shear forces. 
Usually we consider the flow of many particles in collective diffusion. The diffusion 
coefficient of a single particle and the collective diffusion coefficient coincides 
at infinite dilution, but can differ at higher concentrations.40 
Cell adhesion mediated by not one but two receptors has been considered by 
Bhatia et al.41; the analysis would also apply to decorated nanoparticles. In their 
study, the two receptors were selectin and integrin ICAM; "the state diagram" 
evolved shows the area of firm adhesion as opposed to rolling adhesion for leukocytes 
as a function of receptor densities and association rate constants. The fate of 
transport initial adhesion attachment uptake 
Adhesion time 
short range interactions 
' or specific ligand e 
receptor interactions 
Fig. 9. Processes occurring in the deposition of nanoparticles in flow conditions as a function 
of the range of interaction forces (nm) and adhesion times. At the start, mass transport 
to the surface occurs, initial adhesion following through electrostatic attraction and van der 
Waals' forces. Hydrophobic interactions can play their part as well as specific receptorligand 
interactions which are short-range interactions. Drawn after Vacheethasanee and 
20 Florence 
nanoparticles in flowing blood, their adhesion, extravasation and permeation into 
tumors, thus depends on a complex of factors such as diameter, surface ligand density 
and orientation, shape, capillary diameter and rugosity, bifurcations, viscosity 
and flow gradients. 
9. Does Shape Matter? 
Nanosystems can be prepared in a variety of shapes. Nanocrystals42 are often irregular; 
there are asymmetric carbon nanotubes, and surfactant and lipid vesicles can 
be produced as discs, polyhedral structures,40,43 toroids and tubes.21,44 The vesicle 
constructs often have dimensions larger than 500 nm; it must be assumed that 
vesicles in the nanometer size range will be less affected. In these systems, shape is 
less important than membrane properties in controlling the release of encapsulated 
drug, but the flow properties of vesicular suspensions are clearly determined by 
shape and elasticity As most particulate delivery vectors have been spherical, little 
attention has been paid to the influence of shape on fate; yet it is known that the 
shape of environmental particles and fibres, for example, influences their fate and 
As discussed above, there are two different but related effects of particle flow: 
the effect of particle shape and size and characteristics on flow, as well as the effect of 
flow on flexible particles, as discussed by Bruinsma.46 With elastic vesicles, we have 
argued44 that shape matters because it affects flow and potential fate in vivo through 
extravasation for instance; elasticity also allows vesicles to be transported in vessels 
which would be blocked by solid particles. The elasticity and visco-elasticity of such 
systems may be important in differentiating them from solid nanoparticles. 
Much of the debate on whether the shape of vesicles matters, is dependent on 
the knowledge of the nature of the capillary blood supply and the forces exerted on, 
and the damage done to, vesicles as they move in capillaries.44 In studies conducted 
in our laboratories with doxorubicin loaded niosomes, 60% of the drug remained in 
the vesicles 8 hrs after intravenous administration.47 The extent to which the drug 
loss was due to diffusion or to damage is not known, but vesicles subjected to deliberate 
stress can lose considerable amounts of their payload, simply by extrusion of 
the vesicles through capillaries of reducing diameter.48 Reduction in diameter of 
systems below 1 micron will clearly reduce such stresses and allow flexible systems 
to retain their loads intact. 
Vasanthi et al.49 treated the anisotropic diffusion of oblate spheroids, explaining 
that because non-spherical molecules rotate as they translate, their motion differs 
significantly from that of a sphere. For rods, theory predicts that the diffusion 
coefficient in the direction parallel to the major axis of the rod (Dn) is twice that in 
the perpendicular direction (Di.). 
Nanoparticle Flow: Implications for Drug Delivery 21 
B IJ.  I T' '    -l I 
3*M x/2 *!* 0 
Platelet angle a 
Fig. 10. The non-dimensional bond force as a function of the angle of an ellipsoidal platelet 
passing through zero when the platelet is 90 to the surface. From Mody et a/.50 
There are few studies which have considered the motion of ellipsoidal particles 
near a plane wall, although this is relevant to platelet flow and adhesion to the walls 
of vessels. Mody and colleagues50 have addressed the issue, observing the effects 
of shear stress on platelet adhesion. Platelets, unlike leukocytes, do not roll but 
display a flipping motion in the direction of flow, due to their flattened ellipsoidal 
structure. The bond force between the ellipse and the surface is dependent on the 
platelet angle as defined in Fig. 10. 
Flexible systems such as vesicles have been widely studied, while being forced 
under pressure in capillaries smaller than the vesicle diameter. The elasticity of 
the membranes can be estimated from the extent of deformation. Vesicle flow in 
linearly forced motion has been followed. Flexible vesicles adjust their shape to 
equilibrate the applied force51; locally in some cases, two-dimensional flow of lipids 
in the vesicle membrane occurs,52 clearly influencing the position of the embedded 
surface ligands. 
There are many nanoparticulates which are produced in non-spherical forms, 
hence the transport properties of asymmetric particles is important.53 
10. Speculations on Flow and the EPR Effect 
Erythrocyte velocity in normal vessels depends on vessel diameter (see Fig. 4 
above), but there is no such dependence in tumors (Fig. 11), even though flow 
may be an order of magnitude slower. According to Jain,18,52 "to reach cancer cells 
in a tumor, a blood-borne therapeutic molecule, particle or cell must make its way 
22 Florence 
t 0.5 
i 0.2 
o.i H 
MCalV J U7 
i r*I1 ri 1 I i ""> r 1 1 T" 
0 10 20 30 40 SO 60 70 0 10 20 30 40 50 60 70 
Tumor Vessel Diameter Qua) 
Fig. 11. Diagram from Jain18 showing the lack of a clear relationship between erythrocyte 
velocity and tumor vessel diameter in two tumor types, MCalV and U87. The low and 
variable velocities compared to those shown in Fig. 4 are evident. 
into blood vessels of the tumor and cross the vessel wall into the interstitium and 
finally migrate through the interstitium". While blood flow is reduced in tumor 
vessels, nonetheless cancer cells have been reported to compress tumor vessels and 
this will have consequences on fluid flow.54 This is highly relevant to the enhanced 
permeation and retention effect (EPR) which allows entry of macromolecules into 
tumors from spaces in the ill-formed tumor vasculature.55 Access of nanoparticles 
to tumors is equally important and must be critically size-dependent. 
In convective flow, stable colloidal particles may be captured by the process of 
hydrodynamic bridging,52,56 events which may be relevant to the first process in 
the enhanced permeation and retention (EPR) effect. At high velocities but in the 
low Re regime, hydrodynamic forces acting on the particles at an entrance to a pore 
(or a defect in a tumor vessel) may overcome colloidal repulsive forces and result 
in flocculation of the particles and the plugging of the pore. The effects of velocity, 
particle concentration, and the ratio of pore size to particle size (the aspect ratio) on 
retention by hydrodynamic bridging have been studied. The effect of velocity on 
retention by bridging is opposite to that of retention by deposition. There is a critical 
flow velocity necessary for particle bridging to occur, a function of the net colloidal 
interparticle and particleporous medium repulsion that must be overcome by the 
hydrodynamic forces for bridging to occur. Figure 12 demonstrates the effect for an 
aspect ratio of 3.7 (220 nm particles) 
11. Intra-tumoral Injection 
Direct injection of delivery systems into tumors has both been a mode of experimental 
and clinical drug delivery. Solutions allow the drugs to diffuse or leach out 
Nanoparticle Flow: Implications for Drug Delivery 23 
f l 
v^ o- 
Fig. 12. Particle behavior prior to entry to a pore of radius, rp: (a) a discrete nanoparticle, 
(b) aggregate, (c) individual particles converging on the pore opening demonstrating 
hydrodynamic bridging, as discussed by Ramachandran.56 We speculate that events such 
as bridging might occur during entry of nanoparticles into tumors through fenestrations in 
the tumor capillary blood supply, aspects of the enhanced permeation and retention effect. 
of the tumor, especially through the needle track, whereas suspensions might allow 
some greater residence time. Viral vectors have been administered by intra-tumoral 
injection.57 To decrease the extent of viral dissemination into the systemic circulation, 
a viscous alginate solution was used as the viral vehicle. However, transgene 
expression was not increased perhaps because, as the authors speculate, the diffusion 
of the virus is reduced by the viscous medium once in situ. The transport 
of particles of viral dimensions requires, according to Higuchi et al.,16 convective 
rather than diffusional transport. "The early transport of colloids into the vascular 
and lymphatic vessels relies largely on an extracellular pathway which depends 
on convective transport (i.e. solvent drag)". "Thus the particle uptake in the period 
immediately after injection is relatively insensitive to particle size; it is expected 
that viruses will be carried in the tissue towards lymphatics and microvessels with 
great efficacy leading to enhanced escape compared with the relatively low levels" 
for 1 and 0.4 /xm particles.16 The question of how resistance to convective transport 
in the interstitial space (the interstitial fluid plus the extracellular matrix) has been 
considered at least for molecules.58 Clearly, the spacing between the cells or between 
fibres will be a significant factor in determining the size cut-off for transport. 
12. Conclusions 
This phenomenological survey of possible factors affecting the flow and hence the 
mass transport of nanoparticles has explored a range of scenarios. It is by no means 
a comprehensive survey, but there is sufficient in the literature to stimulate further 
analyses to provide a better overall prediction of the influence of particle characteristics, 
particularly, diameter and surface nature, shape and flexibility on delivery 
and targeting to remote sites in the body. Conf ocal microscopy and other techniques 
24 Florence 
will allow experimental study of nanoparticles so that their movement and fate can 
be studied in a variety of tissues. Atomic force microscopy allows measurement of 
forces of interaction of particles with cells and receptors to aid a more quantitative 
approach. However, it is wrong to underestimate the challenges ahead if nanoparticulate 
carriers are to be designed to overcome the various biological barriers and 
survive transit in the conduits of capillary blood or lymph, extravasation and tissue, 
and subsequently intracellular transport.59 One cannot help but conclude that as 
many properties including flow are dictated by particle diameter, one of the most 
important strategies is to ensure the maintenance of particle stability in vivo. 
1. El-Sayed M, Kiani MF, Naimark MD, Hikal AH and Ghandehari H (2001) Extravasation 
of poly(amidoannine) (PAMAM) dendrimers across microvascular network endothelium. 
Pharm Res 18:23-28. 
2. Health and Safety Executive, Health Effects of particles produced for nanotechnologies. 
Sudbury, UK, pp. 1-37. 
3. Hussain N, Jaitley V and Florence AT (2001) Recent advances in the understanding of 
uptake of microparticulates across the gastrointestinal lymphatics. Adv Drug Del Rev 
4. Florence AT (1997) The oral absorption of micro- and nanoparticulates: Neither exceptional 
nor unusual. Pharm Res 14:259-266. 
5. Florence AT and Hussain N (2001) Transcytosis of nanoparticle and dendrimer delivery 
systems: Evolving vistas. Adv Drug Del Rev 50 (Suppl 1):S69-S89. 
6. Fortina P, Kricka LJ, Surrey S and Grodzinski P (2005) Nanobiotechnology: The 
promise and reality of new approaches to molecular recognition. Trends Biotechnol 23: 
7. Warheit DB, Laurence BR, Reed KL, Roach DH, Reynolds GA and Webb TR (2004) 
Comparative pulmonary toxicity assessment of single-wall carbon nanotubes in rats. 
Toxicol Sci 77:117-125. 
8. Sugihara-Seki M and Skalak R (1997) Asymmetric flows of spherical particles in a 
cylindrical tube. Biorheology 34:155-159. 
9. Wang H and Skalak R (1969) Viscous flow in a cylindrical tube containing a line of 
spherical particles. / Fluid Mech 38:75-96. 
10. Jani P, Halbert GW, Langridge J and Florence AT (1989) The uptake and translocation of 
latex nanospheres and microspheres after oral-administration to rats.}Pharm Pharmacol 
11. Nasseri B and Florence AT (2003) Microtubules formed by capillary extrusion and 
fusion of surfactant vesicles. Int} Pharm 266:91-98. 
12. Rowland RES, Taylor PW and Florence AT (2005) / Drug Del Sci Tech 
13. Ruenraroengsak P, Hartell N and Florence AT (2005) unpublished. 
14. Fokin AA, Robicsek F and Masters TN (2000) Transport of viral-size particulate matter 
after intravenous versus intralymphatic entry. Microcirculation 7:357-365. 
Nanoparticle Flow: Implications for Drug Delivery 25 
15. Fokin AA, Robicsek F, Masters TN, Schmid-Schonbein GW and Jenkins SH (2000) 
Propagation of viral-size particles in lymph and blood after subcutaneous inoculation. 
Microcirculation 7:193-200. 
16. Higuchi M, Fokin A, Masters TN, Robicsek F and Schmid-Schonbein GW (1999) Transport 
of colloidal particles in lymphatics and vasculature after subcutaneous injection. 
JAppl Physiol 86:1381-1387. 
17. Ilium L, Davis SS, Wilson CG, Thomas NW, Frier M and Hardy JG (1982) Blood clearance 
and organ deposition of intravenously administered colloidal particles. The effects of 
particle size, nature and shape. Int ] Pharm 12:135-146. 
18. Jain RK (2001) Delivery of molecular medicine to solid tumors: Lessons from in vivo 
imaging of gene expression and function. / Control Rel 74:7-25. 
19. Silebi CA and DosRamos JG (1989) Separation of submicrometer particles by capillary 
hydrodynamic fractionation (CHDF). / Coll Interf Sci 130:14-24. 
20. Chambers E and Mitragotri S (2004) Prolonged circulation of large polymeric nanoparticles 
by non-covalent adsorption on erythrocytes. / Control Rel 100:111-119. 
21. Nunez ADR, Pinto R and Paredes VME (2002) Viscosity minimum in bimodal concentrated 
suspensions under shear. Eur Phys } E 9:327-334. 
22. Ding Y and Wen D (2005) Particle migration in a flow of nanoparticle suspensions. 
Powder Technol 149:84-92. 
23. Odde D (1998) Diffusion inside microtubules. Eur Biophys J 27:514-520. 
24. Sinton D (2004) Microscale flow visualization. Microfluid Nanofluid 1:2-21. 
25. Chiu J-J, Chen C-N, Lee P-L, Yang CT, Chuang HS and Chien SUS (2003) Analysis 
of the effect of disturbed flow in monocytic adhesion to endothelial cells. / Biomech 
26. Shankar A, Loizidou M, Burnstock G and Taylor I (1999) Noradrenaline improves the 
tumour to normal blood flow ratio and drug delivery in a model of liver metastases. 
Br } Surgery 86:453^57. 
27. Goldberg JA, Murray T, Kerr DJ, Willmott N, Bessent RG, McKillop JH and McCardle 
CS (1991) The use of angiotensin II as a potential method of targeting cytotoxic microspheres 
in patients with intrahepatic tumours. Br J Cancer 63:308-310. 
28. Zhang Z, Kleinstreuer C, Donohue JF and Kim CS (2005) Comparison of micro- and 
nano-size particle depositions in a human upper airway model. Aerosol Sci 36:211-233. 
29. Shi HKC, Zhang Z and Kim CS (2004) Nanoparticle transport and deposition in bifurcating 
tubes with different inlet conditions. Phys Fluids 16:2199-2213. 
30. James SC and Chrysikopoulos CV (2004) Dense colloid transport in a bifurcating fracture. 
/ Coll Interf Sci 270:250-254. 
31. Kim D, El-Shall H, Dennis D and Morey T (2005) Interaction of PLGA nanoparticles 
with human blood constituents. Coll SurfB 40:83-91. 
32. Gorodetsky R, Peylan-Ramu N, Reshef A, Gaberman E, Levdansky L and Marx G 
(2005) Interactions of carboplatin with fibrin(ogen), implications for local slow release 
chemotherapy. / Control Rel 102:235-245. 
33. Florence AT (2005) Issues in oral nanoparticle drug carrier uptake and targeting. / Drug 
Targ 12:65-70. 
26 Florence 
34. Singh B, Hussain N, Sakthivel T and Florence AT (2003) Effect of physiological media on 
the stability of surface-adsorbed DNA-dendron-gold nanoparticles. / Pharm Pharmacol 
35. Gabor F, Bogner E, Weissenboeck A and Wirth M (2004) The lectin-cell interaction and its 
implications to intestinal lectin-mediated drug delivery. Adv Drug Del Rev 56:459-480. 
36. Adamczyk Z, Siwek B, Jaszczolt K and Weronski P (2004) Deposition of latex particles 
at heterogeneous surfaces. Colloids Surface A: Physicochem Eng Aspects 249:95-98. 
37. Adamczyk Z (1989) Particle transfer and deposition from flowing colloid suspensions. 
Coll Surf 35:283-308. 
38. Vacheethasanee K and Marchant RE (2000) Non-specific staphylococcus epidermidis 
adhesion: Contribtuions of biomaterial hydrophobicity and charge, in An, YH, 
Friedman RJ (eds.) Handbook of Bacterial Adhesion: Principles, Methods and Applications. 
Humana Press, Totowa, NJ, pp. 73-90. 
39. Patil VRS, Campbell CJ, Yun YH, Slack SM and Goettz DJ (2001) Particle diameter 
influences adhesion under flow. Biophys J 80:1733-1743. 
40. Bowen WR and Mongruel A (1998) Calculation of the collective diffusion coefficient of 
electrostatically stabilised colloidal particles. Coll Surface A 138:161-172. 
41. Bhatia SK, King MR and Hammer DA (2003) The state diagram for cell adhesion mediated 
by two receptors. Biophys J 84:2671-2690. 
42. Akerman ME, Chan WC, Laakkonen P, Bhatia SN and Ruoslahti E (2002) Nanocrystal 
targeting in vivo. Proc Natl Acad Sci USA 99:12617-12621. 
43. Uchegbu IF, Schatzlein A, Vanlerberghe GMN and Florence AT (1997) Polyhedral nonionic 
surfactant vesicles. J Pharm Pharmacol 49:606-610. 
44. Florence AT, Nasseri B and Arunothyanun P (2004) Does shape matter? Spherical, 
polyhedral and tubular vesicles, in Sonke S (ed.) Carrier-based Drug Delivery. American 
Chemical Society, Washington, pp. 75-84. 
45. Schins RP (2002) Mechanisms of genotoxicity of particles and fibers. Inhal Toxicol 14: 
46. Bruinsma R (2005) Rheology and shape transitions of vesicles under capillary flow. 
Physica A 234:249-270. 
47. Uchegbu IF, Double JA, Turton JA and Florence AT (1995) Distibution, metabolism and 
tumoricidal activity of doxorubicin administered in sorbitan monostearate (Span 60) 
niosomes in the mouse. Pharm Res 12:1019-1024. 
48. Nasseri B and Florence AT (2003) Some properties of extruded non-ionic surfactant 
micro-tubes. Int f Pharm 254:11-16. 
49. Vasanthi R and Bhattacharyya S (2005) Anisotropic diffusion of spheroids in liquids: 
Slow orientational relaxation of the oblates. / Chem Phys 116:1092-1096. 
50. Mody NA, Lomakin O, Doggett TADTG and King MR (2005) Mechanics of transient 
platelet adhesion to von Willebrand factor under flow. Biophys J 88:1432-1443. 
51. Kern N and Fourcade B (1999) Vesicles in linearly forced motion. Europhys Lett 
52. Nasseri B and Florence AT (2005) The relative flow of the walls of phospholipid tethers. 
Int J Pharm 298:372-377. 
Nanoparticle Flow: Implications for Drug Delivery 27 
53. Naess SN and Elgsaeter A (2005) Transport properties of non-spherical nanoparticles 
studied by Brownian dynamics: Theory and numerical simulations. Energy 30:831-844. 
54. Padera TP, Stoll BR, Tooredman JB, Capen D, di Tomaso E and Jain RK (2004) Cancer 
cells compress intratumour vessels. Nature 427:695. 
55. Maeda H (2001) The enhanced permeability and retention (EPR) effect in tumor vasculature: 
the key role of tumor-selective macromolecular drug targeting. Adv Enzyme 
Regul 41:189-207. 
56. Ramachandran VV, Venkatesan R, Tryggvason G and Scott FH (2000) Low Reynolds 
Number Interactions between Colloidal Particles near the Entrance to a Cylindrical 
Pore. / Coll Interf Sci 229:311-322. 
57. Wang Y, Hu JK, Krol A, Li YP, Li CY and Yuan F (2003) Systemic dissemination of viral 
vectors during intratumoral injection. Mol Cancer Ther 2:1233-1242. 
58. McGuire S and Yuan F (2001) Quantitative analysis of intratumoral infusion of color 
molecules. Am J Physiol Heart Circ Physiol 281:H715-H721. 
59. Jones AT, Gumbleton M and Duncan R (2003) Understanding endocytic pathways and 
intracellular trafficking; a prerequisite for effective design of advanced drug delivery 
systems. Adv Drug Del Rev 55:1353-1357. 
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Polymeric Nanoparticles as Drug 
Carriers and Controlled Release 
Implant Devices 
SM Moghimi, E Vega, ML Garcia, 
OAR Al-Hanbali and KJ Rutt 
1. Introduction 
Polymeric nanoparticles are submicron size entities, often ranging from 10-1000 nm 
in diameter, and are assembled from a wide variety of biodegradable (e.g. albumin, 
chitosan, alginate) and non-biodegradable polymers (Tables 1 and 2). The most 
active area of research using polymeric nanoparticles is in controlled delivery of 
pharmaceuticals following parenteral, oral, pulmonary, nasal, and topical routes 
of administration.1-6 Indeed, therapeutic agents can be encapsulated, covalently 
attached, or adsorbed onto such nanocarriers. These approaches can easily overcome 
drug solubility issues; this is particularly important as a significant proportion 
of new drug candidates arising from high-throughput screening initiatives are 
water insoluble. Polymeric nanoparticles, however, differ from nanosuspensions 
of drugs which are sub-micron colloidal dispersions of pure particles of drug that 
are stabilized by surfactants.7 By virtue of their small size and by functionalizing 
their surface with polymers and appropriate ligands, polymeric nanoparticles can 
also be targeted to specific cells and locations in the body.1,3'5'8-10 Thus, polymeric 
nanoparticles may overcome stability issues for certain drugs and minimize druginduced 
side effects. The extent of drug encapsulation/incorporation, as well as 
30 Moghimi etal. 
the release profile from polymeric nanocarriers, however, depends on the polymer 
type and its physicochemical properties, the particle size and its morphology (e.g. 
solid nanospheres as opposed to polymeric nanocapsules).4 In addition, depending 
on the polymer characteristics, polymeric nanocarriers can also be engineered 
in such a way that they can be activated by changes in the environmental pH, 
chemical stimuli, or temperature.1112 Such modifications offer control over particle 
integrity, drug delivery rates, and the location of drug release, for example, 
within specific organelles. For instance, nanoparticles made from poly(lactide-coglycolide), 
PLGA, can escape the endo-lysosomal compartment within minutes 
of internalization in intact cells and reach the cytosol.12 This is due to the selective 
reversal of the surface charge of nanoparticles from the anionic to the cationic state in 
endo-lysosomes, resulting in a local particle-membrane interaction with subsequent 
cytoplasmic release. This is an excellent approach for channelling antigens into the 
highly polymorphic MHC class-I molecules of macrophages and dendritic cells 
for subsequent presentation to CD8+ T lymphocytes. Other applications include 
cytoplasmic release of plasmid vectors and therapeutic agents (e.g. for combating 
cytoplasmic infections and for slow cytoplasmic release of drugs that act on nuclear 
Polymeric nanoparticles are also beginning to make a significant impact on 
global pharmaceutical planning (life-cycle management) and market intelligence. 
For example, due to imminent expiration of patents, pharmaceutical companies 
may launch follow-up or nano-formulated versions of a product to minimize 
generic threats to best-selling medicines. This could lead to an extension of as much 
as 20 years from a new patent on the nanoparticulate formulation of the drug. 
By coalescing certain polymeric nanoparticles carefully from an aqueous 
suspension, shape retentive hydrogels can be formed to erode partially or 
completely.1113 Drugs and macromolecules may be trapped within interstitial 
spaces between particles during aggregate formation. Thus, hydrogel nanoparticles 
have potential as controlled release implant devices following local administration 
or implantation, and may also serve as tissue engineering scaffolds with concurrent 
morphogenic protein release. 
This article will briefly review some of the most commonly used laboratory 
scale methods for the production of polymeric nanoparticles and drug encapsulation 
procedures. The importance of the nanometre scale size range and surface 
engineering strategies for site-specific targeting of polymeric nanoparticles, following 
different routes of administration, are also discussed. 
2. Nanoparticle Engineering 
Polymeric nanoparticles are usually prepared either directly from preformed 
polymers such as aliphatic polyesters (Table 1) and block copolymers (Table 2), 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 31 
Table 1 Chemical properties of some commonly used aliphatic polyesters in nanoparticle 
Polymer Type Melting Point (C) Glass Transition Resorption Time 
Temperature (C) (Months) 
DL-PLA Amorphous 50-60 12-16 
PGA 220-230 35^0 6-12 
DL-PLGA (50/50) Amorphous 45-50 1-2 
DL-PLGA (75/25) Amorphous 45-50 4-5 
PCL 55-65 (-65)-(-60) >24 
DL-PLA: poly(L-lactide); PGA: poly(glycolide); DL-PLGA: poly(DL-lactide-co-glycolide); PCL: poly- 
Table 2 Selected examples of block copolymers for production of biodegradable 
or by polymerization of monomers.4 Commonly used methodologies include 
the solvent evaporation,14-15 the spontaneous emulsification/ solvent diffusion,16 
nanoprecipitation or solvent displacement17'18 and emulsion polymerization 
techniques.19-21 The method of choice depends on the polymer and the drug 
type, as well as the required particle size distribution and polydispersity 
indices. However, some polymers, such as comb-like polyesters, the di-block 
copolymer poly(ethylene oxide-b-sebacic acid) and tri-block copolymer poly(2- 
methyloxazoline)-fr-poly(dimethylsiloxane)-fr-poly(2-methyloxazoline) can spontaneously 
form stable nanoparticles (core-shell type nanospheres).22-24 
In the solvent evaporation method, the polymer is simply dissolved together 
with the drug in an organic solvent and the mixture is then emulsified to form either 
an oil-in-water nanoemulsion (for encapsulation of hydrophobic drugs) or waterin-
oil nanoemulsion (for encapsulation of hydrophilic drugs) using suitable surfactants. 
Nanoparticles are then obtained following evaporation of the solvent and 
can be concentrated by filtration, centrifugation or lyophilization. The spontaneous 
emulsification/solvent diffusion method is a modified version of the solvent evaporation 
technique, which utilizes a water-soluble solvent (e.g. methanol or acetone) 
along with a water-insoluble one such as chloroform. As a result of the spontaneous 
32 Moghimi etal. 
diffusion of the water-soluble solvent into the water-insoluble phase, an interfacial 
turbulence is created leading to the formation of nanoparticles. Nanoprecipitation, 
however, is a versatile and simple method. This is based on spontaneous formation 
of nanoparticles during phase separation (the Marangoni effect), which is induced 
by slow addition of the diffusing phase (polymer-drug solution) to the dispersing 
phase (a non-solvent of the polymers, which is miscible with the solvent that solubilizes 
the polymer). The dispersing phase may contain surfactants. Depending 
on the solvent choice and solvent/non-solvent volume ratio, this method is suitable 
for encapsulation of both water-soluble and hydrophobic drugs, as well as 
protein-based pharmaceuticals.17'18 
In emulsion polymerization, the monomer is dispersed into an aqueous phase 
using an emulsifying agent. The initiator radicals are generated in the aqueous 
phase and they diffuse into the monomer-swollen micelles. Anionic polymerization 
in the micelles is then initiated by the hydroxyl ions of water. Chain transfer 
agents are abundant and termination occurs by radical combination. The size and 
molecular masses of nanoparticles are dependent on the initial pH of the polymerization 
medium.20 Drugs are incorporated during the polymerization step or can 
be adsorbed into the nanosphere surface afterwards. The addition of cyclodextrins 
to the polymerization medium can promote the encapsulation of poorly watersoluble 
drugs.25 Depending on the monomer used, some drugs can also initiate the 
polymerization step, resulting in the covalent attachment of drug molecules to the 
nanospheres. For instance, photosensitizers such as naphthalocyanines, can initiate 
the polymerization of alkylcyanoacrylates.26 
A number of specialized approaches (e.g. dialysis, salting-out, supercritical 
fluid technology, denaturation, ionic interaction, ionic gelation, and interfacial 
polymerization) have also been described for the preparation of polymeric 
nanoparticles, based on the choice of the starting material and the biological 
2.1. Drug release mechanisms 
The release profile of drugs from nanoparticles depends on the physicochemical 
nature of the drug molecules as well as the matrix.4'16'28,33-36 Factors include mode of 
drug attachment and/or encapsulation (e.g. surface adsorption, dispersion homogeneity 
of drug molecules in the polymer matrix, covalent conjugation), the physical 
state of the drug within the matrix (such as crystal form), and parameters controlling 
matrix hydration and/or degradation. Generally, rapid release occurs by desorption, 
where the drug is weakly bound to the nanosphere surface. If the drug is 
uniformly distributed in the polymer matrix, the release occurs either by diffusion 
(if the encapsulated drug is in crystalline form, the drug is first dissolved locally 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 33 
and then diffuses out) or erosion of the matrix, or a combination of both mechanisms. 
Erosion can be further subdivided into either homogeneous (with uniform 
degradation rates throughout the matrix) or heterogeneous (where degradation is 
confined at the surface) processes. Parameters such as polymer molecular weight 
distribution, crystallinity, hydrophobicity/hydrophilicity, melting and glass transition 
temperature, polymer blends and prior polymer treatment (e.g. oxygen-plasma 
treatment) all control the extent of matrix hydration and degradation. For instance, 
in the case of aliphatic polyesters, their degradation time is shorter for low molecular 
weight polymers, more hydrophilic polymers, more amorphous polymers and 
copolymers with high glycolide content (Table 1). 
3. Site-specific Targeting with Nanoparticles: Importance 
of Size and Surface Properties 
Numerous articles have recently discussed the importance of nanoparticle size and 
surface characteristics in controlling their biodistribution, following different routes 
of administration.1 ~3/5 Only a brief overview is provided here. 
Following intravenous injection, liver (Kupffer cells) and spleen (marginal zone 
and red pulp) macrophages clear polymeric nanoparticles rapidly from the blood 
circulation.1 Opsonization, which is surface deposition of blood opsonic factors 
such as fibronectin, immunoglobulins, C-reactive and certain complement proteins, 
often aid particle recognition by these macrophages. Indeed, the propensity 
of macrophages of the reticuloendothelial system for rapid recognition and 
clearance of particulate matter has provided a rational approach to macrophagespecific 
targeting with nanoparticles (e.g. for the treatment of obligate intracellular 
microorganisms, delivery of toxins for macrophage killing, and diagnostic agents).1 
However, the rapid sequestration of nanoparticles by macrophages in contact with 
blood is problematic for the efficient targeting of polymeric nanoparticles to nonmacrophage 
sites. Thus, inherent in nanoparticle design is the precision surface 
manipulation and engineering with synthetic polymers; this affords control over 
nanoparticle interaction and fate within biological systems. There are numerous 
examples where the surface of nanocarriers is carefully assembled with projected 
"macromolecular hairs" made from poly(ethyleneglycol), PEG, or its derivatives 
(e.g. methoxyPEG-albumin, PLA-PEG) or other related polymers [e.g. block 
copolymers such as selected poloxamers and poloxamines, poly(phosphazene)- 
poly(ethyleneoxide)].3,5 This is achieved either during the particle assembly procedures 
or polymerization step, or post particle manufacturing. This strategy 
suppresses macrophage recognition by an array of complex mechanisms, which 
collectively achieve reduced protein adsorption and surface opsonization. Therefore, 
such entities, provided that they are below 150 nm in size, exhibit prolonged 
34 Moghimi et al. 
residency time in the circulation, and are referred to as "stealth" or "macrophageevading" 
nanoparticles.1,5 The efficiency of the "macrophage-evading" process is 
dependent on polymer type and its surface stability, reactivity, and physics (e.g. 
surface density and assumed conformation).5 Prolonged circulation properties are 
ideal for slow or controlled release of therapeutic agents in the blood to treat 
vascular disorders. Long circulating polymeric nanoparticles may have application 
in vascular imaging too (e.g. detection of vascular bleeding or abnormalities). 
Long-circulating nanoparticles can also escape from vasculature and this is normally 
restricted to sites where the capillaries have open fenestration or when the 
integrity of the endothelial barrier is perturbed by inflammatory processes or by 
tumor growth.5 However, extravasated nanoparticles, as in tumour interstitium, 
distribute heterogeneously in perivascular clusters that do not move significantly; 
these particles may therefore act as depot systems, particularly for the sustained 
release of antiangiogenic agents, and to some extent, for drug delivery to multidrug 
resistant tumors (e.g. by co-encapsulation of both anticancer drugs and the competitive 
inhibitors of active drug efflux pumps).1 The surface of long-circulating 
nanoparticles is also amenable for modification with targeting ligands. Such entities 
can navigate capillaries and escape routes in search of signature molecules 
expressed by the target; this process is often referred to as "active targeting".1-5 
For example, certain cancer cells express folate receptors and these receptors have 
the ability to endocytose stealth nanoparticles that are decorated with folic acid. 
Delivery of anti-cancer agents to tumor cells by such means could overcome the 
possibility of multi-drug resistance.1,37 
Non-deformable "stealth" nanoparticles, however, are prone to splenic filtration 
at interendothelial cell slits, if their size exceeds that of the width of the cell 
slits (200-250 nm).38,39 Indeed, these "splenotropic" vehicles can deliver their cargo 
efficiently to the red-pulp regions of the sinusoidal spleen. Activated or stimulated 
macrophages are also known to rapidly phagocytose stealth nanoparticles; 
stealth nanospheres may therefore have applications as diagnostic/imaging tools 
for the identification of stimulated or newly recruited hepatic macrophages.40 Such 
diagnostic procedures may prove useful for patient selection or for monitoring 
the progress of treatment with long-circulating nanoparticles carrying anti-cancer 
agents, thus minimizing damage to hepatic macrophages.41 
Polymeric nanospheres can also target endothelial cells on the bloodbrain 
barrier. For instance, following intravenous injection polysorbate 80-coated 
poly(alkylcyanoacrylate), PACA, nanospheres attract apolipoprotein E from the 
blood, thus mimicking low density lipoprotein (LDL) and become recognizable 
by LDL receptors expressed by the blood-brain barrier endothelial cells.10 Another 
related example is PEG-coated PACA nanoparticles, with the ability to localize 
mainly in the ependymal cells of the choroid plexus and the epithelial cells of pia 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 35 
region and the ventricles of the mouse and the rat brain.42 The molecular basis of 
this deposition pattern remains to be unravelled. 
Others have administered nanoparticles directly to pathological sites for 
optimal biological performance.43 One example is intramurally delivered PLGA 
nanoparticles to an injured artery following angioplasty, using a cardiac infusion 
catheter. Here, nanoparticles penetrate the dilated arterial wall under pressure and 
once the pressure is released, the artery returns to its normal state resulting in particle 
immobilization in the arterial wall, where they may act as a sustained release 
system for drugs and genetic materials.43 Again, particle size is an important parameter; 
the smaller the size, the greater the arterial deposition and cellular entry, as 
well as lower inflammatory responses. 
Polymeric nanospheres also provide intriguing opportunities for lymphatic 
drug delivery, as well as for diagnostic imaging of the lymphatic vessels and their 
associated lymph nodes when injected interstitially.44 The extent of lymphatic delivery 
and lymph node localization of nanospheres depends on their size and surface 
characteristics. For instance, hydrophilic nanoparticles, in the size range of 
30-100 nm, as opposed to their hydrophobic counterparts, repulse each other and 
interact poorly with the ground substance of the interstitium and drain rapidly into 
the initial lymphatics through patent junctions in the lymphatic capillaries.45,46 The 
drained particles are conveyed to the nodes via the afferent lymph. Macrophages 
of medullary sinuses and paracortex are mainly responsible for particle capture 
from the lymph, but this also depends on nanoparticle surface properties. Larger 
nanospheres (>150nm), however, are retained at interstitial sites for prolonged 
periods of time and may therefore act as sustained release systems for drugs and 
antigens.47,48 For example, large-sized PLGA particles can provide antigen release 
over weeks and months following continuous or pulsatile kinetics. By mixing particle 
types with different degradation and pulsatile release kinetics, multiple discrete 
booster doses of encapsulated antigens can be provided after a single administration 
of the formulation (e.g. 1-2 and 6-12 months).48 An alternative approach is the use 
of nanoparticle hydrogels for slow and local antigen release. For example, by controlling 
the ionic strength of the dispersion medium, monodisperse nanoparticles of 
poly-2-hydroxyethylmethacrylate, poly(HEMA), and poly[HEMA-co-methacrylic 
acid] coalesce together to form a shape retentive hydrogel suitable for interstitial 
implantation.13 Macromolecules may be trapped between the particle aggregates 
and their release is controlled by a combination of diffusion (larger particles packed 
together have larger spaces in the lattice, and this allows for faster diffusion) and 
erosion (arising from aggregates that contain particles with methacrylic acid).13 
Nanoparticles that erode from the aggregate are drained into the lymphatic system 
and may be retained by the regional nodes. Similarly, by controlling the inherent 
physical attractive forces between model polystyrene nanoparticles, ordered lattices 
36 Moghimi et al. 
Fig. 1. Scanning electron micrographs of uncoated and surface-modified polystyrene 
nanoparticles. Due to surface hydrophobicity uncoated nanospheres (A), 350 nm in size, 
tend to aggregate. By controlling the physical attractive forces between the nanoparticles (by 
surface coating with an appropriate concentration of a block copolymer), ordered structures 
are formed and these can be deposited onto the surface of large microspheres (B). 
can be deposited on the surface of very large microspheres (Fig. 1). Following subcutaneous 
localization, surface adsorbed nanospheres may gradually detach from 
the parent microsphere and gain entry into the lumen of the lymphatic capillaries. 
Polymeric nanoparticles also have numerous applications following oral delivery. 
Evidence suggests that the adsorption of particulates in the intestine following 
oral administration take place at the Peyer's patches.49-50 The epithelial cell 
layer overlying the Peyer's patches contains specialized M cells. These cells can 
Polymeric Nanoparticles as Drug Carriers and Controlled Release Implant Devices 37 
sample particles from the lumen and transport them to the underlying macrophages 
and dendritic cells. Indeed, numerous studies have confirmed protective immunity 
induced by mucosal immunization with PACA, PLGA and chitosan based particulate 
systems.3,32,48'50-53 Part of the success is due to the encapsulation of antigens in 
polymeric particulate systems, which provides better protection for the antigen during 
intestinal transit. The immune outcomes have included mucosal (secretory IgA) 
and serum antibody (IgG and IgM) responses, as well as systemic cytotoxic T lymphocyte 
responses in splenocytes. Induction of an appropriate immune response 
following oral administration depends primarily on factors that affect uptake and 
particle translocation by M cells. These include particle size, dose, composition, and 
surface chemistry, as well as the region of the intestine where particles are taken up, 
membrane recycling from intracellular sources and the species.50 Tolerance to orally 
administered microparticulate encapsulated antigens is another potential outcome, 
but it has received little attention. 
The bioavailability of some drugs can be improved after oral administration 
by means of polymeric nanoparticles.54-57 This is a reflection of drug protection 
by the nanoparticle against hostile conditions of the gastrointestinal tract, as well 
as the mode of nanoparticle interaction with mucosal layers. However, the bioadhesive 
properties of nanoparticles may vary with their size and surface characteristics 
(e.g. surface charge, surface polymer density and conformation), as well as 
the location and type of the mucosal surface in the gastrointestinal tract. Similarly, 
improved drug bioavailability has also been reported following ocular administration 
with PLA, PACA, poly(butylcyanoacrylate) and Eudragit nanoparticles.6,58-61 
For example, loading of tamoxifen in PEGylated nanoparticles proved successful 
in the treatment of autoimmune uveortinitis following intraocular injection.59 
Interaction of surface-modified polymeric nanoparticles with nasal associated lymphoid 
tissue and their transport across nasal mucosa have also received attention, 
particularly with respect to peptide-based pharmaceuticals and antigen 
4. Conclusions 
Polymeric nanoparticles are promising vehicles for site-specific and controlled 
delivery of therapeutic agents, following different routes of administration and 
these trends seem to continue with advances in materials and polymer chemistry 
and pharmaceutical nanotechnology. However, nanoparticles do not behave similarly; 
their encapsulation capacity, drug release profile, biodistribution and stability 
vary with their chemical makeup, morphology and size. Inherently, nanosphere 
design and targeting strategies may vary according to physiological and therapeutic 
needs, as well as in relation to the type, developmental stage and location of 
38 Moghimietal. 
the disease. Attention should also be paid to toxicity issues that may arise from 
nanoparticle administration and the release of their polymeric contents and degradation 
products. These issues are discussed elsewhere.1,63~66 
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Genetic Vaccines: A Role for Liposomes 
Gregory Gregoriadis, Andrew Bacon, 
Brenda McCormack and Peter Laing 
1. Introduction 
Prevention of microbial infections by the use of vaccines is a preferred alternative 
to treatment. Vaccines have been applied successfully, for example, in the eradication 
of smallpox as well as against tetanus, diphtheria, whooping cough, polio and 
measles, thus preventing millions of deaths each year. However, vaccines made 
of attenuated organisms, which mimick natural infections usually without the disease, 
can be potentially unsafe. For instance, there is a risk of reversion during replication 
of live viruses or even mutation to a more pathogenic state. Furthermore, 
with immunocompromised individuals, some of the attenuated viruses may still 
provoke disease. On the other hand, with killed virus vaccines, their extracellular 
localization and subsequent phagocytosis by professional antigen presenting cells 
(APC) or antigen-specific B cells, lead to MHC-II class restricted presentation and to 
T helper cell and humoural immunity. However, they do not elicit significant cytotoxic 
T cell (CTL) responses. Moreover, subunit vaccines produced from biological 
fluids may not be entirely free of infectious agents. Even with subunit and peptide 
vaccines produced recombinantly or synthetically (and thus considered safe), 
immune responses are weak and often not of the appropriate kind. The great variety 
of immunological adjuvants1'2 that are now available go a long way in rendering 
subunit and peptide vaccines stronger and more efficient. However, more than seventy 
years after the introduction of aluminium salts as an adjuvant, only two other 
adjuvants, liposomes3 and MF59,1 have been approved for use in humans.4 Thus, 
44 Gregoriadis et al. 
inspite of considerable progress, the road to the ideal vaccine appears as elusive as 
ever, until recently. 
Recent developments have led to a novel and exciting concept, namely de novo 
production of the required vaccine antigen by the host's cells in vivo, which promises 
to revolutionize vaccination especially where vaccines are either ineffective or 
unavailable. The concept entails the direct injection of antigen-encoding plasmid 
DNA which, on uptake by cells, localizes to some extent into the nucleus where 
it transfects the cells episomally. The produced antigen is recognized as foreign 
by the host and is thus subjected to pathways similar to those observed for antigens 
of internalized viruses (but without their disadvantages), leading to protective 
humoural and cell mediated immunity.5-9 A series of publications since 1992 first 
established the ability of plasmid DNA to induce an immune (antibody) response 
to the encoded foreign protein10; in experiments with DNA encoding influenza 
nucleoprotein, immunity was both humoural and cell-mediated, and also protective 
in mice challenged with the virus.11,12 This was the first demonstration of an 
experimental DNA vaccine. Another observation was the induction of humoural 
and cell-mediated immunity against HIV-1 using plasmids encoding the HIV rev 
and env proteins.13 Similar results were obtained with a gene for the hepatitis B 
surface antigen (HBsAg).14 DNA immunization was also found to apply in cancer 
treatment. For instance, injection of plasmids encoding tumor antigens promoted 
immune responses15,16 which were protective in an animal model.6 The concept 
of DNA immunization has now been adopted by vaccinologists worldwide using 
an ever increasing number of plasmids encoding immunogens from bacterial, viral 
and parasitic pathogens, and a variety of tumors.8,9 In many of these studies, genetic 
immunization has led to the protection of animals from infection.5-9 A number of 
clinical trials for the therapy of, or prophylaxis against, a variety of infections are 
in progress.8,9 
2. The DNA Vaccine 
A plasmid DNA vaccine is usually6 supercoiled and consists of the gene encoding 
the vaccine antigen (the section of the target pathogen which elicits protective 
immunity), a promoter sequence which is often derived from cytomegalovirus 
(CMV) or Rous sarcoma virus (RSV)), an mRNA stability polyadenylation region 
at the 3' end of the insert, and the plasminogen activator gene which controls the 
secretion of the recombinant product. In addition, there are an origin of replication 
for the amplification of the plasmid in bacteria, and a gene for antibiotic resistance 
to select the transformed bacteria. 
Immunization procedures with DNA vaccines are carried out by the intramuscular 
and, to a lesser extent, the intraepidermal route. Other routes include the 
Genetic Vaccines: A Role for Liposomes 45 
oral, nasal, vaginal, intravenous, intraperitoneal and subcutaneous routes.8,9 Intramuscular 
injection of DNA vaccines leads to such types of immunity as CTL.5'9,11'12 
This was unexpected because antigen presentation requires the function of professional 
APC.17 However, myocytes which were shown5 to take up the plasmid only 
to a small extent and with only a fraction of cells participating in the uptake, are 
not professional APCs. Although myocytes carry MHC class I molecules and can 
present endogenously produced viral peptides to the CD8+ cells to induce CTLs, 
they do so inefficiently18 as they lack vital costimulatory molecules (e.g. the B7-1 
molecule). It is thus difficult to accept that antigen presentation, leading to a CTL 
response, occurs via myocytes. Instead, it was reported18 that CTL responses occur 
as a result of the transfer of antigenic material between the myocytes and professional 
APC to some extent. In parallel, it could also be that plasmid secreted by 
the myocytes or as such, is taken up directly by APC infiltrating the injected site. 
Such APC would include dendritic cells which will express and present peptides 
to CD8+ cells following transport to the lymph nodes or spleen. On the other hand, 
CD4+ cells may be activated by APCs via MHC class II presentation of antigen 
secreted by the myocytes (or released from them after their destruction via a Tc 
response) and captured by the cells. Such events would lead to both cellular (Th 1) 
and humoural (Th 2) immunity. Indeed, it has been shown6 that dendritic cells are 
the essential APC involved in immune responses elicited by intramuscularly given 
DNA vaccines. 
3. DNA Vaccination via Liposomes 
Vaccination with naked DNA by the intramuscular route is dependent on the ability 
of myocytes to take up the plasmid. However, some of the DNA may also be 
engulfed by APC infiltrating the site of injection, or in the lymph nodes following 
migration of the DNA to the lymphatics. The extent of DNA degradation by 
extracellular deoxyribonucleases is unknown, but depending on the time of its residence 
interstitially, degradation could be considerable. Therefore, approaches that 
protect DNA from the extracellular nucleases and promote DNA uptake by cells 
more efficiently, or target it to APC, should contribute to the optimal design of DNA 
It has been suggested19 that as APC are a preferred alternative to muscle cells 
for DNA vaccine uptake and expression, liposomes (known3 to be taken up avidly 
by APC infiltrating the site of injection or in the lymphatics, an event that has 
been implicated3 in their immunoadjuvant activity) would be a suitable means 
of delivery of entrapped DNA to such cells. Liposomes would also protect20 their 
DNA content from deoxyrubonuclease attack. Moreover, the structural versatility21 
of the system would ensure that its tranfection efficiency is further improved 
46 Cregoriadis ct al. 
by the judicial choice of its structural characteristics or by the co-entrapment of 
cytokine genes, other adjuvants (e.g. immunostimulatory sequences), or indeed 
protein antigens (see later) together with the plasmid vaccine. As a number of 
injectable liposome-based drug formulations, including vaccines against hepatitis A 
and influenza, have been already licensed for clinical use,21 acceptance of the system 
clinically would be less problematic than with other systems that are still at an 
experimental stage. 
3.1. Procedure for the entrapment of plasmid DNA into liposomes 
A variety8,22,23 of plasmid DNAs have been quantitatively entrapped into liposomes 
by a mild dehydration-rehydration procedure.20'22,23 The procedure (Fig. 1) 
consists of mixing preformed small unilamellar vesicles (SUV) with a solution 
of the DNA destined for entrapment, freeze-drying of the mixture, followed by 
controlled rehydration of the formed powder, and centrifugation to remove nonentrapped 
material. Formed liposomes are multilamellar.20 However, when an 
appropriate amount of sucrose is added to the SUV and DNA mixture prior 
to dehydration,24 the resulting liposomes are much smaller (about 100-160 nm 
in diameter). As expected, DNA incorporation values8'23-26 were higher (up to 
90% of the amount used) when a cationic lipid was present in the bilayers. No 
apparent relationship was observed between amount of DNA used (10-500//g) 
and the values of incorporation for the compositions and lipid mass used.8,23,26 
The possibility that DNA was not entrapped within the bilayers of cationic liposomes, 
but was rather complexed with their surface (as suggested by the high 
Fig. 1. Entrapment of DNA and/or protein into cationic liposomes. The procedure entails 
mixing up empty SUV with the solute(s) destined for entrapment and subsequent dehydration. 
On rehydration, most of the solute(s) is recovered entrapped within the generated 
multilamellar liposomes. 
Genetic Vaccines: A Role for Liposomes 47 
Naked DNA 
"Complexed** DNA 
Naked DNA 
taraXeB (DNA) g 
"Complexed" DNA 
Fig. 2. Gel electrophoresis of a mixture of cationic SUV and pRc/CMV HBS before (complexed 
DNA) and after (entrapped DNA) dehydration-rehydration of the mixture. 
"incorporation" values obtained on mixing)20 was examined by treating liposomeentrapped 
and liposome-complexed DNA with deoxyribonuclease. Substantially, 
more liposome-entrapped DNA remained intact than when it was complexed,20 
presumably because of the inability of the enzyme to reach its substrate in the former 
case. The significant resistance of complexed DNA (despite its accessibility) 
to the enzyme could be attributed to its condensed state.25 Additional evidence 
that the DNA was entrapped within liposomes was obtained by gel electrophoresis 
of a mixture of cationic SUV and plasmid DNA before (complexed DNA) and 
after dehydration-rehydration of the mixture (entrapped DNA). When the anionic 
sodium dodecylsulphate (SDS) was incorporated in the gel, complexed DNA was 
dissociated from the SUV, presumably because of ionic competition for the cationic 
charges. As expected, "entrapped" DNAretained its association with the liposomes, 
suggesting its unavailability to the competing SDS anions26 (Fig. 2). 
3.2. DNA immunization studies 
Previously,20 liposome-entrapped plasmid found to transfect cells in vitro regardless 
of the vesicle surface charge was tested in immunization experiments,19,27 using 
a plasmid (pRc/CMV HBS) encoding the S region of the hepatitis B surface antigen 
(HBsAg; subtype ayw). Mice (Balb/c) that are repeatedly injected intramuscularly 
with 5 or 10/ig plasmid entrapped in cationic liposomes, exhibited at all 
times much greater (up to 100-fold) antibody (IgGi) responses (Fig. 3) against the 
48 Gregoriadis et al. 
c 5 a. I 
8 I 
26 34 44 
Days after first injection 
Fig. 3. Immune responses in mice injected with naked, or liposome-entrapped pRc/CMV 
HBS. Balb/c mice were injected intramuscularly on days 0, 10, 20, 27 and 37 with 5 /xg of 
DNA entrapped in cationic liposomes composed of PC, DOPE and DOTAP (A), DC-Chol 
(B) or SA (C) (molar ratios 1:0.5:0.25), or in the naked form (D). Animals were bled 7, 15, 
26, 34 and 44 days after the first injection and sera tested by ELISA for IgGT (black bars), 
IgG2a (white bars) or IgG2b (grey bars) responses against the encoded hepatitis B surface 
antigen (HBsAg; S region, ayw subtype). Values are means SD of log10 of reciprocal end 
point serum dilutions required for OD to reach readings of about 0.2. Sera from untreated 
mice gave log10 values of less than 2.0. IgGj responses were mounted by all mice injected 
with liposomal DNA but became measurable only at 26 days. Differences in log10 values 
(all IgG subclasses at all time intervals) in mice immunized with liposomal DNA and mice 
immunized with naked DNA were statistically significant (P < 0.0001-0.002). (Reproduced 
with permission from Ref. 19.) 
Genetic Vaccines: A Role for Liposomes 49 
encoded antigen than animals immunized with the naked plasmid. Values of other 
subclasses (IgG2a and IgG2b) were also greater (up to 10-fold) (Fig. 3). Moreover, 
IgGj responses for the liposome-entrapped plasmid DNA were higher (up to 10- 
fold) than those obtained with DNA complexed with similar cationic liposomes.19 
This was also true for IFN-y and IL-4 levels in the spleens of immunized mice.19 
In other experiments,8 the effect of the route of injection of the pRc/CMV HBS 
plasmid was examined with respect to both humoural and cell-mediated immunity, 
using Balb/c mice and an outbred mouse strain (T.O.). Results8 comparing 
responses for liposome-entrapped and naked plasmid DNA showed greater antibody 
(IgGi) responses for the entrapped DNA, not only by the intramuscular route, 
but also the subcutaneous and the intravenous routes. As there were no significant 
differences in the titers between the two strains,8 it was concluded that immunization 
with liposomal pRc/CMV HBS is not MHC restricted. Results obtained 
on the testing of IFN-y and IL-4 in the spleens (not shown) exhibited a similar 
Involvement of muscle cells in the mechanism by which liposomes promote 
greater immune responses to the encoded antigen than seen with the naked plasmid, 
is rather unlikely. Although, cationic liposomes could in theory bind to and 
be taken up by the negatively charged myocytes, the negatively charged proteins 
present in the interstitial fluid would neutralize21 the cationic liposomal surface 
and thus interfere with such binding. In addition, vesicle size (about 600-700 nm 
average diameter; Ref. 26) would render access to the cells difficult, if not impossible. 
It is therefore more likely that cationic liposomes are endocytosed by APC, 
including dendritic cells, in the lymphatics where liposomes are expected to end 
up.28 Uptake of liposomal plasmid DNA is supported in studies where mice were 
injected intramuscularly or subcutaneously with liposomes entrapping the plasmid 
(pCMV- EFGP), encoding the enhanced fluorescent green protein or with the 
naked plasmid. Fluorescence microscopy of sections of the lymph nodes draining 
the injected site revealed (Fig. 4) much more green fluorescence when the plasmid 
was administered in the liposomal form.27 It appears8'19 that the key ingredient of 
the DNA-containing liposomes as used in Fig. 3, contributing to enhanced immune 
responses, is the cationic lipid. The mechanism by which liposomal DNAreaches the 
nucleus for episomal transfection is poorly understood. It is conceivable, however, 
that some of the endocytosed liposomal DNA escapes the endocytic vacuoles prior 
to their fusion with lysosomes (in a way similar to that proposed29 for vesicle-DNA 
complexes) to enter the cytosol for eventual episomal transfection and presentation 
of the encoded antigen. It is perhaps at this stage of intracellular trafficking of DNA, 
spanning its putative escape from endosomes and access to the nucleus, that the 
cationic lipid, possibly together within the fusogenic phosphatidylethanolamine 
(PE) component, plays a significant role. 
50 Gregoriadis et al. 
Kttmam w* ii'm<|* 
ff!f b ((lit 
Fig. 4. Fluorescence images of muscle and lymph node sections from mice injected intramuscularly 
with 10/xg liposome-entrapped or naked pCMVEGFP and killed 48h later. 
Sections from untreated animals were used as controls. (Reproduced with permission from 
Ref. 27.) 
3.3. Induction of a cytotoxic T lymphocyte (CTL) response 
by liposome-entrapped plasmid DNA 
Immunization studies with liposome-entrapped DNA vaccines were expanded30 
to include the cytotoxic T lymphocyte (CTL) component of the immune response. 
This was measured by the specific killing of syngeneic target cells pulsed with a 
recognized CTL epitope peptide derived from the antigen tested. To that end, the 
type and degree of immune response induced following subcutaneous injection of 
DNA in cationic liposomes was monitored and compared with that obtained with 
DNA alone injected by the same route. 6-8 week old, female C57/BL6 (H-2d) mice 
were injected subcutaneously with one or two doses of 2.5 or 10 ^g ovalbumin 
(OVA)-encoding plasmid DNA (pCI-OVA), either alone or entrapped in liposomes. 
Animals immunized subcutaneously with 100 /xg of OVA protein complexed with 
1 /xg of cholera toxin (CT) served as a positive control. Blood samples and spleens 
were collected from all animals one week after the last injection and tested for 
anti-OVA total IgG (serum), CTL activity and cytokine release (spleen). After a 
single dose of antigen, only animals immunized with either protein or 10/xg of 
liposomal DNA showed significant anti-OVA antibody titres by ELBA. After two 
doses of antigen, only animals immunized with either protein or liposomal DNA 
(both 2.5 and 10 ttg DNA) showed significant levels of seroconversion and serum 
antibody titres against OVA by ELBA.30 Similarly, no anti-OVA CTL activity was 
detected in animals immunized with DNA alone. However, animals immunized 
with two doses of 10 /xg of liposomal DNA displayed a CTL response higher (60% 
cell killing vs 50%) than that obtained in the positive control group immunized 
Genetic Vaccines: A Role for Liposomes 51 
with OVA protein and adjuvant (CT).30 Thus, delivery of a small dose of liposomal 
plasmid DNA subcutaneously, a route of immunization not normally inducing 
significant plasmid DNA mediated immune activation,9 results in a strong antigenspecific 
cellular response which is greater than that achieved by higher doses of a 
conventional protein antigen together with a powerful adjuvant (CT). 
4. The Co-delivery Concept 
Proteins that are synthesized within a cell (e.g. from plasmid DNA having a 
mammalian-active promoter) are continuously sampled as peptides by the 
proteosome / class-I MHC antigen presenting pathway. Conversely, proteins that are 
acquired exogenously by antigen-presenting cells are sampled in an analogous way 
by the endosomal/MHC-class-II pathway. It follows that the delivery of both protein 
and plasmid-DNA-encoded forms of a protein antigen to the same individual 
antigen-presenting cell would result in the simultaneous presentation of the antigen 
via both class-I and class-II pathways, thereby providing an opportunity for synergy 
in the resulting immune response to the antigen. Several appropriate liposomal 
formulations were designed to test the "co-delivery" hypothesis, exploiting the 
advantages of the dehydration-rehydration liposome technology that entraps both 
DNA and protein immunogens efficiently. The formulations, described in Table 1, 
comprise various test and control permutations of plasmid DNA and protein, either 
free or entrapped (together or separately) in the liposomal vehicle. 
Immunization with DNA encoding the influenza haemagglutinin protein 
has been explored previously with naked31 or liposomally formulated DNA.32 
Although immune responses elicited by DNA alone were adequate to achieve protective 
efficacy against influenza virus challenge in preclinical studies, only weak 
anti-HA antibody responses were elicited.31 The present "co-delivery" concept was 
designed to rectify this deficiency of DNA-based influenza vaccines. In a series of 
experiments, plasmid DNA encoding the haemagglutinin (HA) antigen [referred to 
in Table 1 and Fig. 5 as DNA(ha)] of the influenza virus (A/Sichuan/87 or A/PR/8 
strains) was co-entrapped with the corresponding whole inactivated virus (referred 
to as HA) within the same liposomes using the dehydration-rehydration method 
(for details on lipid composition and method see Refs. 26 and 27). A variety of control 
preparations including liposomes co-entrapping irrelevant DNA (i.e. plasmid 
DNA encoding ovalbumin) with HA or irrelevant protein (i.e. ova) with DNA (ha), 
entrapping DNA(ha) or HA alone, a mixture of the latter two preparations, and 
a mixture of the naked DNA(ha) and HA were used to immunize mice. Results 
shown in Fig. 5 demonstrate that the "co-delivery" hypothesis formulation (comprising 
both HA and its corresponding DNA in the same liposomes), elicited a 
greater response than all other formulations at each time point in the series, and it 
52 Gregoriadis et al. 
Table 1 Liposomal formulations of DNA and protein used in immunization experiments. 
Dose (/ig/animal (0.2 ml S/O) 
Liposomes (samples 4.1 & 5.1) 
DNA and protein (mixed) 
DNA and protein (mixed) 
DNA and protein (mixed) 
DNA alone 
Protein alone 
Control (PBS) 
ha (10) 
ova (11) 
ha (10) 
ha (10) 
ha (10) 
ha (10) 
ova (11) 
ha (10) 
ha (10) 
HA (0.6) 
HA (0.6) 
OVA (0.76) 
HA (0.6) 
HA (0.6) 
OVA (0.76) 
HA (0.6) 
HA (0.6) 
HA (10) 
Plasmid DNA encoding the HA antigen [DNA(ha)] and the HA antigen (HA) were entrapped in liposomes 
either together (co-entrapped; sample 1.1) or separately in different formulations (sample 6.1) 
mixed before injection. In some formulations, DNA(ha) and HA were entrapped alone (samples 4.1 and 
5.1 respectively). In others, ovalbumin (OVA) and plasmid DNA encoding ha fDNA(ha)] (sample 7) 
or HA and plasmid DNA encoding OVA (sample 8) were entrapped separately and then mixed. Mice 
were injected subcutaneously on days 0 and 28 and blood samples analyzed by ELISA for Ig responses. 
1OD0O - 
Ig response 
;r***OM burton) 
DNA (10 (ig) / Protein (0.6 ng) 
-  - Up(DNA(HA)/HA) 
- * - Lip(ONA{OVA)/HA) 
- * - Up(DNA(HA)/OVA) 
-  - Up (DNA (HA)/no protein) 
-  - Up(noDNA/HA) 
- * - Up(DNA(H;)) + Up(HA) 
 DNA {HA )  OVA 
- * - DNA ( n  ) no protein 
- * - HP (protein alone) 
 control (negative J 
20 A 3 0 
40 50 
Day post 1st dose 
Fig. 5. Serum Ig endpoint titres in Balb/c mice immunized on days 0 and 28 with DNA 
and/or antigen formulations as described in Table 1 a nd bled at time intervals. 
Genetic Vaccines: A Role for Liposomes 53 
is by far the strongest response after a single dose. Notably, the formulation "Lip 
(OVA/ha)", which is a control for the CpG adjuvant effect of plasmid DNA,33 gave 
a response which was much lower than that of "co-delivery" with the appropriate 
homologous pair of HA DNA and protein. Likewise, Lip (HA/ova) (an inappropriate 
pairing according to the hypothesis), gave a markedly weaker response. Figure 5 
also demonstrates that separately entrapped HA DNA and protein (in neighbouring 
vesicles) gave rise to an inferior response, supporting the hypothesis that delivery 
of both payloads to the same cell (which is best achieved by co-entrapment 
in the same liposome) is important in achieving the optimal antibody response. It 
is also remarkable that, inspite the modest DNA dose (10 /xg) and small number 
(2) of immunizations used, several formulations completely failed to generate an 
anti-HA response. These included HA DNA alone, and liposomally entrapped HA 
DNA. These findings serve to emphasize the striking degree of superiority of "codelivery" 
over previous methods of DNA-based immunization against influenza 
In conclusion, the present studies demonstrate that very small doses of protein 
as an additive in DNA immunization can dramatically improve the antibody 
response to the target protein, provided that the protein and DNA are homologous 
to one-another (i.e. that the DNA can express the protein), and that the payloads 
are delivered in the same individual liposomal vehicle. The simplest hypothesis 
to explain our observation is that the synergy observed between the appropriately 
delivered "homologous pair" of protein and DNA involves delivery of both 
payloads to the same antigen-presenting cell. The application of the co-delievery 
concept to alternative delivery systems, e.g. niosomes, dendimers, PLA/PLGA, chitosans, 
alginates and other microparticles awaits investigation. It is anticipated that 
the "co-delivery" approach will lead to better DNA-based vaccines for prophylactic 
and therapeutic use, particularly where vaccines require the elicitation of antibody 
responses (e.g. influenza vaccines). 
1. Powel MF and Newman MJ (eds.) (1995) Vaccine Design: The Subunit and Adjuvant 
Approach. Plenum Press: New York. 
2. Gregoriadis G, McCormack B, Allison AC and Poste G (eds.) (1993) New Generation 
Vaccines: The Role of Basic Immunology. Plenum Press: New York. 
3. Gregoriadis G (1990) Immunological adjuvants: A role for liposomes. Immunol Today. 
4. Gluck R, Mischler R, Brantschen S, Just M, Althans B and Cryz SJ, Jr (1992) Immunopotentiating 
reconstituted influenza virome vaccine delivery system for immunization 
against hepatitis A. / Clin Invest 90:2491-2495. 
54 Gregoriadis et al. 
5. Davis HL, Whalen RG and Demeneix BA (1993) Direct gene transfer in skeletal muscle 
in vivo: Factors influencing efficiency of transfer and stability of expression. Hum Gene 
Ther 4:151-156. 
6. Manickan E, Karem KL and Rouse BT (1997) DNA vaccines  A modern gimmick or a 
boon to vaccinology? Crit Rev Immunol 17:139-154. 
7. Chattergoon M, Boyer J and Weiner DB (1997) Genetic immunization: A new era in 
vaccines and immune therapeutics. FASEB 11:754-763. 
8. Gregoriadis G (1998) Genetic vaccines: Strategies for optimization. Pharm Res 15:661-670. 
9. Lewis PJ and Babiuk LA (1999) DNA vaccines: A review. Adv Virus Res 54:129-188. 
10. Tang DC, Devit M and Johnston SA (1992) Genetic immunization is a simple method for 
eliciting an immune response. Nature 356:152-154. 
11. Ulmer JB, Donnelly J, Parker SE, et al. (1993) Heterologous protection against influenza 
by injection of DNA encoding a viral protein. Science 259:1745-1749. 
12. Fynan EF> Webster RG, Fuller DH and Haynes JR (1993) DNA vaccines: Protective immunizations 
by parenteral, mucosal and gene-gun inoculations. Proc Natl Acad Sci USA 
13. Wang B, Ugen K, Srikantan V, et al. (1993) Gene inoculation generates immune responses 
against HIV-I. Proc Natl Acad Sci USA 90:4156^160. 
14. Davis HL, Michel ML, Mancini M, Schleef M and Whalen RG (1994) Direct gene transfer 
in skeletal muscle: Plasmid DNA based immunization against the hepatitis B virus 
surface antigen. Vaccine 12:1503-1509. 
15. Conry R, LoBuglio A, Loechel F, et al. (1995) A carcinoembryonic antigen polynucleotide 
vaccine for human clinical use. Cancer Gene Ther 2:33-38. 
16. Bright RK, Beames B, Shearer MH and Kennedy RC (1996) Protection against lethal 
tumor challenge with SV40-transformed cells by the direct injection of DNA encoding 
SV-40 large tumor antigen. Cancer Res 56:1126-1130. 
17. Matzinger P (1994) Tolerance, danger and the extended family. Annu Rev Immunol 12: 
18. Spier E (1996) Meeting Report: International meeting on the nucleic acid vaccines for 
the prevention of infectious disease and regulatory nuclear acid (DNA) vaccines. Vaccine 
19. Gregoriadis G, Saffie R and de Souza B (1997) Liposome-mediated DNA vaccination. 
FEES Lett 402:107-110. 
20. Gregoriadis G, Saffie R and Hart SL (1996) High yield incorporation of plasmid DNA 
within liposomes: Effect on DNA integrity and transfection efficiency. / Drug Targ 3: 
21. Gregoriadis G (1995) Engineering targeted liposomes: Progress and problems. Trends 
Biotechnol 13:527-537. 
22. Gregoriadis G, McCormack B, Obrenovic M and Perrie Y (1999) Entrapment of plasmid 
DNA vaccines into liposomes by dehydration/rehydration, in Lowrie DB and Whalen R. 
(eds.) Methods in Molecular Medicine, DNA Vaccines: Methods and Protocols. Humana Press 
Inc.: Totowa, NJ. pp. 305-312. 
Genetic Vaccines: A Role for Liposomes 55 
23. Gregoriadis G, McCormack B, Obrenovic M, Saffie R, Zadi B and Perrie Y (1999) Liposomes 
as immunological adjuvants and vaccine carriers. Methods 19:156-162. 
24. Zadi B and Gregoriadis G (2000) A novel method for high-yield entrapment of solutes 
into small liposomes. J Lipos Res 10:73-80. 
25. Feigner PL and Rhodes G (1991) Gene therapeutics. Nature 349:351-352. 
26. Perrie Y and Gregoriadis G (2000) Liposome-entrapped plasmid DNA: Characterization 
studies. Biochim Biphys Acta 1475:125-132. 
27. Perrie Y and Gregoriadis G (2001) Liposome mediated DNA vaccination: The effect of 
vesicle composition. Vaccine 19:3301-3310. 
28. Velinova M, Read N, Kirby C and Gregoriadis G (1996) Morphological observations 
on the fate of liposomes in the regional lymphs nodes after footpad injection into rats. 
Biochim Biophys Acta 1299:207-215. 
29. Szoka FC, Xu Y and Zelpati O (1996) How are nucleic acids released in cells from cationic 
lipid-nucleic acid-complexes? / Lipos Res 6:567-587. 
30. Bacon A, Caparros-Wanderley W, Zadi B and Gregoriadis G (2002) Induction of a cytotoxic 
T lymphocyte (CTL) response to plasmid DNA delivered by Lipodine. / Lipos 
Res 12:173-183. 
31. Johnson PA, Conwey MA, Daly J, Nicolson C, Robertson J and Mills KH (2000) Plasmid 
DNA encoding influenza virus haemagglutinin induces Th 1 cells and protection against 
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Polymer Micelles as Drug Carriers 
Elena V. Batrakova, Tatiana K. Bronich, 
Joseph A. Vetro and Alexander V. Kabanov 
1. Introduction 
It has long been recognized that improving one or more of the intrinsic adsorption, 
distribution, metabolism, and excretion (ADME) properties of a drug is a critical 
step in developing more effective drug therapies. As early as 1906, Paul Ehrlich 
proposed altering drug distribution by conjugating toxic drugs to "magic bullets" 
(antibodies) having high affinity for cancer cell-specific antigens, in order to both 
improve the therapeutic efficacy of cancer while decreasing its toxicity.1 Since then, 
it has become clear that directly improving intrinsic ADME through modifications 
of the drug is limited or precluded by structural requirements for activity. In other 
words, low molecular mass drugs are too small and have only limited number of 
atomic groups that can be altered to improve ADME, which often adversely affects 
drug pharmacological activity. In turn, the modifications of many low molecular 
mass drugs, aimed to increase their pharmacological activity, often adversely 
affect their ADME properties. For example, the potency and specificity of drugs 
can be improved by the addition of hydrophobic groups.2 The associated decrease 
in water solubility, however, increases the likelihood of drug aggregation, leading 
to poor absorption and bioavailability during oral administration2 or lowered systemic 
bioavailability, high local toxicity, and possible pulmonary embolism during 
intravenous administration.3 
Although there have been considerable difficulties for improving some existing 
drugs through chemical modifications, the problem became even more obvious 
58 Batrakova et al. 
with the development of high-throughput drug discovery technologies. Almost 
half of lead drug candidates identified by high-throughput screening have poor 
solubility in water, and are abandoned before the formulation development stage.4 
In addition, newly synthesized drug candidates often fail due to poor bioavailability, 
metabolism and/or undesirable side effects, which together decrease the 
therapeutic index of the molecules. Furthermore, a new generation of biopharmaceuticals 
and gene therapy agents are emerging based on novel biomacromolecules, 
such as DNA and proteins. The use of these biotechnology-derived drugs is completely 
dependent on efficient delivery to the critical site of the action in the body. 
Therefore, drug delivery research is essential in the translation of newly discovered 
molecules into potent drug candidates and can significantly improve therapies of 
existing drugs. 
Polymer-based drugs and drug delivery systems emerged from the laboratory 
bench in the 1990s as a promising therapeutic strategy for the treatment of certain 
devastating human diseases.5'6 A number of polymer therapeutics are presently 
on the market or undergoing clinical evaluation to treat cancer and other diseases. 
Most of them are low molecular weight drug molecules or therapeutic proteins that 
are chemically linked to water-soluble polymers to increase drug solubility, drug 
stability, or enable targeting to tumors. 
Recently, as a result of rapid development of novel nanotechnology-derived 
materials, a new generation of polymer therapeutics has emerged, using materials 
and devices of nanoscale size for the delivery of drugs, genes, and imaging 
molecules.7-12 These materials include polymer micelles, polymer-DNA complexes 
("polyplexes"), liposomes, and other nanostructured materials for medical use that 
are collectively known as nanomedicines. Compared with first generation polymer 
therapeutics, the new generation nanomedicines are more advanced. They 
entrap small drugs or biopharmaceutical agents such as therapeutic proteins and 
DNA, and can be designed to trigger the release of these agents at the target 
site. Many nanomedicines are constructed using self-assembly principles such as 
the spontaneous formation of micelles or interpolyelectrolyte complexes, driven 
by diverse molecular interactions (hydrophobic, electrostatic, etc.). This chapter 
considers polymeric micelles as an important example of the new generation of 
nanomedicines, which is also perhaps among the most advanced approach toward 
clinical applications in diagnostics and the treatment of human diseases. 
2. Polymer Micelle Structures 
2.1. Self-assembled micelles 
Self-assembled polymer micelles are created from amphiphilic polymers that 
spontaneously form nanosized aggregates when the individual polymer chains 
Polymer Micelles as Drug Carriers 59 
Single polymer chains Polymeric micelle 
Fig. 1. Self-assembly of block copolymer micelles. 
("unimers") are directly dissolved in aqueous solution (dissolution method)13 
above a threshold concentration (critical micelle concentration or CMC) and solution 
temperature (critical micelle temperature or CMT) (Fig. 1). Amphiphilic polymers 
with very low water solubility can alternatively be dissolved in a volatile 
organic solvent, then dialyzed against an aqueous buffer (dialysis method).14 
Amphiphilic di-block (hydrophilic-hydrophobic) or tri-block (hydrophilichydrophobic-
hydrophilic) copolymers are most commonly used to prepare selfassembled 
polymer micelles for drug delivery,9'15,16 although the use of graft 
copolymers has been reported.17-19 For drug delivery purposes, the individual 
unimers are designed to be biodegradable20,21 and/or have a low enough molecular 
mass (< ~40 kDa) to be eliminated by renal clearance, in order to avoid polymer 
buildup within the body that can potentially lead to toxicity.22 The most developed 
amphiphilic block copolymers assemble into spherical core-shell micelles approximately 
10 to 80 nm in diameter,23 consisting of a hydrophobic core for drug loading 
and a hydrophilic shell that acts as a physical ("steric") barrier to both micelle 
aggregation in solution, and to protein binding and opsonization during systemic 
administration (Fig. 2). 
The most common hydrophilic block used to form the hydrophilic shell 
is the FDA-approved excipient poly(ethylene glycol) (PEG) or poly(ethylene 
oxide) (PEO).24 PEG or PEO consists of the same repeating monomer subunit 
CH2-CH2-O, and may have different terminal end groups, depending on the 
synthesis procedure, e.g. hydroxyl group HO-(CH2-CH2-0)n-H; methoxy group 
CH30-(CH2-CH2-0)n-H, etc. PEG/PEO blocks typically range from 1 to 15 kDa.16,24 
In addition to its FDA approval, PEG is extremely soluble and has a large 
excluded volume. This makes it especially suitable for physically interfering with 
intra-micelle interactions and subsequent micelle aggregation. PEG also blocks 
protein and cell surface interactions, which greatly decreases nanoparticle uptake 
by the reticuloendothelial system (RES), and consequently increases the plasma 
60 Batrakova etal. 
No self assembly 
A n n n 
Di-block copolymer 
Tri-block copolymer 
Graft copolymer 
i n n n***** + ^ ^ , "w>>2^i?^s /^' 
Charged copolymer / ? S 
(unimolecular micelles) 
Star Dendritic 
hydrophilic block 
hydrophobic block 
cation ic block 
anionic block 
^ ^ H 
i i i i . 
Fig. 2. Polymer micelle structures. 
half life of the polymer micelle.25 The degree of steric protection by the hydrophilic 
shell is a function of both the density and length of the hydrophilic PEG blocks.25 
Unlike the hydrophilic block, which is typically PEG or PEO, different 
types of hydrophobic blocks have been sufficiently developed as hydrophobic 
drug loading cores.16 Examples of diblock copolymers include (a) poly(L-amino 
acids), (b) biodegradable poly(esters), which includes poly(glycolic acid), poly(D 
lactic acid), poly(D,L-lactic acid), copolymers of lactide/glycolide, and poly(ecaprolactone), 
(c) phospholipids/long chain fatty acids26; and for tri-block 
copolymers, (d) polypropylene oxide (in Pluronics/poloxamers).9 The choice of 
hydrophobic block is largely dictated by drug compatibility with the hydrophobic 
core (when drug is physically loaded, as described later) and the kinetic stability of 
the micelle. 
The self-assembly of amphiphilic copolymers is a thermodynamic and, consequently, 
a reversible process that is entropically driven by the release of ordered 
water from hydrophobic blocks; it is either stabilized or destabilized by solvent 
interactions with the hydrophilic shell. As such, the structural potential of 
amphiphilic copolymer unimers to form micelles is determined by the mass ratio 
of hydrophilic to hydrophobic blocks, which also affects the subsequent morphology 
if aggregates are formed.14 If the mass of the hydrophilic block is too great, 
the copolymers exist in aqueous solution as unimers, whereas, if the mass of the 
hydrophobic block is too great, unimer aggregates with non-micellar morphology 
are formed.27 If the mass of the hydrophilic block is similar or slightly greater than 
the hydrophobic block, then conventional core shell micelles are formed. 
An important consideration for drug delivery is the relative thermodynamic 
(potential for disassembly) and kinetic (rate of disassembly) stability of the polymer 
Polymer Micelles as Drug Carriers 61 
micelle complexes, after intravenous injection and subsequent extreme dilution in 
the vascular compartment.28 This is because the polymer micelles must be stable 
enough to avoid burst release of the drug cargo, as in the case of a physically loaded 
drug, upon systemic administration and remain as nanoparticles long enough to 
accumulate in sufficient concentrations at the target site. 
The relative thermodynamic stability of polymer micelles (which is inversely 
related to the CMC) is primarily controled by the length of the hydrophobic block.13 
An increase in the length of the hydrophobic block alone significantly decreases the 
CMC of the unimer construct (i.e. increases the thermodynamic stability of the polymer 
micelle), whereas an increase in the hydrophilic block alone slightly increases 
the CMC (i.e. decrease the thermodynamic stability of a polymer micelle).14 
Although the CMC indicates the unimer concentration below which polymer 
micelles will begin to disassemble, the kinetic stability determines the rate 
at which polymer micelle disassembly occurs. Many diblock copolymer micelles 
possess good kinetic stability and only slowly dissociate into unimers after extreme 
dilution.29 Thus, although polymer micelles are diluted well below typical unimer 
CMCs29 (10~6-10-7M) after intravenous injection, their relative kinetic stability 
might still be suitable for drug delivery. The kinetic stability depends on several 
factors, including the size of a hydrophobic block, the mass ratio of hydrophilic to 
hydrophobic blocks, and the physical state of the micelle core.14 The incorporation 
of hydrophobic drugs may also further enhance micelle stability. 
2.2. Unimolecular micelles 
Unimolecular micelles are topologically similar to self-assembled micelles, but consist 
of single polymer molecules with covalently linked amphiphile chains. For 
example, copolymers with star-like or dendritic architecture, depending on their 
structure and composition, can either aggregate into multimolecular micelles,30-32 
or exist as unimolecular micelles.33 Dendrimers are widely used as building blocks 
to prepare unimolecular micelles, because they are highly-branched, have welldefined 
globular shape and controled surface functionality.34-40 For example, unimolecular 
micelles were prepared by coupling dendritic hypercores of different 
generations with PEO chains.40'41 The dendritic cores can entrap various drug 
molecules. However, due to the structural limitations involved in the synthesis 
of dendrimers of higher generation, and relatively compact structure of the dendrimers, 
the loading capacity of such micelles is limited. Thus, to increase the loading 
capacity, the dendrimer core can be modified with hydrophobic block, followed 
by the attachment of the PEO chains. For example, Wang et al. recently synthesized 
an amphiphilic 16-arm star polymer with a polyamidoamine dendrimer core 
and arms composed of inner lipophilic poly(e-caprolactone) block and outer PEO 
62 Batrakova et al. 
block.42 These unimolecular micelles were shown to encapsulate a hydrophobic 
drug, etoposide, with high loading capacity. 
Multiarm star-like block copolymers represent another type of unimolecular 
micelles.42-46 Star polymers are generally synthesized by either the arm-first or 
core-first methods. In the arm-first method, monofunctional living linear macromolecules 
are synthesized and then cross-linked either through propagation, using 
a bifunctional comonomer,47 or by adding a multifunctional terminating agent 
to connect precise number of arms to one center.45 Conversely, in the core-first 
method, polymer chains are grown from a multifunctional initiator.43'44'46'48 One 
of the first reported examples of unimolecular micelles, suitable for drug delivery, 
was a three-arm star polymer, composed of mucic acid substituted with fatty acids 
as a lipophilic inner block and PEO as a hydrophilic outer block.44 These polymers 
were directly dispersible in aqueous solutions and formed unimolecular micelles. 
The size and solubilizing capacity of the micelles were varied by changing the ratio 
of the hydrophilic and lipophilic moieties. In addition, star-copolymers with polyelectrolyte 
arms can be prepared to develop pH-sensitive unimolecular micelles as 
drug carriers.46 
2.3. Cross-linked micelles 
The multimolecular micelles structure can be reinforced by the formation of crosslinks 
between the polymer chains. These resulting cross-linked micelles are, in 
essence, single molecules of nanoscale size that are stabile upon dilution, shear 
forces and environmental variations (e.g. changes in pH, ionic strength, solvents 
etc.). There are several reports on the stabilization of the polymer micelles by crosslinking 
either within the core domain49-53 or throughout the shell layer.54-56 In 
these cases, the cross-linked micelles maintained small size and core-shell morphology, 
while their dissociation was permanently suppressed. Stable nanospheres 
from the PEO-b-polylactide micelles were prepared by using polymerizable group 
at the core segment.49 In addition to stabilization, the core polymerized micelles 
readily solubilized rather large molecules such as paclitaxel, and retained high 
loading capacity even upon dilution.50 Formation of interpenetrating network 
of a temperature-sensitive polymer (poly-N-isopropylacrilomide) inside the core 
was also employed for the stabilization of the Pluronic micelles.53 The resulting 
micelle structures were stable against dilution, exhibited temperature-responsive 
swelling behavior, and showed higher drug loading capacity than regular Pluronic 
Recently, a novel type of polymer micelles with cross-linked ionic cores was 
prepared by using block ionomer complexes as templates.57 The nanofabrication of 
these micelles involved condensation of PEO-b-poly(sodium methacrylate) diblock 
Polymer Micelles as Drug Carriers 63 
copolymers by divalent metal cations into spherical micelles of core-shell morphology. 
The core of the micelle was further chemically cross-linked and cations 
removed by dialysis. Resulting micelles represent hydrophilic nanospheres of coreshell 
morphology. The core comprises a network of the cross-linked polyanions 
and can encapsulate oppositely charged therapeutic and diagnostic agents, while a 
hydrophilic PEO shell provides for increased solubility. Furthermore, these micelles 
displayed the pH- and ionic strength-responsive hydrogel-like behavior, due to the 
effect of the cross-linked ionic core. Such behavior is instrumental for the design of 
drug carriers with controled loading and release characteristics. 
3. Drug Loading and Release 
In general, there are three major methods for loading drugs into polymer micelle 
cores: (1) chemical conjugation, (2) physical entrapment or solubilization, and 
(3) polyionic complexation (e.g. ionic binding). 
3.1. Chemical conjuga tion 
Drug incorporation into polymer micelles via chemical conjugation was first proposed 
by Ringsdorf's group58 in 1984. According to this approach, a drug is chemically 
conjugated to the core-forming block of the copolymer via a carefully designed 
pH- or enzyme-sensitive linker, that can be cleaved to release a drug in its active 
form within a cell.59,60 The polymer-drug conjugate then acts as a polymer prodrug 
which self assembles into a core-shell structure. The appropriate choice of 
conjugating bond depends on specific applications. 
The nature of the polymer-drug linkage and the stability of the drug conjugate 
linkage can be controled to influence the rate of drug release, and therefore, the 
effectiveness of the prodrug.61-63 For instance, recent work by Kataoka's group proposed 
pH-sensitive polymer micelles of PEO-b-poly(aspartate hydrazone doxorubicin), 
in which doxorubicin was conjugated to the hydrophobic segments through 
acid-sensitive hydrazone linkers that are stable at extracellular pH 7.4, but degrade 
and release the free drug at acidic pH 5.0 to 6.0 in endosomes and lysosomes.63,64 
The original approach developed by this group used doxorubicin conjugated to the 
poly(aspartic acid) chain of PEO-b-poly(aspartic acid) block copolymer through an 
amide bond.65 Adjusting both the composition of the block copolymer and the concentration 
of the conjugated doxorubicin, led to improved efficacy, as evidenced by 
a complete elimination of solid tumors implanted in mice.66 It was later determined 
that doxorubicin physically encapsulated within the micellar core was responsible 
for antitumor activity. This finding led to the use of PEO-b-poly(aspartate doxorubicin) 
conjugates as nanocontainers for physically entrapped doxorubicin.67 
64 Batrakova et al. 
3.2. Physical entrapment 
The physical incorporation or solublization of drugs within block copolymer 
micelles is generally preferred over micelle-forming polymer-drug conjugates, 
especially for hydrophobic drug molecules. Indeed, many polymers and drug 
molecules do not contain reactive functional groups for chemical conjugation, 
and therefore, specific block copolymers have to be designed for a given type 
of drug. In contrast, a variety of drugs can be physically incorporated into the 
core of the micelles, by engineering the structure of the core-forming segment. In 
addition, molecular characteristics (i.e. molecular weight, composition, presence 
of functional groups for active targeting) within a homologous copolymer series 
can be designed to optimize the performance of a drug for a given drug delivery 
situation.9,14 This concept was introduced by our group in the late 1980s and was 
initially termed "micellar microcontainer",68 but is now widely known as a "micellar 
nanocontainer".9,10 Haloperidol was encapsulated in Pluronic block copolymer 
micelles,68 the micelles were targeted to the brain using brain-specific antibodies 
or insulin, and enhancement of neuroleptic activity by the solubilized drug was 
observed. During the last 25 years, a large variety of amphiphilic block copolymers 
have been explored as nanocontainers for various drugs. 
Different loading methods can be used for physical entrapment of the drug into 
the micelles, including but not limited to dialysis,69-72 oil in water emulsification,69 
direct dissolution,42,73,74 or solvent evaporation techniques.75,76 Depending on the 
method, drug solubilization may occur during or after micelle assembly. The loading 
capacity of the polymer micelles, which is frequently expressed in terms of the 
micelle-water partition coefficient, is influenced by several factors, including both 
the structure of core-forming block and a drug, molecular characteristics of the 
copolymer such as composition, molecular weight, and the solution temperature.13 
Many studies indicate that the most important factor related to the drug solubilization 
capacity of a polymer micelle is the compatibility between the drug and 
the core-forming block.9,14,77-80 For this reason, the choice of the core-forming block 
is most critical. One parameter that can be used to assess the compatibility between 
the polymer and a drug is the Flory-Huggins interaction parameter, Xsp/ defined as 
Xsp= (Ss - <5p)2Vs/kT; where Ss and <5p are Scatchard-Hildebrand solubility parameters, 
and Vs is the molecular volume of the solubilizate. It was successfully used as a 
correlation parameter for the solubilization of aliphatic and aromatic hydrocarbons 
in block copolymer micelles.80,81 Recently, Allen's group82 elegantly demonstrated 
that the calculation and comparison of partial solubility parameters of polymers 
and drugs could be used as a reliable means to predict polymer-drug compatibility 
and to guide formulation development. Polymer micelles, possessing core-forming 
blocks predicted to be compatible with the drug of interest (Ellipticine), were able 
Polymer Micelles as Drug Carriers 65 
to increase the solubility of the drug up to 30,000 times, compared with its saturation 
solubility in water.82 The degree of compatibility between the drug and the 
core-forming block has also been shown to influence the release rate of the drug 
from the micelles. When the environment within the core of the micelle becomes 
more compatible with the drug, it results in a considerable decrease in the rate of 
drug release. 
For a given drug, the extent of incorporation is a function of factors that also 
control the micelle size and/or aggregation number. Such factors include the ratio of 
hydrophobic to hydrophilic block length and the copolymer molecular weight. For 
example, the loading capacity of Pluronic micelles was found to increase with the 
increase in the hydrophobic PPO block length. This effect is attributed to a decrease 
in CMC, and therefore, an increase in aggregation number and micelle core size. 
Also, but to a lesser extent, the hydrophilic block length affects the extent of solubilization, 
such that an increase in percentage of PEO in Pluronic block copolymers 
results in a decrease in the loading capacity of the micelles.80,83-85 For a given ratio of 
PPO-to-PEO, higher molecular weight polymers form larger micelles, and therefore, 
show a higher drug loading capacity. Therefore, the total amount of loaded drug can 
be adjusted as a function of the micellar characteristics as clearly was demonstrated 
by Nagaradjan83 and Kozlov et al.85 Several studies indicate that both the copolymer 
concentration as well as the drug to polymer ratio upon loading, have a complex 
effect on the loading capacity of polymer micelles.79,84,86 In general, more polymer 
chains provide more absorption sites. As a result, solubilization is increased 
with polymer concentration.82 However, the solubilization capacity was found to 
reach a saturation level with an increase of polymer concentration.79 The maximum 
loading level is largely influenced by the interaction between the solubilizate and 
core-forming block, and stronger interactions enable saturation to be reached at 
lower polymer concentration. It was also demonstrated in the studies by Hurter 
and Hatton84'86 that the loading capacity of micelles formed from copolymers with 
high hydrophobic content was independent of the polymer concentration. In addition, 
the location of the incorporated molecules within polymer micelles (micelle 
core or the core-shell interface) determines the extent of solubilization, as well as the 
rate of drug release.87,88 It has been found that more soluble compounds are localized 
at the core-shell interface or even in the inner shell, whereas more hydrophobic 
molecules have a tendency to solubilize in the micelle core.85,87,88 The release rate 
of drug localized in the shell or at the interface appears to account for the "burst 
release" from the micelles.87 In general, for drugs physically incorporated in polymer 
micelles, release is controled by the rate of diffusion of the drug from the micellar 
core, stability of the micelles, and the rate of biodegradation of the copolymer. 
If the micelle is stable and the rate of polymer biodegradation is slow, the diffusion 
rate of the drug will be mainly determined by the abovementioned factors, 
66 Batrakova et al. 
i.e. the compatibility between the drug and core forming block of copolymer,69,82 
the amount of drug loaded, the molecular volume of drug, and the length of the 
core forming block.89 In addition, the physical state of the micelle core and drug 
has a large influence on release characteristics. It was demonstrated that the diffusion 
of incorporated molecules from the block copolymer micelles with glassy 
cores is slower, in comparison to the diffusion out of the cores that are more 
3.3. Poly ionic complexation 
Charged therapeutic agents can be incorporated into block copolymer micelles, 
through electrostatic interactions with an oppositely charged ionic segment of block 
copolymer. Since it was being proposed independently by Kabanov and Kataoka 
in 1995,90,91 this approach is now widely used for the incorporation of various 
polynucleic acids into block ionomer complexes, for developing non-viral gene 
delivery systems. Ionic block lengths, charge density, and ionic strength of the 
solution affect the formation of stable block ionomer complexes, and therefore, 
control the amount of drug that can be incorporated within the micelles.8'92 The pHand 
salt-sensitivity of such block ionomer micelles provide a unique opportunity to 
control the triggered release of the active therapeutic agent.1563,93-96 Furthermore, 
block ionomer complexes can participate in the polyion interchange reactions which 
are believed to account for the release of the therapeutic agent and DNA in an active 
form inside cells.7 Several comprehensive reviews can be found in the literature that 
focus on block ionomer micelles as drug and gene delivery systems.8,92 In addition, 
physicochemical aspects of the DNA complexes with cationic block copolymers 
have also been recently reviewed.97 
As an example, the metal-complex formation of ionic block copolymer, PEOb-
poly(L-aspartic acid), was explored to prepare polymer micelles incorporating 
cz's-dichlorodiamminoplatinum (II) (CDDP);98,99 a potent chemotherapeutic agent 
widely used in the treatment of a variety of solid tumors, particularly, testicular, 
ovarian, head and neck, and lung tumors.100,101 The CDDP-loaded micelles 
had a size of approximately 20 nm. These micelles showed remarkable stability 
upon dilution in distilled water, while in physiological saline, they displayed sustained 
release of the regenerated Pt complex over 50hrs, due to inverse ligand 
exchange from carboxylate to chloride. The release rate was inversely correlated 
with the chain length of poly(L-aspartic acid) segments in the block copolymer. 
The stability of CDDP-loaded micelle against salt was shown to be improved by 
the addition of homopolymer, poly(L-aspartic acid), in the micelles.102 Recently, 
CDDP-loaded micelles were newly prepared using another block copolymer, 
PEO-b-poly(glutamic acid) to improve and optimize the micellar stability, as well 
Polymer Micelles as Drug Carriers 67 
as the drug release profile.103 The drug loading in the micelles was as high as 39% 
(w/w), and these micelles released the platinum in physiological saline at 37C in 
sustained manner > 150 hrs, without initial burst of the drug. 
The principle of polyionic complexation can also be used to design new photosensitizers 
for photodynamic therapy of cancer. The group of Kataoka reported 
formation of micelles, as a result of mixing of oppositely charged dendrimer porphyrin 
and block ionomer, based on electrostatic assembly104 or combination of 
electrostatic and hydrogen bonding interactions.95'105 The micelles were stabile at 
physiological conditions and released the entrapped dendrimers in the acidic pH 
environment (pH 5.0), suggesting a possibility of pH-triggered drug release in the 
intracellular endosomal compartments. Overall, the photodynamic efficacy of the 
dendrimer porphyrins was dramatically improved by inclusion into micelles. This 
process resulted in more than two orders of magnitude increase in the photocytotoxicity, 
compared with that of the free dendrimer porphyrins. 
In addition, the polyionic complexation has been used to immobilize charged 
enzymes such as egg white lysozyme106 or trypsin,107 which were incorporated 
in the core of polyion micelles, after mixing with oppositely charged ionic block 
copolymer. A remarkable enhancement of enzymatic activity was observed in 
the core of the micelles. Furthermore, the on-off switching of the enzyme activity 
was achieved through the destabilization of the core domain by applying a 
pulse electric field.108 These unique features of the polyion micelles are relevant 
for their use as smart nanoreactors in the diverse fields of medical and biological 
Last, but not the least, a special class of polyion complexes has been synthesized 
by reacting block ionomers with surfactants of opposite charge, resulting in the 
formation of environmentally responsive nanomaterials, which differ in sizes and 
morphologies, and include micelles and vesicles.109-113 These materials contain a 
hydrophobic core formed by the surfactant tail groups, and a hydrophilic shell 
formed, for example, by PEO chains of the block ionomer. These block ionomer 
complexes can incorporate charged surfactant drugs such as retinoic acid, as well 
as other drugs via solubilization in the hydrophobic domains formed by surfactant 
molecules.114 They display transitions induced by changes in pH, salt concentration, 
chemical nature of low molecular mass counterions, as well as temperature. They 
can also be fine tuned to respond to environmental changes occurring in a very wide 
range of conditions that could realize during delivery of biological and imaging 
agents.94115 The unique self-assembly behavior, the simplicity of the preparation, 
and the wide variety of available surfactant components that can easily produce 
polymer micelles with a very broad range of core properties, make this type of 
materials extremely promising for developing vehicles for the delivery of diagnostic 
and therapeutic modalities. 
68 Batrakova et al. 
4. Pharmacokinetics and Biodistribution 
Incorporation of a low molecular mass drug into polymer micelles drastically alters 
pharmacokinetics and biodistribution of the drug in the body, which is crucial 
for the drug action. Low molecular mass drugs, after administration in the body, 
rapidly extravasate to various tissues affecting them almost indiscriminately, and 
then are rapidly eliminated from the body via renal clearance, often causing toxicity 
to kidneys.116 Furthermore, many drugs display low stability and are degraded in 
the body, often forming toxic metabolites. An example is doxorubicinol, a major 
metabolite of doxorubicin, which causes cardiac toxicity.117 These impediments to 
the therapeutic use of low molecular mass drugs can be mitigated by encapsulating 
drugs in polymer micelles. Within the micelles, the drug molecules are protected 
from enzymatic degradation by the micelle shell. The pharmacokinetics and biodistribution 
of the micelle-incorporated drugs are mainly determined by the surface 
properties, size, and stability of the micelles, and are less affected by the properties 
of the loaded drug. The surface properties of the micelles are determined by 
the micelle shell. The shell from PEO effectively masks drug molecules and prevents 
interactions with serum proteins and cells, which contributes to prolonged 
circulation of the micelles in the body.16 From the size standpoint, polymer micelles 
fit an ideal range of sizes for systemic drug delivery. On the one hand, micelles 
are sufficiently large, usually exceeding 10 nm in diameter, which hinders their 
extravasation in nontarget tissues and prevents renal glomerular excretion. On the 
other hand, the micelles are not considered large, since their size usually does not 
exceed 100 nm. As a result, micelles avoid scavenging by the mononuclear phagocytes 
system (MPS) in the liver and spleen. To this end, "stealth" particles whose 
surface is decorated with PEO are known to be less visible to macrophages and 
have prolonged half-lives in the blood.64,118,119 
The contribution of the micelle stability to pharmacokintetics and biodistribution 
is much less understood, although it is clear that micelle degradation should 
result in a decrease of the size and drug release, perhaps, prematurely. Degradation 
of the micelles, resulting in the formation of block copolymer unimers, could also be 
a principal route for the removal of the polymer material from the body. The molecular 
mass of the unimers of most block copolymers is below the renal excretion limit, 
i.e. less than ~ 20 to 40 kDa,22,120121 while the molecular mass of the micelles, which 
usually contain several dozen or even hundreds of unimers molecules, is above 
this limit. Thus, the unimers are sufficiently small and can be removed via renal 
excretion, while the micelles cannot. A recent study by Batrakova et al. determined 
pharmacokinetic parameters of an amphiphilic block copolymer, Pluronic P85, 
and perhaps provided first evidence that the pharmacokinetic behavior of a block 
copolymer can be a function of its aggregation state.119 Specifically, the formation 
Polymer Micelles as Drug Carriers 69 
of micelles increased the half-life of the block copolymer in plasma and decreased 
the uptake of the block copolymer in the liver. However, it had no effect on the total 
clearance, indicating that the elimination of Pluronic P85 was controled by the renal 
tubular transport of unimers, but not by the rate of micelles disposition or disintegration. 
Furthermore, the values of the clearance suggested that a significant portion 
of the block copolymer was reabsorbed back into the blood, probably, through the 
kidney's tubular membranes. Chemical degradation of the polymers comprising 
the micelles, followed by renal excretion of the relatively low molecular mass products 
of degradation, may be another route for the removal of the micelle polymer 
material from the body. This route could be particularly important in the case of 
the cross-linked or unimolecular micelles, micelles displaying very high stability, 
and / or micelles composed from very hydrophobic polymer molecules that can bind 
and retain considerably biological membranes and other cellular components. 
The delivery of chemotherapeutic drugs to treat tumors is one of the most 
advanced areas of research using polymer micelles. Two approaches have been 
explored to enhance delivery of drug-loaded polymer micelles to the tumor sites: 
(1) passive targeting and (2) vectorized targeting. The passive targeting involves 
enhanced permeability and retention (EPR) effect.122,123 It is based on the fact that 
solid tumors display increased vascular density and permeability caused by angiogenesis, 
impaired lymphatic recovery, and lack of a smooth muscle layer in solid 
tumor vessels. As a result, micellar drugs can penetrate and retain in the sites of 
tumor lesions. At the same time, extravasation of micellar drugs in normal tissues 
is decreased, compared with low molecular drug molecules. Among normal 
organs, spleen and liver can accumulate polymer drugs, but the drugs are eventually 
cleared via the lymphatic system. The increased circulation time of the micellar 
drugs should further enhance exposure of the tumors to the micellar drug, compared 
with the low molecular mass drugs. Along with passive targeting, the delivery 
of micellar drugs to tumors can potentially be enhanced by the modification of 
the surface of the polymer micelles with the targeting molecules, vectors that can 
selectively bind to the surface of the tumor cells. Potential vectors include antibodies, 
aptamers and peptides, capable of binding tumor-specific antigens and other 
molecules diplayed at the surfaces of the tumors.124-126 
Altered biodistribution of a common antineoplastic agent was demonstrated 
for CDDP encapsulated in polyionic micelles with PEO-b-poly(glutamic acid) block 
copolymers.103 Free CDDP is rapidly distributed to each organ, where its levels 
peak at about one hr after i.v. administration. In contrast, in the case of the CDDPincorporated 
micelles, due to their remarkably prolonged blood circulation time, 
the drug level in the liver, spleen and tumor continued to increase up to at least 
24 hrs after injection. Consequently, the CDDP-incorporated micelle exhibited 4-, 
39-, and 20-fold higher accumulation in the liver, spleen and tumor respectively, 
70 Batrakova et al. 
than the free CDDP. At the same time, the encapsulation of CDDP into the micelles 
significantly decreased drug accumulation in the kidney, especially during first hr 
after administration. This suggested potential for the decrease of severe nephrotoxicity 
observed with the free drug, which is excreted through the glomerular 
filtration, thus affecting the kidney.127 
Promising results were also demonstrated for doxorubicin incorporated into 
styrene-maleic acid micelles.128 In this case, as a result of drug entrapment into 
micelles, the drug was redirected from the heart to the tumor, and the doxorubicin 
cardiotoxicity was diminished. Complete blood counts and cardiac histology for 
the micellar drug showed no serious side effects for i.v. doses as high as 100 mg/kg 
doxorubicin equivalent in mice. Similar results were reported for doxorubicin incorporated 
in mixed micelles of PEO-b-poly(L-histidine) and PEO-b-poly(L-lactic acid) 
block copolymers.129 Tissue levels of doxorubicin administered in the micellar formulation 
were decreased in the blood and the liver, and considerably increased in 
the solid tumor, compared with the free drug. Further increase in the tumor delivery 
was achieved by modifying the surface of the micelles with the folate molecules. 
The accumulated doxorubicin levels observed using folate-modified micelles was 
20 times higher than those for free doxorubicin, and 3 times higher than those for 
unmodified micelles. 
The first micellar formulation of doxorubicin to reach clinical evaluation stage, 
used the micelles composed of triblock copolymer, PEO-b-poly (propylene oxide)-b- 
PEO, Pluronic.130 Analysis of pharmacokinetics and biodistribution of doxorubicin 
incorporated into mixed micelles of Pluronics L61 and F127, SP1049C, demonstrated 
more efficient accumulation of the micellar drug in the tumors, compared with the 
free drug. Specifically, the areas under the curves (AUC) in the Lewis lung carcinoma 
3LL M-27 solid tumors in C57B1 / 6 mice were increased about two fold using SP1049, 
compared with the free doxorubicin. Furthermore, this study indicated that the peak 
levels of doxorubicin formulated with SP1049 in the tumor were delayed and the 
drug residence time was increased, in comparison with the free doxorubicin.130 
A clear visualization of drug delivery to the tumor site was shown for doxorubicin 
covalently incorporated through the pH-sensitive link into polymer micelles 
of PEO-poly(aspartate hydrazone doxorubicin).64 A phase-contrast image showed 
that the tumor blood vessels containing the micelles leaked into extra vascular compartments 
of the tumors, resulting in the infiltration of the micelles into tumor 
sites. The micelles circulated in the blood for a prolonged time, and the AUC for 
micellar doxorubicin was 15-fold greater than the AUC for the free doxorubicin. 
Furthermore, the AUC values of the micellar doxorubicin in the heart and kidney 
decreased, compared with the free drug. Thus, the selectivity of drug delivery to 
the tumor, compared with heart and kidney (AUCtumor/AUC0rgan) was increased 
by 6- and 5-folds respectively. This may result in the reduction of side effects of 
Polymer Micelles as Drug Carriers 71 
doxorubicin such as cardiotoxicity and nephrotoxicity. Moreover, the micellar doxorubicin 
showed relatively low uptake in the liver and spleen, despite very long 
residence time in the blood. 
Biodistribution of paclitaxel incorporated into biodegradable polymer micelles 
of monomethoxy-PEO-b-poly(D,L-lactide) block copolymer, Genexol-PM, was 
compared with the regular formulation of the drug in Cremophor EL.131 Two to 
three-fold increases in drug levels were demonstrated in most tissues including 
liver, spleen, kidneys, lungs, heart and tumor, after i.v. administration of Genexol- 
PM, compared with paclitaxel. Nevertheless, acute dose toxicity of Genexol-PM 
was about 25 times lower than that of the conventional drug formulation, which 
appears to be a result of the reformulation avoiding the use of Chremophor EL and 
dehydrated ethanol that are toxic. 
Selective tumor targeting with paclitaxel encapsulated in micelles, modified 
with tumor-specific antibodies 2C5 ("immunomicelles"), was reported using Lewis 
lung carcinoma solid tumor model in C57B1/6J mice.26 These micelles were prepared 
from PEO-distearyl phosphatidylethanolamine conjugates with the free PEO 
end activated with the p-nitrophenylcarbonyl group for the antibody attachment. 
The amount of micellar drug accumulated in the tumor exceeded that in the nontarget 
tissue (muscles) by more than ten times. It is worth noting that the highest 
accumulation in the tumor was demonstrated in the micelles containing the longest 
PEO chains, which also had the longest circulation time in the blood. Furthermore, 
the immunomicelles displayed the highest amount of tumor-accumulated drug, 
compared with either free paclitaxel or non-vectorized micelles. It was demonstrated 
that paclitaxel delivered by plain micelles in the interstitial space of the 
tumor was eventually cleared after gradual micellar degradation. In contrast, 
paclitaxel-loaded 2C5 immunomicelles were internalized by cancer cells and the 
retention of the drug inside the tumor was enhanced.132 
Unexpected results were found using pH-sensitive polymer micelles of Nisopropylacrylamide 
and methacrylic acid copolymers randomly or terminally 
alkylated with octadecyl groups.64,133 It was demonstrated that aluminium chloride 
phthalocyanine (AlClPc) incorporated in such micelles was cleared more rapidly 
and less accumulated in the tumor, than the AlClPc formulated with Cremophor 
EL. Furthermore, significant accumulation in the liver and spleen (and lungs for 
most hydrophobic copolymers) was observed, compared with Cremophor EL formulation. 
The enhanced uptake of such polymer micelles by the cells of mononuclear 
phagocyte system (MPS) could be due to micelle aggregation in the blood 
and embolism in the capillaries. Thus, it attempted to reduce the uptake of the 
micelles in MPS by incorporating water soluble monomers, N-vinyl-2-pyrrolidone 
in the copolymer structure.134 The modified formulation displayed same levels of 
tumor accumulation and somewhat higher antitumor activity than the Cremophor 
72 Batrakova et al. 
EL formulation. This work serves as an example reinforcing the need of proper 
adjustment of the polymer micelle structure, and perhaps the need of using block 
copolymers to produce a defined protective hydrophilic shell to facilitate evasion 
of the polymer micelles from MPS. 
5. Drug Delivery Applications 
The studies on the application of polymer micelles in drug delivery have mostly 
focused on the following areas that are considered below: (1) delivery of anticancer 
agents to treat tumors; (2) drug delivery to the brain to treat neurodegenerative diseases; 
(3) delivery of antifungal agents; (4) delivery of imaging agents for diagnostic 
applications; and (5) delivery of polynucleotide therapeutics. 
5.1. Chemotherapy of cancer 
To enhance chemotherapy of tumors using polymer micelles, four major approaches 
were employed: (1) passive targeting of polymer micelles to tumors due to EPR 
effect; (2) targeting of polymer micelles to specific antigens overexpressed at the 
surface of tumor cells; (3) enhanced drug release at the tumor sites having low pH; 
and (4) sensitization of drug resistant tumors by block copolymers. 
A series of pioneering studies by Kataoka's group used polymer micelles for 
passive targeting of various anticancer agents and chemotherapy of tumors.102,103'135 
One notable recent example reported by this group involves polymer micelles of 
PEO-b-poly(L-aspartic acid) incorporating CDDP. Evaluation of anticancer activity 
using murine colon adenocarcinoma C26 as an in vivo tumor model, demonstrated 
that CDDP in polymer micelles had significantly higher activity than the free CDDP, 
resulting in complete eradication of the tumor.103 A formulation of paclitaxel in 
biodegradable polymer micelles of monomethoxy-PEO-b-poly(D,L-lactide) block 
copolymer, Genexol-PM, also displayed elevated activity in vivo against human 
ovarian carcinoma OVCAR-3 and human breast carcinoma MCF7, compared with 
a regular formulation of the drug in Cremophor EL.131 In addition, anthracycline 
antibiotics, doxorubicin and pirarubicin, incorporated in styrene-maleic acid 
micelles each revealed potent anticancer effects in vivo against mouse sarcoma 
S-180, resulting in complete eradication of tumors in 100% of tested animals.128 
Notably, animals survived for more than one year, after treatment with the micelleincorporated 
pirarubicin at doses as high as lOOmg/kg of pirarubicin equivalent. 
Complete blood counts, liver function test, and cardiac histology showed no 
sign of adverse effects for intravenous doses of the micellar formulation. In contrast, 
animals receiving free pirarubicin had a much reduced survival and showed 
serious side effects.136 Collectively, these studies suggested that various micelleincorporated 
drugs display improved therapeutic index in solid tumors, which 
Polymer Micelles as Drug Carriers 73 
correlates with enhanced passive targeting of the drug to the tumor sites, as well as 
decreased side effects, compared with conventional formulations of these drugs. 
Tumor-specific targeting of polymer micelles to molecular markers expressed 
at the surface of the cancer cells has also been explored to eradicate tumor cells. 
For example, a recent study by Gao's group developed a polymer micelle carrier 
to deliver doxorubicin to the tumor endothelial cells with overexpressed Xvfi3 
integrins.137 A cyclic pentapeptide, cRGD was used as a targeting ligand that is 
capable of selective and high affinity binding to the Xvfio, integrin. Micelles of PEOb-
poly(e-caprolactone) loaded with doxorubicin were covalently bound with cRGD. 
As a result of such modification, the uptake of doxorubicin-containing micelles in 
in vitro human endothelial cell model derived from Kaposi's sarcoma, was profoundly 
increased. In addition, folate receptor often overexpressed in cancer cells 
has been evaluated for targeting various drug carriers to tumors.138 This strategy 
has also been evaluated to target polymer micelles. For example, mixed micelles 
of PEO-b-poly(L-histidine) and PEO-b-poly(L-lactic acid) block copolymers with 
solubilized doxorubicin129 or micelles of PEO-b-poly(DL-lactic-co-glycolic acid) 
block copolymer with covalently attached doxorubicin,139 were each surface modified 
by conjugating folate molecules to the free PEO ends. In both cases, in vitro 
and in vivo studies demonstrated increased antitumor activity of the micelleincorporated 
drug resulting from such modification. The enhanced delivery of the 
micellar drugs through the folate receptor, and the enhanced retention of the modified 
micelles at the tumor sites are possible explanations for the effects of these folate 
Micelles conjugated with antibodies or antibody fragments capable to 
recognize tumor antigens were shown to improve therapeutic efficacy in vivo over 
non-modified micelles.23 This approach can result in high selectivity of binding, 
internalization, and effective retention of the micelles in the tumor cells. In addition, 
recent advances in antibody engineering allow for the production of humanized 
antibody fragments, reducing problems with immune response against mouse 
antibodies.140 For example, micelles of PEO-distearyl phosphatidylethanolamine 
were covalently modified with the monoclonal antibody 2C5 that binds to microsomes, 
displayed at the surface of many tumor cells. The micelles were then 
used for incorporating various poorly soluble anticancer drugs including tamoxifen, 
paclitaxel, dequalinium, and chlorine e6 trimethyl ester.26'132'141 It was shown 
that paclitaxel-loaded 2C5-immunomicelles could specifically recognize a variety 
of tumor types. The binding of these immunomicelles was observed for all 
cancer cell lines tested, i.e. murine Lewis lung carcinoma, T-lymphoma EL4, and 
human breast adenocarcinomas, BT-20 and MCF7.141 Moreover, paclitaxel-loaded 
2C5 immunomicelles demonstrated highest anticancer activity in Lewis lung carcinoma 
tumor model in mice, compared with plain paclitaxel-loaded micelles and 
74 Batrakova et al. 
the free drug.132 The increased antitumor effect of immunomicelles in vivo correlated 
with the enhanced retention of the drug delivered with the immunomicelles 
inside the tumor. 
Tumors often display low pH of interstitial fluid, which is mainly attributed 
to higher rates of aerobic and anaerobic glycolysis in cancer cells than in normal 
cells.142,143 This phenomenon has been employed in the design of various 
pH-sensitive polymer micelle systems for the delivery of anticancer drugs to the 
tumors. One approach consisted in the chemical conjugation of anticancer drugs 
to the block copolymers through pH sensitive cleavable links that are stable at 
neutral pH, but are cleavable and release the drug in the mildly acidic pH. For 
example, several groups used hydrasone-based linking groups, to covalently attach 
doxorubicin to PEO-b-poly(DL-lactic-co-glycolic acid) block copolymer,21,144 PEOb-
block-poly(allyl glycidyl ether)145 or PEO-b-poly(aspartate hydrazone) block 
copolymer.63,64 It was suggested that doxorubicin will remain in the micelles in 
the blood stream, and will be released at tumor sites at lower pH. For example, 
in vitro and in vivo studies using PEO-b-poly(aspartate hydrazone doxorubicin) 
micelles demonstrated that the micelles display an intracellular pH-triggered drug 
release capability, tumor-infiltrating permeability, and effective antitumor activity 
with extremely low toxicity.63,64 Overall, the animal studies suggested that such 
polymer micelle drug has a wide therapeutic window due to increased efficacy 
and decreased toxicity, compared with free doxorubicin.64 
An alternative mechanism for pH-induced triggering of drug release at the 
tumor sites consists of using pH sensitive polyacids or polybases as building 
blocks for polymer micelles.94,146,147 For example, mixed micelles of PEO-bpoly(
L-histidine) and PEO-b-poly(L-lactic acid) block copolymers incorporate pHsensitive 
poly-base, poly(L-histidine) in the hydrophobic core.147 The core can also 
solubilize hydrophobic drugs such as doxorubicin. The protonation of the polybase 
at acidic conditions resulted in the destabilization of the core and triggered 
release of the drug. This system was also targeted to the tumors through the folate 
molecules as described earlier and has shown significant in vivo antitumor activity 
and less side effects, compared with the free drug.129 Notably, it was also effective 
in vitro and in vivo against multidrug resistant (MDR) human breast carcinoma 
MCF7/ADR that overexpresses P-glycoprotein (Pgp). Pgp is a drug efflux transport 
protein that serves to eliminate drugs from the cancer cells and significantly 
decreases the anticancer activity of the drugs. The micelle incorporated drug was 
released inside the cells, and thus avoided the contact with Pgp localized at the 
cell plasma membrane, which perhaps contributed to the increased activity of pH 
sensitive doxorubicin micelles in the MDR cells. 
A different approach using Pluronic block copolymer micelles to overcome 
MDR in tumors has been developed by our group.130,148-151 Studies by Alakhov 
Polymer Micelles as Drug Carriers 75 
et al. demonstrated that Pluronic block copolymers can sensitize MDR cells, 
resulting in an increased cytotoxic activity of doxorubicin, paclitaxel, and other 
drugs by 2,3 orders of magnitude.148'149 Remarkably, Pluronic can enhance drug 
effects in MDR cells through multiple effects including (1) inhibiting drug efflux 
transporters, such as Pgp149-152 and multidrug resistance proteins (MRPs),153'154 
(2) abolishing drug sequestration within cytoplasmic vesicles,149'153 (3) inhibiting 
the glutathione/glutathione S-transferase detoxification system,154 and (4) enhancing 
proapoptotic signaling in MDR cells.155 Similar effects of Pluronics have also 
been reported using in vivo tumor models.130,150 In these studies, mice bearing 
drug-sensitive and drug-resistant tumors were treated with doxorubicin alone 
and with doxorubicin in Pluronic compositions. The tumor panel included i.p. 
murine leukemias (P388, P388-Dox), s.c. murine myelomas (Sp2/0, Sp2/0-Dnr), 
i.v. and s.c. Lewis lung carcinoma (3LL-M27), s.c. human breast carcinomas (MCF7, 
MCF7/ADR), and s.c. human oral epidermoid carcinoma (KBv).130 Using the NCI 
criteria for tumor inhibition and increased lifespan, Pluronic/doxorubicin has met 
the efficiency criteria in all models (9 of 9), while doxorubicin alone was only effective 
in selected tumors (2 of 9) .130 Results showed that the tumors were more responsive 
in the Pluronic /doxorubicin treatment groups than in doxorubicin alone. These 
studies demonstrated improved treatment of drug resistant cancers with Pluronics. 
The mechanisms of effects of Pluronic on Pgp have been studied in great 
detail.151 In particular, exposure of MDR cells to Pluronics has resulted in the 
inhibition of Pgp-mediated efflux,149 and this overcomes defects in intracellular 
accumulation of Pgp-dependent drugs,148,149,152 and abolishes the directionality 
difference in the flux of these drugs across polarized cell monolayers.156-158 
The lack of changes in membrane permeability with Pluronics to (1) non-Pgp 
compounds in MDR cells,158,159 and (2) Pgp probes in non-MDR cells,149,153 suggested 
that Pluronic effects were specific to the Pgp efflux system. These effects 
were observed at Pluronic concentrations less than or equal to the critical micelle 
concentration (CMC).152,159 Thus, Pluronic unimers rather than the micelles were 
responsible for these effects. Specifically, Pluronic molecules displayed a dual function 
in MDR cells.160-162 Firstly, they incorporated into the cell membranes and 
decreased the membrane microviscosity. This was accompanied by the inhibition 
of Pgp ATPase activity. Secondly, they translocated into cells and reached intracellular 
compartments. This was accompanied by the inhibition of respiration,163 
presumably due to Pluronic interactions with the mitochondria membranes. As a 
result, within 15 min after exposure to select Pluronics, intracellular levels of ATP in 
MDR cells were drastically decreased.160-162 Remarkably, such ATP depletion was 
not observed in non-MDR cells, suggesting that the Pluronic was "selective", with 
respect to the MDR phenotype.160'164 Combining these two effects, Pgp ATPase inhibition 
and ATP depletion, resulted in the shut-down of the efflux system in MDR 
76 Batrakova et al. 
cells.160-162 The Pgp remained functionally active when (1) ATP was restored using 
an ATP supplementation system in the presence of a Pluronic, or (2) when ATP was 
depleted, but there was no direct contact between the Pluronic and Pgp (and no 
ATPase inhibition). Overall, these detailed studies which resulted in the development 
of a micellar formulation of doxorubicin that is evaluated clinically, reinforce 
the fact that block copolymers, comprising the micelles, can serve as biological 
response modifying agents that can have beneficial effects in the chemotherapy of 
5.2. Drug delivery to the brain 
By restricting drug transport to the brain, the blood brain barrier (BBB) represents a 
formidable impediment for the treatment of brain tumors and neurodegenerative 
diseases such as HIV-associated dementia, stroke, Parkinson's and Alzheimer's 
diseases. Two strategies using polymer micelles have been evaluated to enhance 
delivery of biologically active agents to the brain. The first strategy is based on 
the modification of polymer micelles with antibodies or ligand molecules capable 
of transcytosis across brain microvessel endothelial cells, comprising the BBB. The 
second strategy uses Pluronic block copolymers to inhibit drug efflux systems, 
particularly, Pgp, and selectively increase the permeability of BBB to Pgp substrates. 
An earlier study used micelles of Pluronic block copolymers for the delivery 
of the CNS drugs to the brain.68'73 These micelles were surface-modified by attaching 
to the free PEO ends, either polyclonal antibodies against brain-specific antigen, 
a2-glycoprotein, or insulin to target the receptor at the lumenal side of BBB. 
The modified micelles were used to solubilize fluorescent dye or neuroleptic drug, 
haloperidol, and these formulations were administered intravenously in mice. Both 
the antibody and insulin modification of the micelles resulted in enhanced delivery 
of the fluorescent dye to the brain and drastic increases in neuroleptic effect of 
haloperidol in the animals. Subsequent studies using in vitro BBB models demonstrated 
that the micelles, vectorized by insulin, undergo receptor-mediated transport 
across brain microvessel endothelial cells.156 Based on one of these observations, 
one should expect development of novel polymer micelles that target specific 
receptors at the surface of the BBB to enhance transport of the incorporated drugs 
to the brain. 
The studies by our group have also demonstrated that selected Pluronic block 
copolymers, such as Pluronic P85, are potent inhibitors of Pgp, and they have the 
increased entry of the Pgp-substrates to the brain across BBB.156'158'159'165 Pluronic 
did not induce toxic effect in BBB, as revealed by the lack of alteration in paracellular 
permeability of the barrier,156'158 and in histological studies, using specific markers 
for brain endothelial cells.166 Overall, this strategy has potential in developing 
Polymer Micelles as Drug Carriers 77 
novel modalities for the delivery of various drugs to the brain, including selective 
anti cancer agents to treat metastatic brain tumors, as well as HIV protease 
inhibitors to eradicate HIV virus in the brain.167'168 
5.3. Formulations of antifungal agents 
The need for safe and effective modalities for the delivery of chemotherapeutic 
agents to treat systemic fungal infections in immunocompromised AIDS, surgery, 
transplant and cancer patients is very high. The challenges to the delivery of antifungal 
agents include low solubility and sometimes high toxicity of these agents. 
These agents, such as amphotericin B, have low compatibility with hydrophobic 
cores of polymer micelles formed by many conventional block copolymers. Thus, to 
increase solubilization of amphotericin B, the core-forming blocks of methoxy-PEOb-
poly(L-aspartate) were derivatized with stearate side chains.169-172 The resulting 
block copolymers formed micelles. Amphotericin B interacted strongly with the 
stearate side chains in the core of the micelles, resulting in an efficient entrapment 
of the drug in the micelles, as well as subsequent sustained release in the external 
environment. As a result of solubilization of amphotericin B in the micelles, the 
onset of hemolytic activity of this drug toward bovine erythrocytes was delayed, 
relative to that of the free drug.171 Using a neutropenic murine model of disseminated 
Candidas, it was shown that micelle-incorporated amphotericin B retained 
potent in vivo activity. Pluronic block copolymers were used by the same group 
for incapsulation of another poorly soluble antifungal agent, nystatin.172 This is a 
commercially available drug that has shown potential for systemic administration, 
but has never been approved for that purpose, due to toxicity issues. The possibility 
to use Pluronic block copolymers to overcome resistance to certain antifungal 
agents has also been demonstrated.173-176 Overall, one should expect further scientific 
developments using polymer micelle delivery systems for the treatment of 
fungal infection. 
5.4. Delivery of imaging agents 
Efficient delivery of imaging agents to the site of disease in the body can improve 
early diagnostics of cancer and other diseases. The studies in this area using polymer 
micelles as carriers for imaging agents were initiated by Torchilin.177 For example, 
micelles of amphiphilic PEO-lipid conjugates were loaded with i n In and 
gadolinium diethylenetriamine pentaacetic acid-phosphatidylethanolamine (Gd- 
DTPA-PE) and then used for visualization of local lymphatic chain after subcutaneous 
injection into the rabbit's paw.178 The images of local lymphatics were 
acquired using a gamma camera and a magnetic resonance (MR) imager. The 
78 Batrakova et al. 
injected micelles stayed within the lymph fluid, thus serving as lymphangiographic 
agents for indirect MR or gamma lymphography. Another polymer micelle system 
composed of amphiphilic methoxy-PEO-b-poly[epsilon,N-(triiodobenzoyl)-Llysine] 
block copolymers, labeled with iodine, was administered systemically in 
rabbits and visualized by X-ray computed tomography.179 The labeled micelles 
displayed exceptional 24 hrs half-life in the blood, which is likely due to the coreshell 
architecture of the micelle carriers that protected the iodine-containing core. 
Notably, small polymer micelles (<20nm) may be advantageous for bioimaging 
of tumors, compared with PEG-modified long-circulating liposomes (ca. lOOnm). 
In particular, the micelles from PEO-distearoyl phosphatidyl ethanolamine conjugates 
containing m In-labeled model protein were more efficacious in the delivery of 
protein to Lewis lung carcinoma than larger long-circulating liposomes.180 Overall, 
polymer micelles loaded with various agents for gamma, magnetic resonance, and 
computed tomography imaging represent promising modalities for non-invasive 
diagnostics of various diseases. 
5.5. Delivery of polynucleotides 
To improve the stability of polycation-based DNA, delivery complexes in dispersion 
block and graft copolymers containing segments from polycations and nonionic 
water-soluble polymers, such as PEO, were developed.90,181,182 Binding of 
these copolymers with DNA results in the formation of micelle-like block ionomer 
complexes ("polyion complex micelles"), containing hydrophobic sites formed by 
the polycation-neutralized DNA and hydrophilic sites formed by the PEO chains. 
Despite neutralization of charge, complexes remain stable in aqueous dispersion 
due to the effect of the PEO chains.183 Overall, the PEO modified polycation-DNA 
complexes form stable dispersions and do not interact with serum proteins.183,184 
These systems were successfully used for intravitreal delivery of an antisense 
oligonucleotide and the suppression of gene expression in retina in rats.185 Furthermore, 
they displayed extended plasma clearance kinetics and were shown to 
transfect liver and tumor cells, after systemic administration in the body.186-188 In 
addition, there is a possibility targeting such polyplexes to the specific receptors at 
the surface of the cell, for example, by modifying the free ends of PEO chains with 
specific targeting ligands.189-191 Alternatively, to increase the binding of the complexes 
with the cell membrane and the transport of the polynucleotides inside the 
cells, the polycations were modified with amphiphilic Pluronic molecules.192,193 One 
recent study has shown a potential of Pluronic-polyethyleneimine-based micelles 
for in vivo delivery of antisense oligonucleotides to tumors, and have demonstrated 
sensitization of the tumors to radiotherapy as a result of systemic administration 
of the oligonucleotide-loaded micelles.194 
Polymer Micelles as Drug Carriers 79 
6. Clinical Trials 
Three polymer micelle formulations of anticancer drugs have been reported to 
reach clinical trials. The doxorubicin-conjugated polymer micelles developed by 
Kataoka's group195 have progressed recently to Phase I clinical trial at the National 
Cancer Center Hospital, Tokyo, Japan. The micelle carrier NK911 is based on PEO-bpoly(
aspartic acid) block copolymers, in which the aspartic acid units were partially 
(ca. 45%) substituted with doxorubicin to form hydrophobic block. The resulting 
substituted block copolymer forms micelles that are further noncovalently loaded 
with free doxorubicin. Preclinical studies in mice demonstrated higher NK911 activity 
against Colon 26, M5076, and P388, compared with the free drug. Moreover, 
NK911 has less side effects, resulting in less animal body and toxic death than the 
free drug.196 
Clinically, the Pluronic micelle formulation of doxorubicin has been most 
advanced. Based on the in vivo efficacy evaluation, Pluronic L61 was selected for 
clinical development for the treatment of MDR cancers. The final block copolymer 
formulation is a mixture of 0.25% Pluronic L61 and 2% Pluronic F127, formulated 
in isotonic buffered saline.130 This system contains mixed micelles of L61 and F127, 
with an effective diameter of ca. 22 to 27 nm and is stable in the serum. Prior to 
administration, doxorubicin is mixed with this system, which results in spontaneous 
incorporation of the drug in the micelles. The drug is easily released by diffusion 
after dilution of the micelles. The formulation of doxorubicin with Pluronic, 
SP1049C, is safe, following systemic administration based on toxicity studies in 
animals.130 A two-site Phase I clinical trial of SP1049C has been completed.197 Based 
on its results, the dose-limiting toxicity of SP1049C was myelosuppression, reached 
at 90mg/m2 (maximum tolerated dose was 70mg/m2). Phase II study of this formulation 
to treat inoperable metastatic adenocarcinoma of the esophagus is near 
completion as well.198 
Finally, Phase I studies were reported for Genexol-PM, a Cremophor-free polymer 
micelle-formulated paclitaxel.199 Twenty-one patient entered into this study 
with lung, colorectal, breast, ovary, and esophagus cancers. No hypersensitivity 
reaction was observed in any patient. Neuropathy and myalgia were the most 
common toxicities. There were 14% partial responses. The paclitaxel area under 
the curve and peak of the drug concentration in the blood were increased with the 
escalating dose, suggesting linear pharmacokinetics for Genexol-PM.199 
7. Conclusions 
Approximately two decades have passed since the conception of the polymer 
micelle conjugates and nanocontainers for drug delivery. During the first decade, 
80 Batrakova et al. 
only a few studies were published; however, more recently, the number of publications 
in this field has increased tremendously. During this period, novel biocompatible 
and/or biodegradable block copolymer chemistries have been researched, 
the block ionomer complexes capable of incorporating DNA and other charged 
molecules have been discovered, the pH and other chemical signal sensitive micelles 
have been developed. Many studies focused on the use of polymer micelles for 
delivery of poorly soluble and toxic chemotherapeutic agents to the tumors to 
treat cancer. There has been considerable advancement in understanding the processes 
of polymer micelle delivery into the tumors, including passive and vectorized 
targeting of the polymer micelles. Notable achievements also include the studies 
demonstrating the possibilities for overcoming multidrug resistance in cancer, and 
enhancing drug delivery to the brain using block copolymer micelles systems. Overall, 
it is clear that this area has reached a mature stage, reinforced by the fact that 
several human clinical trials using polymer micelles for cancer drug delivery have 
been initiated. At the same time, it is obvious that the possibilities for delivery of 
the diagnostic and therapeutic agents using polymer micelles are extremely broad, 
and one should expect further increase in the laboratory and clinical research in 
this field during the next decade. Targeting polymer micelles to cancer sites within 
the body will address an urgent need to greatly improve the early diagnosis and 
treatment of cancer. Capabilities for the discovery and use of targeting molecules 
will support the development of multifunctional therapeutics that can carry and 
retain antineoplastic agents within tumors. This will also be instrumental in developing 
novel biosensing and imaging modalities for the early detection of cancer 
and other devastating human diseases. 
The authors acknowledge the support of the research using polymer micelles by 
grants from the National Institutes of Health CA89225, NS36229 and EB000551, 
as well as the National Science Foundation DMR0071682, DMR0513699 and BES- 
9907281. We also acknowledge financial support of Supratek Pharma, Inc. (Montreal, 
Canada). AVK and EVB are shareholders and AVK serves as a consultant to 
this Company. 
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172. Croy SR and Kwon GS (2004) The effects of Pluronic block copolymers on the aggregation 
state of nystatin. / Control Rel 95:161-171. 
173. Jagannath C, Sepulveda E, Actor JK, Luxem F, Emanuele MR and Hunter RL 
(2000) Effect of poloxamer CRL-1072 on drug uptake and nitric-oxide-mediated 
killing of Mycobacterium avium by macrophages. Immunopharmacology 48: 
92 Batrakova et al. 
174. Jagannath C, Emanuele MR and Hunter RL (2000) Activity of poloxamer CRL- 
1072 against drug-sensitive and resistant strains of Mycobacterium tuberculosis in 
macrophages and in mice. Int J Antimicrob Agents 15:55-63. 
175. Jagannath C, Emanuele MR and Hunter RL (1999) Activities of poloxamer CRL-1072 
against Mycobacterium avium in macrophage culture and in mice. Antimicrob Agents 
Chemother 43:2898-2903. 
176. Jagannath C, Wells A, Mshvildadze M, Olsen M, Sepulveda E, Emanuele M, Hunter 
RL, Jr. and Dasgupta A (1999) Significantly improved oral uptake of amikacin in FVB 
mice in the presence of CRL-1605 copolymer. Life Sci 64:1733-1738. 
177. Torchilin VP (2002) PEG-based micelles as carriers of contrast agents for different imaging 
modalities. Adv Drug Del Rev 54:235-252. 
178. Trubetskoy VS, Frank-Kamenetsky MD, Whiteman KR, Wolf GL and Torchilin VP (1996) 
Stable polymeric micelles: Lymphangiographic contrast media for gamma scintigraphy 
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179. Trubetskoy VS, Gazelle GS, Wolf GL and Torchilin VP (1997) Block-copolymer of 
polyethylene glycol and polylysine as a carrier of organic iodine: Design of longcirculating 
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180. Weissig V, Whiteman KR and Torchilin VP (1998) Accumulation of protein-loaded longcirculating 
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181. Katayose S and Kataoka K (1997) Water-soluble polyion complex associates of DNA 
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Characterization of vectors for gene therapy formed by self-assembly of DNA with 
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183. Vinogradov SV, Bronich TK and Kabanov AV (1998) Self-assembly of polyaminepoly(
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184. Itaka K, Harada A, Nakamura K, Kawaguchi H and Kataoka K (2002) Evaluation by 
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185. Roy S, Zhang K, Roth T, Vinogradov S, Kao RS and Kabanov A (1999) Reduction of 
fibronectin expression by intravitreal administration of antisense oligonucleotides. Nat 
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186. Ogris M, Steinlein P, Kursa M, Mechtler K, Kircheis R and Wagner E (1998) The size of 
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cells. Gene Ther 5:1425-1433. 
187. Oupicky D, Ogris M, Howard KA, Dash PR, Ulbrich K and Seymour LW (2002) Importance 
of lateral and steric stabilization of polyelectrolyte gene delivery vectors for 
extended systemic circulation. Mol Ther 5:463-472. 
Polymer Micelles as Drug Carriers 93 
188. Harada-Shiba M, Yamauchi K, Harada A, Takamisawa I, Shimokado K and Kataoka K 
(2002) Polyion complex micelles as vectors in gene therapy-pharmacokinetics and 
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protein-modified corona for receptor-mediated delivery of oligonucleotides into cells. 
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This page is intentionally left blank
Vesicles Prepared from Synthetic 
Amphiphiles  Polymeric Vesicles and 
Ijeoma Florence Uchegbu and Andreas G. Schatzlein 
1. Introduction 
This chapter will examine what is known about vesicles prepared from synthetic 
amphiphiles and will encompass a review of the data published on polymeric vesicles 
and non-ionic surfactant vesicles (niosomes). Schematic representations of the 
molecular arrangements in these systems are as depicted in Fig. 1. Examples of 
drug delivery applications will also be presented. 
Vesicular systems arise when amphiphilic molecules self assemble in aqueous 
media in an effort to reduce the high energy interaction between the hydrophobic 
portion of the amphiphile and the aqueous disperse phase, and maximize the low 
energy interaction between the hydrophilic head group and the disperse phase 
(Fig. 1). These self assemblies reside in the nanometre to micrometre size domain. 
Excellent reviews exist on the self assembly of amphiphiles/16 and hence this topic 
will not be dealt with in great detail here. Vesicles are important pharmaceutical 
systems, especially as liposomes, the result of phospholipid self assembly,19 are 
licensed for the clinical delivery of anti cancer drugs.21 It is thus possible that the 
vesicles described here may be incorporated into licensed medicines at some point 
in future. 
96 Uchegbu & Schatzlein 
t % 
If 111 
Self assembling 
polymerisable monomers 
(a) Polymerised 
isfi? M# 
*   * ^ 
is m 
(b) Self assembling 
amphiphilic polymers 
80% favors dense 
nanoparticles, while a polydactic acid) fraction of 58-80% favors bilayer vesicle 
assemblies, and a polydactic acid) fraction of less than 50% favors the production 
of micellar self assemblies.31 
The sizes of the vesicle and dense nanoparticle assemblies formed from 
amphiphilic poly(ethylenimines) are also dependent on polymer levels of 
hydrophobic modification (mole % cetylation) and the relationships shown in 
Eqs. (1) and (2) have been developed,18 
dv = 1.95Ct + 139 (1) 
dn = 2.31Ct + 5.6 (2) 
where dv = vesicle z-average mean hydrodynamic diameter, Ct = mole% cetylation 
(number of cetyl groups per 100 monomer units), and dn  nanoparticle 
z-average mean hydrodynamic diameter. 
The molecular weight of the polymer is also an important factor to consider 
when choosing vesicle forming polymers. The importance of this parameter has 
been demonstrated with the poly(L-lysine) vesicle system20 [e.g. Compound 6, 
Fig. 3(a)]. With these amphiphiles a vesicle formation index (F') has been computed: 
F' = - ^ = (3) 
where H = mole% unreacted L-lysine units, L  mole% L-lysine units substituted 
with palmitic acid and DP = the degree of polymerisation of the polymer. An F' 
value in excess of 0.168 is necessary for vesicle formation.20 
Additionally, not only does the molecular weight of the polymer impact on 
vesicle formation, but it is also a direct controller of the vesicle mean size; the 
relationship shown in Eq. (4) has been developed for the palmitoyl glycol chitosan 
VMW = 0.782dv + 107 (4) 
Vesicles Prepared from Synthetic Amphiphiles 101 
where MW = polymer molecular weight, and dv = vesicle z-average mean 
hydrodynamic diameter. 
2.3. Block copolymers 
Block copolymer vesicles, termed "polymersomes" are fairly new discoveries, being 
first reported in the 1990s.32 Polymersomes have been prepared from a variety of 
block copolymers, some examples of which are given in Fig. 4. There is a clear 
relationship between the hydrophobic content of polymers and self assembly. Low 
levels of hydrophobicity (less than 50% of the polymer consisting of hydrophobic 
HO, X N^- 
Fig. 4. Examples of some vesicle forming block copolymers Compound 7,1 Compound 8,7 
and Compound 9.13 
102 Uchegbu & Schatzlein 
moieties) favors the formation of micelles33 and intermediate levels of hydrophobicity 
(50-80%) favors the formation of bilayer vesicles.31,33'34 For the self assembly 
of block copolymers, it has been established that generally the critical packing 
parameter (CPP): 
CPP = ^ (5) 
should approach unity for vesicular self assemblies to prevail,24 where v = volume 
of the hydrophobic block, 1  length of the hydrophobic block and a = the area of 
the hydrophilic block. 
Vesicle sizes are varied and range from tens of nanometres35 to tens of 
microns.36 Polymersome membranes are 8-21 nm thick; 2-5 times thicker than 
the 4nm membrane thickness displayed by conventional low molecular weight 
amphiphiles.16'27,31'34,35 The thickness of the membrane is determined by the degree 
of polymerization in the hydrophobic block34 and these extra thick membranes 
confer, on the vesicle, exceptional stability to soluble surfactantS24 and mechanical 
stress.24'27,37'38 With these vesicles, there is an asymmetric distribution of the 
polymers in the inner and outer leaflets of the bilayer and polymers with a large 
hydrophilic chain length are preferentially localized to the exterior leaflet and vice 
versa.39 Preferred residence in the outer leaflet is favored by the more hydrophilic 
polymers, because the greater repulsion between the longer hydrophilic corona 
molecules on the outer leaflet stabilize the vesicle curvature.39 
Vesicle stability is a desirable characteristic for pharmaceutical vesicles and as 
such, a great deal of effort has been expended on producing stable systems. As the 
drive for nanomedicines (medicines incorporating functional nanoparticles) grows, 
stability issues will need to be adequately addressed to ensure the widespread 
adoption of such systems. In actual fact, the early workers in the polymeric vesicle 
field were primarily driven by this need to produce stable drug carriers. Extremely 
stable systems are possible on polymerization of block copolymers subsequent to 
self assembly. Poly(ethylene oxide)-WocA:-poly[3-(trimethoxysilyl)propyl methacrylate] 
copolymer vesicles in water, methanol, triethylamine mixtures produced 
polymerized polymersomes that are stable for up to one year.40 Triethylamine 
hydrolyzes the trimethoxysilyl groups and then catalyzes their polycondensation 
to yield an extremely stable hydrophobic polysilsesquioxane core.40,41 Additionally, 
poly(ethylene oxide)-Wocfc-poly(butadiene) vesicles on cross linking produce 
vesicles which are organic solvent resistant.42 
2.4. Preparing vesicles from self-assembling polymers 
Polymeric vesicles are relatively simple to prepare. The input of energy is achieved 
in the laboratory by probe sonication of the amphiphilic polymer in the disperse 
Vesicles Prepared from Synthetic Amphiphiles 103 
phase.1120 However, clearly the energy required for self assembly is not trivial as 
vesicles are not easily formed by hand shaking, unlike low molecular weight surfactant 
formulations.4 Vesicles once formed are morphologically stable for months11 
and may be loaded with hydrophilic43-45 and hydrophobic [see Fig. 6(b) below] 
solutes, by probe sonicating in the presence of such solutes. Commercially, it is 
envisaged that polymeric vesicles may be fabricated by microfluidization and high 
pressure homogenization techniques. 
2.5. Self assembling polymerizable monomers 
Polymerized vesicles may also be prepared by utilizing self assembling polymerizable 
amphiphiles, followed by the polymerization of the resulting vesicular self 
assembly (Fig. 1). Examples of some polymerizable vesicle forming monomers are 
shown in Fig. 5. This method of producing polymerised vesicles is the oldest form 
of polymeric vesicle technology.12,46 
l O 
Fig. 5. Polymerizable vesicle forming monomers used to make polymerized vesicles 
by Jung and others (Compound 10),5 Cho and others (Compound ll),8 Hub and others 
(Compound 12)12 and Bader and others (Compound 13).15 
104 Uchegbu & Schatzlein 
Polymerized vesicles prepared using polymerized self assembling monomers 
are essentially polymer shells and it is unclear how much of the bilayer assembly 
actually survives the polymerization step. The advantage, however, is that 
they are extremely stable, resisting degradation by detergents47-49 or organic 
solvents.8'48,50,51 They are also less leaky,50 thermostable,52 and because the vesicle 
forming components are kinetically trapped by the polymerization process, they 
have improved colloidal stability.8 A major advantage of these nanosystems is that 
they may be isolated as dry powders which are readily dispersible in water to give 
50-100 nm particles;48 thus potentially allowing the formulation of solid vesicle 
dosage forms. Polymerization involves fairly reactive species and hence vesicles 
are best prepared prior to drug loading, which may be a limitation. 
3. Polymeric Vesicle Drug Delivery Applications 
Polymeric vesicles, which are the focus of this chapter, exist in two main varieties as 
illustrated in Fig. 1. These technologies are suitable candidates for the development 
of robust, controllable and responsive nanomedicine drug carriers. 
3.1. Drug targeting 
Poly(oxyethylene) amphiphiles, when incorporated into liposomal26 and niosomal6 
bilayers, prolong vesicle circulation and facilitate tumor targeting,6'53 due to the 
leaky nature of the poorly developed tumor vascular endothelium.54 Only 10 
mole % poly(ethylene oxide)  lipid amphiphiles may be incorporated into 
liposomes55 or niosomes,56'57 without a loss of vesicle integrity due to the preferred 
tendency of the hydrophilic poly(oxyethylene) amphiphiles to form micelles. Polymersomes 
composed of poly(ethylene oxide)-Wocfc-polybutadiene or poly(ethylene 
oxide)-Wocfc-poly(ethylethylene), in which the entire vesicle surface is covered with 
the poly(ethylene oxide) coat, have been studied as long circulating nanocarriers 
for drug delivery58 The circulation time of poly(ethylene oxide) polymersomes is 
directly dependent on the length of the poly(ethylene oxide) block and polymersome 
half lives of up to 28 hrs have been recorded in rats with a poly(ethylene oxide) 
degree of polymerization of 50.58 This half life compares favorably with a half life 
of 14 hrs recorded for poly(oxyethylene) coated liposomes.59 It is assumed that 
the 100% surface coverage of the polymeric vesicles is responsible for the reduced 
clearance of these polymersomes from the blood.38 The long half life of these polymersomes 
makes them excellent candidates for the development of anti tumor 
Furthermore, drug release may be controlled in the polymersomes by controlling 
the hydrolysis rate of the hydrophobic blocks.31 This has been demonstrated 
Vesicles Prepared from Synthetic Amphiphiles 105 
with poly(L-lactic acid)-fr/ocfc-poly(ethylene glycol) and poly(caprolactone)-Wod> 
poly(ethylene glycol) vesicles.31 Hydrolysis of the hydrophobic block causes the 
polymer to move from a vesicular to a micellar assembly, as the overall level of 
hydrophobic content diminishes, and this in turn leads to drug release.31 Hydrolysis 
rates and implicitly release rates may be controlled by varying the relative level 
of the hydrophobic blocks. 
Carbohydrate polymeric vesicles may also be used as drug targeting agents. 
Vesicles prepared from glycol chitosan vesicles improve the intracellular delivery 
of hydrophilic macromolecules44 and anti cancer drugs,45 the latter is achieved with 
the help of a transferrin ligand attached to the surface of the vesicle. 
3.2. Gene delivery 
Poly(L-lysine) based vesicles, prepared from Compound 6 [Fig. 3(a)] have been used 
for gene delivery,29,60 as these vesicles are less toxic than unmodified poly(L-lysine) 
and produce higher levels of gene transfer (Table l).29 The production of polymeric 
vesicles and the resultant reduction in cytotoxicity enables poly(L-lysine) to be used 
in in vivo gene, as the unmodified polymer is too toxic for in vivo use. When the 
targeting ligand, galactose, was bound to the distal ends of the poly(oxyethylene) 
chains, gene expression was increased in HepG2 cells in vitro.60 However, in vivo 
targeting to the liver hepatocytes was not achieved with these systems.60 
A similar procedure with the poly(ethylenimine) vesicles prepared using 
Compound 5 [Fig. 3(a)] also resulted in a reduction in the cytotoxicity of the polymer 
(Table l),17 although in this case, the poly(ethylenimine) vesicles were not as 
efficient gene transfer agents as the free polymer. 
Table 1 Biological Activity of poly(ethylenimine)17 and poly(L-lysine)29 Vesicles. 
Polymer A431 cells A549 
IC50 Gene Transfer IC50 Gene Transfer 
(AtgmL-1) Relative to Parent (jiigmL-1) Relative to Parent 
Polymer Polymer 
Poly(ethylenimine) 1.9 1 5.2 1 
Polymer 5 (Fig. 6(a)) 16.9 0.2 12.6 0.08 
Polymer 5, cholesterol 15.9 0.2 11 0.08 
vesicles 2:1 (gg_1) 
Poly(L-lysine) 7 1 7 1 
Polymer 6 (Fig. 6(a)) 74 7.8 63 2.3 
106 Uchegbu & Schatzlein 
3.3. Responsive release 
The ultimate goal of all drug delivery efforts is the simple fabrication of responsive 
systems that are capable of delivering precise quantities of their pay load in response 
to physiological or more commonly pathological stimuli. Pre-programmable pills, 
implants and injectables are so far merely the unobtainable ideal, however, polymeric 
systems have been fabricated with responsive capability and it is possible 
that in the future, these may be fine tuned to produce truly intelligent and dynamic 
drug delivery devices or systems. 
The various environmental stimuli that may be used to trigger the release of 
encapsulated drug are outlined below and examples are given of existing developments 
in the area. However, in addition to the areas covered below, it may 
be possible in future for pathology specific molecules to interact with polymeric 
vesicles to trigger release. 
3.3.1. pH 
Diblock polypeptides, in which the hydrophilic block consists of ethylene glycol 
derivatised amino acids (L-lysine), and the hydrophobic block consists of poly 
(L-leucine), form pH responsive vesicles which disaggregate at low pH, providing 
the level of L-leucine and polymer chain length is maintained within defined limits 
of about 12-25 mole% and the polymer has a degree of polymerization of less than 
200.13 These L-lysine based systems may be applied to facilitate endosome specific 
3.3.2. Enzymatic 
Vesicles which release their contents in the presence of an enzyme may be formed 
by loading polymeric vesicles with an enzyme activated prodrug (Fig. 6). The 
particulate nature of the drug delivery system should allow the drug to accumulate 
in tumors, for example, where it may then be activated by an externally 
applied enzyme in a similar manner to the antibody directed enzyme prodrug 
therapeutic strategy. The antibody directed enzyme prodrug therapeutic strategy 
enables an enzyme to be homed to tumors using antibodies followed by the 
application of an enzyme activated prodrug.61 Alternatively, a membrane bound 
enzyme may be used to control and ultimately prolong the activity of either an 
entrapped hydrophilic drug (entrapped in the vesicle aqueous core) or an entrapped 
hydrophobic drug (entrapped in the vesicle membrane) as illustrated in Fig. 6. It is 
possible that the enzyme may be chosen such that it is activated in the presence of 
pathology specific molecules, thus achieving pathology responsive and localized 
drug activity. 
Vesicles Prepared from Synthetic Amphiphiles 107 
<2 0.5- 
I 1 
-O vesicle bound enzyme + external substrate 
-  external enzyme + vesicle loaded substrate 
-A control solution + substrate 
--? ^""V 
0 20 40 60 80 A 
 Enzy m e Time(min) ^ W 
A Water soluble Substrate 
Fig. 6(a). Enzyme activated polymeric vesicles. Vesicles bearing membrane bound enzyme 
(i) were formed by probe sonicating Compound 2 [Fig. 3(a)], cholesterol, N-biotinylated 
dipalmitoyl phosphatidyl ethanolamine (8: 4: 1 gg"1) in neutral phosphate buffer (2mL), 
isolation of the vesicles by ultracentrifugation (150,000 g), redispersion in a similar volume 
of neutral phosphate buffer and incubation of the vesicles with /S-galactosidase streptavidin 
(3 U). Membrane bound enzyme (0.2 mL) was then incubated with o-nitrophenyl-/J-Dgalactoside 
(2.1 mM, 2 mL) and the absorbance monitored (X = 410 ran). The control solution 
contained similar levels of substrate (o-nitrophenyl-jS-D-galactoside) but no enzyme. Vesicles 
encapsulating O-nitrophenyl-jS-D-galactoside (ii) were prepared by probe sonicating Compound 
2, cholesterol (8: 4gg_1) in the presence of o-nitrophenyl-jS-D-galactoside solution 
(34 mM, 2 mL) and isolation of the vesicles by ultracentrifugation and redispersion in neutral 
phosphate buffer. These latter vesicles (0.4 mL) were then incubated with /J-D-galactosidase 
(2UmL_1, 0.1 mL) and the absorbance once again monitored. 
3.3.3. Magnetic 
Magnetically responsive polymerized liposomes composed of 1,2-di (2,4- 
octadecadienoyl)-sn-glycerol-3-phosphorylcholine, loaded with ferric oxide and 
subsequently polymerized may be localized by an external magnetic field to the 
small intestine, and specifically the Payer's patches.47 These polymerized vesicles 
are stable to the degradative influence of solubilizing surfactants such as triton-X 
100,47 and hence should not suffer excessive bile salt mediated degradation during 
gut transit. These magnetically responsive polymeric vesicles improve the absorption 
of drugs via the oral route.47 
108 Uchegbu & Schatzlein 
0.0 10 20 30 40 50.0 
^k Membrane bound enzyme 
^ ^ Hydrophobic substrate 
Fig. 6(b). Enzyme activated polymeric vesicles. Vesicles bearing membrane bound 
enzyme and containing the hydrophobic substrate fluorescein di-/S-D-galactospyranoside 
were formed by probe sonicating Compound 2 [Fig. 3(a)], cholesterol, N-biotinylated 
dipalmitoyl phosphatidyl ethanolamine, fluorescein di-^-D-galactospyranoside (8: 4: 1: 
0.0005 g g_1) in neutral phosphate buffer (2 mL) and incubation of the resulting vesicles with 
b-galactosidase streptavidin (0.3 U). The fluorescence of the enzyme hydrolysed substrate 
was then monitored (Excitation wavelength = 490 nm, Emission wavelength = 514 nm). 
3.3.4. Oxygen 
Block copolymer vesicles which are destabilized by oxidative mechanisms have 
been constructed from poly(oxyethylene)-Wocfc-poly(propylene sulphide)-fr/ocfcpoly(
oxyethylene) ABA block copolymers.62 These polymeric vesicles are destabilized 
on the oxidation of the central sulphide block to give sulphoxides and 
ultimately sulphones.62 On oxidation, vesicles are transformed to worm-like 
micelles and finally to spherical micelles, eventually releasing their contents. 
4. Non-ionic Surfactant Vesicles (Niosomes) 
4.1. Self assembly 
The self assembly of non-ionic surfactants into niosomes is dependent on the 
hydrophilic  hydrophobic balance of the surfactant and a CPP (Eq. 1) of between 
0.5-1016 enables niosomal self assembly. Some examples of niosome forming 
molecules are given in Fig. 7. Further molecular specifics that govern niosome 
Vesicles Prepared from Synthetic Amphiphiles 109 
0 /k0J 
- O  v ^ ^ O - 
Fig. 7(a). Examples of some niosome forming surfactants: Compound 14,2 Compound 15,6 
Compound 16,9 and Compound 17.14 
Fig. 7(b). Niosomal membrane additives, Compound 18 = cholesterol, Compound 19 
Solulan C24.4 
110 Uchegbu & Schatzlein 
formation by non-ionic surfactants may be found in published reviews.4'63 Compounds 
such as Compounds 15 (from the sorbitan surfactant class) are established 
pharmaceutical excipients,64 and hence formulation scientists looking to prepare a 
niosome formulation for speedy transition to the clinic will do well looking at this 
class of molecules for exploitable materials. Most niosomes will not only contain the 
non-ionic surfactant, but will also contain other molecules such as the membrane 
stabilizer cholesterol [Fig. 7(b)].4 
The bilayer membrane is an ordered structure which may exist in the gel or 
liquid crystal state. Essentially, molecules are more mobile in the liquid crystalline 
state, enjoying lateral diffusion within the bilayer that is denied them in the gel state. 
For any system, the liquid crystal state exists at a higher temperature (T) than the 
gel state. An increase in temperature favors the transition from the gel to the liquid 
state because of the entropy gain (AS) associated with this transition, ultimately 
leading to a lowering of the free energy (AG) of the system. Cholesterol abolishes 
this membrane phase transition, thus fluidizing the gel state.65 
Niosomes are 30 nm to 120 JJLVSX in size4 and often their surfaces must be stabilized 
against aggregation. Molecules such as the cholesteryl poly(oxyethylene 
ether)  Solulan C246 (Compound 19, Fig. 7b) or the ionic molecule dicetyl 
phosphate66 have been used to confer steric and electrostatic stabilization on these 
vesicles respectively. The reader should be aware that the inclusion of minor 
quantities (<10% by actual weight or molar content) of ionic surfactant does not 
prevent these structures from being discussed in this chapter under the niosome 
heading. Niosomes are often formulated with minor quantities of cationic and other 
It can be said that the formulation of liposomes with poly(ethylene oxide) 
amphiphiles such as distearolyphosphatidylethanolamine-poly(ethylene glycol)26 
was the crucial step that allowed liposomes to become clinically relevant drug 
delivery systems. The resulting liposomes possess a hydrophilic polymer surface, 
which prevents recognition and clearance of the particles from the blood by the liver 
and spleen macrophages,26,67 thus increasing the liposomes' circulation time and 
allowing tumor targeting.68 Niosomes (non-ionic surfactant vesicles), when formulated 
with a water soluble poly(oxyethylene) cholesteryl ether  (Solulan C24), 
also circulate for prolonged periods in the blood, accumulate in the tumor tissue 
and improve tumoricidal activity.6 As well as stabilizing vesicles in the blood, 
poly(oxyethylene) amphiphiles also stabilize vesicles against aggregation, thus 
promoting vesicle colloidal stability.56 
Poly(oxyethylene) amphiphiles, such as Solulan C24, have a large hydrophilic 
head group [Fig. 7(b)], and are thus more hydrophilic than the vesicle forming 
amphiphiles, and hence the level of the former must be kept low to avoid solubilization 
of the membrane and the formation of mixed micelles.57 In actual fact, 
Vesicles Prepared from Synthetic Amphiphiles 111 
unusual morphologies57 result from the incorporation of non-micellizing quantities 
of Solulan C24 in vesicles as discussed below. 
4.2. Polyhedral vesicles and giant vesicles (Discomes) 
A series of unusual morphologies have been isolated from the hexadecyl diglycerol 
ether, Solulan C24, cholesterol phase diagram [Fig. 8(a)]. The addition of Solulan 
C24 to hexadecyl diglycerol ether [Compound 16, Fig. 7(a)] niosomes eventually 
results in the formation of mixed micelles.57 At sub-micellar concentrations of Solulan 
C24 (20-40 mole%), however, giant vesicles (discomes) of 25-100 pm in size are 
formed.57 Discomes are thermoresponsive vesicles, which become more leaky as 
the temperature is increased from room temperature to 37C. These vesicles may 
thus be used to construct thermoresponsive controlled release systems. 
In cholesterol low regions of the hexadecyl diglycerol ether, cholesterol, Solulan 
C24 phase diagram, polyhedral vesicles [Figs. 8(a) and 8(b)] are found.9 These 
polyhedral vesicles are able to entrap water soluble solutes and the membrane, 
which is in the gel state contains areas of high and low curvature as shown in 
 Polyhedral Vesicles (2 -10 jim) 
 Spherical, helical, tubular Vesicles 
(0.5 -10 urn) 
3 Discomes (10 - 30 jim) + small 
spherical & helical vesicles (0.5 -10 
@ Discomes (12-60 fun) + mixed 
\ Reverse Micelles 
V \ 
Fig. 1. A hypothetical pseudo-ternary phase diagram of an oil/surfactant/water system 
with emphasis on microemulsion and emulsion phases. Within the phase diagram, existence 
fields are shown where micelles, reverse micelles or water-in-oil (w/o) microemulsions and 
oil-in-water microemulsions are formed along with the bicontinuous microemulsions. At 
very low surfactant concentrations two phase systems are observed (taken from Ref. 107). 
Recent Advances in Microemulsions as Drug Delivery Vehicles 129 
fractions, microemulsions are generally considered to be a dispersion of either oil 
or water droplets stabilized by an interfacial film of surfactant and where appropriate, 
cosurf actant. These droplet structures are probably the most commonly encountered 
type of microemulsion microstructure. It is worth noting that both an emulsion 
and a nanoemulsion can only occur in the form of a droplet, either as an oil-in-water 
or water-in-oil droplet. 
At intermediate oil and water compositions, it is obviously not possible for the 
microstructure to be composed of droplets of one phase dispersed in the other. 
In these cases, it is thought that a bicontinuous structure exists, in which the 
water and oil domains are separated by a regular or topologically chaotic continuous 
amphiphile-rich interfacial layer. A bicontinuous microemulsion is often the 
intermediate microstructure between an oil-in-water and a water-in-oil microemulsion, 
although a number of other microstructures such as cylinders and worm-like 
microemulsions have been reported to exist. 
In terms of its microstructure, a microemulsion is therefore a very complex 
system, and in instances where a microemulsion exists over a wide range of compositions, 
several different types of microstructure may be present.73 It is also important 
to remember that whatever the microstructure, a microemulsion is a dynamic 
system in which the interface is continuously and spontaneously fluctuating.104 
For this reason, microemulsions stabilized by polymeric surfactants may be the 
most long lived. 
1.4. Microemulsions, swollen micelles, micelles 
There is much debate in the literature as to what exactly differentiates a microemulsion 
from a micelle at low volume fractions of disperse phase. Some investigators 
have perceived a difference between microemulsions and micellar systems containing 
solubilized oil or water, and have used the terms "swollen" micellar solutions or 
solubilized micellar solutions to describe such systems. These investigators argue 
that the term microemulsion should be restricted to systems in which the droplets 
are of large enough size such that the physical properties of the dispersed oil or 
water phase are indistinguishable from those of the corresponding oil or water 
phase, thereby theoretically making it possible to distinguish between oil-in-water 
(or water-in-oil) microemulsions and micellar solutions containing small amounts 
of solubilized oil (water). However, in most cases, the transformation between 
micelles progressively swollen with oil (water) and a microemulsion containing 
an isotropic core of oil (water) appears to be gradual with no obvious transition 
point. As a consequence, there is no simple method available for determining the 
oil (water) content at which the core of the swollen micelle becomes identical to that 
of a bulk phase. Many researchers therefore use the term microemulsion to include 
130 Lawrence & Warisnoicharoen 
swollen micelles, but not micelles containing no oil (or water).34'107 In biotechnological 
applications, water-in-oil microemulsions are frequently known as reverse 
micelles and or even as nonaqueous media. 
1.5. Microemulsions and cosolvent systems 
The above broad definition does not require a microemulsion to contain any 
microstructure. In other words, it includes systems that are co-solvents, i.e. systems 
in which the constituent components are molecularly dispersed. Most researchers 
in the field agree, however, that for a microemulsion to be formed, it is important 
that the system contains some definite microstructure. In other words, there is a 
definite boundary between the oil and water phases, and at which the amphiphilic 
molecules are located and that a co-solvent is not a type of microemulsion. The only 
way to distinguish a microemulsion from a co-solvent unambiguously is to perform 
either a scattering study (light, X-rays or neutrons) or PFG-NMR measurements. 
Regions of co-solvent formation generally appear at low concentrations of oil or 
2. Microemulsions as Drug Delivery Systems 
It is clear from its description that microemulsions possess a number of properties 
that make their use as drug delivery vehicles particularly attractive. Indeed, 
microemulsions were first studied with the view of using them as potential vehicles 
for poorly-water soluble drugs, in the mid 1970s by Elworthy and Attwood.17 
However, it was not until the mid to late 1980s that they were widely investigated 
as drug delivery systems; this interest being largely the result of the arrival on the 
market of the cyclosporin A microemulsion preconcentrate, Neoral. 
Among the physical properties that make microemulsions attractive as drug 
delivery vehicle is their transparent nature, which means that the product is not 
only aesthetically pleasing, but allows easy visualization of any contamination. 
The small size of the domains present means that a microemulsion can be sterilized 
by terminal filtration.84 Furthermore, depending on the composition of the 
microemulsion, it may be possible to heat sterilize the microemulsions.39 Since oilin-
water microemulsions are able to incorporate lipophilic substances, they can 
be used to facilitate the administration of water-insoluble drugs.24 Significantly, 
the small droplet size provides a large interfacial area for rapid drug release, and 
so the drug should exhibit an enhanced bioavailability, enabling a reduction in 
dose, more consistent temporal profiles of drug absorption, and the protection of 
drug(s) from the hostile environment of the body. In addition to increasing the rate 
of drug release, microemulsions can also be used as a reservoir and actually slow 
the release of drug and prolong its effect, thereby avoiding high concentrations in 
Recent Advances in Microemulsions as Drug Delivery Vehicles 131 
the blood.64'142 Whether a drug is rapidly or slowly released from a microemulsion 
depends very much on the affinity of the drug for the microemulsion. Since 
microemulsions contain surfactants (cosurfactants) and other excipients, they may 
serve to increase the membrane penetration of drug.163'189 
A number of reviews have been presented, describing the pharmaceutical use of 
microemulsions.16'19-50'105"107'176 Since the last major review in the area was writen in 
2001, the present review will mainly deal with developments henceforth, although 
important work prior to this will be discussed when appropriate. 
2.1. Self-emulsifying drug delivery systems (SEDDS) 
Before discussing how microemulsions are being exploited in drug delivery, it 
is worth making one more distinction, namely the difference between a selfemulsifying 
drug delivery system (SEDDS) and a microemulsion. A SEDDS is a mixture 
of oil(s), and surf actant(s), ideally isotropic, sometimes containing cosolvent(s), 
which when introduced into aqueous phase under gentle agitation, spontaneously 
emulsifies to produce a fine oil-in-water dispersion.36'146 Typically, the size of the 
droplets produced by dilution of a SEDDS is in the range of 100 and 300 nm, while, 
upon dispersal in water, a SMEDDS formulation (a sub-group of the SEDDS) forms 
a transparent microemulsion with particle sizes <100 nm. ASMEEDS is also known 
as a pre-microemulsion concentrate.97 It is worth noticing that this method of producing 
a fine oil-in-water emulsion using a S(M)EEDS is identical to the low energy 
emulsification method for producing oil-in water nanoemulsions.173 It is therefore 
likely that a diluted S(M)EDDS and nanoemulsion are identically the same. 
The technique of low-energy or self-emulsification has been commercially 
exploited for many years in the agrochemical industry, in the form of emulsifiable 
concentrates of lipophilic herbicides and pesticides.146 However, it has only recently 
been introduced in the pharmaceutical industry as a tool to improve the delivery 
of lipophilic drugs by incorporating the drug into a S(M)EDDS formulation which 
is then filled into capsules.65 Once the capsule has been swallowed and its contents 
come into contact with the GI fluid, the drug containing (micro)emulsion should 
be spontaneously formed. Once the drug containing (micro)emulsion is formed, 
there should be little difference between the fate of the drug thus administered and 
the same drug administered in a (pre-formulated) microemulsion, although the 
droplets formed from the S(M)EDDS tend to be of a larger size. One advantage of 
administering a drug in a SMEEDS as opposed to a pre-formulated microemulsion, 
is its relatively small volume which can be incorporated into soft or hard gelatin 
capsules, convenient for oral delivery. 
To date, there has been a good amount of commercial success for the first selfmicroemulsifying 
drug delivery systems (SMEDDS) on the market, namely Neoral 
(cyclosporin A). In addition, the recent commercialization of two self-emulsifying 
132 Lawrence & Warisnoicharoen 
formulations, namely Norvir (ritonavir) and Fortovase (saquinavir), has undoubtedly 
increased the interest in SEDDS and other emulsion-based delivery systems 
to improve the delivery of a range of drugs of varying physico-chemical 
However, there are a number of reasons why S(M)EDDS are not in greater 
widespread use, but the main reason is probably the stability of the diluted SEDDS, 
which is in fact a thermodynamically unstable emulsion (although it may exhibit 
some limited kinetic or "meta" stability). It should be noted however that as a 
SEDDS is either diluted just prior to administration or else in the body, the required 
droplet stability is less than 6 hrs (i.e. the transit time of materials down the small 
Although most studies of SEEDS have utilized isotropic liquids, the earliest 
reports of these self-emulsifying systems using pharmaceutical materials are in 
fact related to pastes based on waxy polyoxyethylene n-alkyl ethers.67 In the context 
of drug delivery via self-emulsifying systems, isotropic liquids are generally 
preferred to waxy pastes because if one or more excipient(s) crystallize(s) on cooling 
to form a waxy mixture, it is very difficult to determine the morphology of 
the materials. Despite this, there is currently a general move towards formulating 
semi-solid SEDDS. For example, attempts have been made to transform SEDDS into 
solid dosage forms by addition of large amounts of solidifying excipients such as 
adsorbents and polymers15,134 Unfortunately, as the ratio of SEDDS to solidifying 
excipients required for this approach is very high, this leads to problems in formulating 
drugs having limited solubility in the oil phase. Recent attempts have been 
made to reduce the amount of solidifying excipients by gelling the SEDDS with 
colloidal silicon dioxide.141 
Khoo et al.93 have recently reported the preparation of a halofantrine-containing 
lipid-based solid self-emulsifying system using either Vitamin E TPGS or a blend of 
Gelucire 44:14:Vitamin E TPGS as the base. Upon dispersal, these systems produced 
dispersions that the authors described as microemulsions. Studies in fasted dogs 
showed that these solid dispersions exhibited a five- to seven-fold improvement 
in absolute oral bioavailability, when compared with the commercially available 
tablet formulation. 
In a different approach, Nazzal et alP2 have determined the potential of a 
reversibly induced re-crystallized semi-solid self-nanoemulsifying drug delivery 
system, based on a eutectic interaction between the drug and the carrying agent, as 
an alternative to a conventional SEDDS. In these eutectic-based self-nanoemulsified 
systems, the melting point depression method allows the oil phase containing the 
drug itself to melt at body temperature from its semisolid consistency, and disperse 
to form emulsion droplets in the nanometer size range. Emulsion systems based on 
a eutectic mixture of lidocaine-prilocaine,135 and lidocaine-menthol87 have been 
Recent Advances in Microemulsions as Drug Delivery Vehicles 133 
used in the preparation of topical formulations. However, little is known of the use 
of eutectic mixtures for the preparation of self-(micro) emulsified formulations. 
2.2. Related systems 
There are a number of other putative delivery systems that are closely related to, 
or are prepared from, a microemulsion. These systems include a variety of gel 
formulations (including microemulsion-based gels, ringing gels, microemulsion 
gels) and double microemulsions. 
2.2.1. Microemulsion gels 
Oil-in-water microemulsions can be readily gelled or thickened by the addition of 
a non-interacting, water-soluble polymer such as polyHEMA,158 Carbopol 94044 or 
carrageenan179 to form clear "microemulsion gels". In these cases, it is the external 
aqueous phase that is gelled, while the microemulsion droplets are unperturbed. 
The structure of the resulting "microemulsion gel" is quite different, if it is prepared 
using an interacting polymer, such as stearate-polyethylene oxide-stearate. 
In this instance, the hydrophilic mid-block of the polymer is located in the continuous 
aqueous phase, while the hydrophobic end blocks are dissolved in the oil 
droplets, thereby connecting the various microemulsion droplets and resulting in 
the formation of a transient gel network.159 Clear, "microemulsion gels" are also 
sometimes obtained at surfactant and/or oil concentrations just outside the oilin-
water microemulsion region.180 Sometimes, the resultant gel "rings" or vibrates 
when tapped.180 The ringing is due to the resonance of shear modes within the gel 
body.167 Neither of these "microemulsion gels", which are water continuous, are 
true microemulsions, which are fluid by definition. 
Clear gels can also be formed in oil continuous systems. For example, a gel can 
be formed when water is added to reverse micellar solutions of lecithin-in-oil.5,12'161 
Here, the water causes the worm-like lecithin reverse micelles to intertwine and 
form a gel. In addition, gels, widely known as microemulsion-based gels, can be 
formed from water-in-oil microemulsions stabilized (predominately) by the dichain 
surfactant sodium bis (2-ethylhexyl) sulfosuccinate (AOT), when gelatin, the 
natural amphiphilic polymer is added.70,148 Microemulsion-based gels have now 
been prepared in systems in which a large amount of the AOT has been replaced 
by nonionic surfactant;88'89 or more recently, using in place of AOT, the single chain 
surfactant, cetyltrimethylammonium bromide in combination with pentanol.116 In 
these gels, the gelatin is thought to form water-continuous channels between the 
microemulsion-droplets. These microemulsion-based gels are very unusual in that, 
although they are oil continuous, they are electrically conducting. In addition, the 
134 Lawrence & Warisnoicharoen 
continuous oil behaves as if it were still a fluid, even though placing a gel in a 
solution of the oil does not dissolve it. 
All of these "microemulsion gels" have potential, or are being explored for use 
as drug delivery systems. Of particular interest is the fact that the gels possess the 
properties of being transparent, infinitely stable and readily prepared using only 
the mildest of mixing. In addition the wide range of microemulsion gels available 
means that it is possible to select the gel of the required consistency for application 
to large areas of skin, the nasal membrane, vaginal and buccal membranes and 
for permeation enhancement. Microemulsion-based gels have been explored as 
vehicles for the iontophorectic delivery of drugs.88 
2.2.2. Double or multiple microemulsions 
Double (or multiple) emulsions have attracted much interest as potential drug delivery 
vehicles. For example, adding a water-soluble drug into the internal aqueous 
phase of a water-in-oil-in-water emulsion may allow the sustained release of the 
water-soluble drug.59 A double microemulsion should offer similar advantages over 
the rate of drug release of entrapped solutes. Double emulsions are notoriously difficult 
to formulate due to the requirement to have one surfactant (or mixture of surfactants) 
to stabilize the first (internal) emulsion and a second surfactant (or mixture 
of surfactants) of quite different physico-chemical properties to stabilize the second 
emulsion. Although a few papers have detailed the production of nanoparticles 
from systems they described as double microemulsions,35,56'184 the term "double 
microemulsion" in this context is very misleading, as it refers to the mixing of two 
water-in-oil microemulsions of comparable composition, but containing different 
solute in the aqueous phase. 
There are however two papers which describe the preparation of double (oilin-
water-in-oil) microemulsions. In the first, Castro et al.,30 report spectroscopic 
studies of nifedipine in Brij 96 based oil/water/oil multiple microemulsions. In the 
second, Carli et al.,29 detail the preparation of an oil-in-water-in-oil microemulsion 
from an oily phase of either polyglycolized glycerides or a mixture of mono-, diand 
tri-glycerides, which is microemulsified using a mixture of water and surfactant 
(soy lecithin and Tween 80). The resultant o /w microemulsion is subsequently redispersed 
in an oily phase to produce the double (o/w/o) microemulsion.29 
2.3. Processed microemulsion formula tions 
2.3.1. Solid state or dry emulsions 
In practical terms, a solid dosage form is preferable to a liquid dosage form in respect 
of convenience, ease of handling and accurate dosing. Consequently, a number of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 135 
researchers have attempted to develop powdered, re-dispersible emulsion-derived 
formulations, known as solid state or dry emulsions. Such solid-state emulsions 
can be used to modulate the release rates of emulsified compound.128 Dry emulsions 
have been variously prepared by removing water from an oil-in-water emulsion, 
using water-soluble182 or -insoluble150 solid carriers or indeed a mixture 
of both water-soluble and water-insoluble carriers,143 by rotary evaporation,128 
lyophilization,115 or spray drying.55'127 
Attempts have also been recently made to prepare dry microemulsions using 
similar methodology. For example, Moreno et al.126 have lyophilized an amphotericin 
B-containing lecithin-based oil-in-water microemulsion in the presence 
of 5wt% mannitol. The lyophilized product was an oily cake from which the 
microemulsion could be easily reconstituted over several months. The rationale 
for developing the water-free formulation was to avoid the hydrolysis of lecithin, 
which occurs upon its dispersal in water, thereby preventing any deterioration of 
the formulation upon storage. Overall, the lyophilized lecithin based oil-in-water 
microemulsions appear to be valuable systems for the delivery of amphotericin B, 
with regard to ease and low-cost of manufacturing and their stability and safety, 
compared with other formulations already in the market. 
In a recent paper, Carli et al.29 reported an alternative approach to prepare a 
"dry" formulation known as a nanoemulsified composite of the coenzyme Q10, 
ubidecarenone. This composite is prepared by incorporating the ubidecarenone 
into the inner phase of the double microemulsion, which is then deposited onto a 
solid microporous carrier such as cross-linked polyvinylpyrrolidone. Among the 
advantages offered by this approach are good processing and storage properties, 
easy re-dispersibility in water, high bioavailability and maintenance of the submicron 
size of the released droplets. 
Kim et al.,97 have prepared entric-coated solid state premicroemulsion concentrates 
by first preparing a pre-microemulsion concentrate containing 10 wt% of the 
drug cyclosporin A, 18.5 wt% of a medium chain triglyceride, 51 wt% of surfactant 
and 20 wt% of cosurfactant. The pre-microemulsion concentrates were then 
enteric-coated as films using polymers, such as sodium alginate, Eudragit L 100 
and cellulose acetate phthalate, and the resulting films were pulverized to produce 
powdered, dry, enteric coated premicroemulsion concentrates. Using this approach, 
the authors successfully prepared a once-a-day oral dose form of cyclosporin A. 
3. Formulation 
Microemulsions are far more difficult to formulate than emulsions because their formation 
is a highly specific process involving spontaneous interactions among their 
constituent molecules. In addition, in a number of cases, effects due to the order of 
136 Lawrence & Warisnoicharoen 
the mixing of the component molecules have been observed. Since no adequate theory 
currently exists to predict from which molecules microemulsions can be formed, 
mainly because of the requirement to determine a number of unknown parameters, 
microemulsion formulations are generally developed empirically, although some 
useful practical guidance as to the choice of the constituent components can be 
found in the literature.105,107 
A recognized and classical approach to microemulsion formulation is to undertake 
a systematic study of the phase behavior of the systems understudying utilizing 
of phase diagrams. A major drawback of this approach is the considerable time 
it takes to develop the phase diagram, especially considering the combination of 
possible oil, surfactant and cosurfactant, and the fact that time may be necessary for 
a system to equilibrate. Heat and sonication are therefore often used, particularly 
with systems containing nonionic surfactants, to speed up the formation process. 
While there are now commercially available automated systems to prepare phase 
diagrams,78 the chief drawback of these systems is their cost. 
A number of attempts have been recently made to use modeling to predict 
microemulsion formation, thereby aiding in the formulation of microemulsions. 
A range of modeling techniques have been used including artificial 
neural networks,3'4'8'149 genetic algorithms3,174 and a combination of data mining, 
computer-aided molecular modeling, descriptor calculation and multiple linear 
regression techniques.174,175 Unfortunately, however, all of these techniques 
require a considerable amount of work prior to prediction, thereby restricting 
their potential usefulness. Furthermore, the amount of work required for the predictions 
increases as the number of components of the microemulsion increase; 
microemulsions formulated from five components (i.e. oil, water, surfactant, cosurfactant, 
electrolyte and drug) are not uncommon in pharmaceutical use. To the 
authors' knowledge, to date, no work has been performed predicting how much 
drug can be incorporated into a microemulsion and whether the presence of drug 
has any effect on microemulsion phase behavior. This is an important ommision 
as microemulsions cannot be considered to be inert, since the presence 
of drug in some instances (greatly) influences phase behavior (see for example 
Ref. 138). 
3.1. Surfactants and cosurfactants 
The selection of components for the preparation of microemulsions suitable for 
pharmaceutical use involves a consideration of their toxicity, and if the systems 
are to be used topically, their irritancy and sensitizing properties as well. There are 
a number of surfactant and cosurfactants that are considered acceptable for use 
as excipients in pharmaceutical formulation.153 Strickley172 has recently reviewed 
Recent Advances in Microemulsions as Drug Delivery Vehicles 137 
those surfactants and cosurfactants currently used in commercially available oral 
and intravenous formulations. 
In the general scientific literature, by far the most widely used surfactant to prepare 
a microemulsion is the double chain, ionic surfactant, sodium bis (2-ethylhexyl) 
sulfosuccinate (AOT), although a large number of studies have used the single 
chain, nonionic surfactants of the type QEj, where i is the number of carbons in the 
alkyl chain, C, and j is the number of ethylene oxide units in the polyoxyethylene 
chain, E. Both AOT and the QEj surfactants possess the important advantage of 
being able to form microemulsions in the absence of a cosurfactant,121'185 unlike 
most other types of surfactant such as the widely studied single chain, ionic surfactant, 
sodium dodecyl sulphate (SDS), which will only form microemulsions in the 
presence of an alcohol cosurfactant. Neither AOT nor SDS would be considered to 
be apprpropriate for the preparation of pharmaceutically acceptable microemulsions, 
even though they are listed in the Pharmaceutical Excipient Handbook, 
Rowe et al.l5i 
As a general rule, nonionic and zwitterionic surfactants tend to be less toxic 
than ionic surfactants and are therefore more widely used as pharmaceutical 
excipients.154 Assuming that the surfactants do not degrade into toxic materials, 
surfactants that posses biodegradable/chemically unstable linkers tend to exhibit 
less chronic toxicity than those that are chemically stable. For example, as a group, 
the polyoxyethylene n-acyl surfactants exhibit ~ten times less chronic toxicity than 
their n-alkyl counterparts, mainly due to their quicker degradation time of days as 
opposed to weeks. When it comes to comparing acute toxicity, the two groups of 
surfactants exhibit comparable toxicity. 
Perhaps the most widely used nonionic surfactants in pharmaceutical formulations 
are the polyoxyethylene sorbitan n-acyl esters, i.e. the Tweens and in particular, 
Tween 20 and Tween 80, both of which are used parenterally and orally. In addition, 
polyoxyethylene derivatives of the triglyceride, castor oil have acceptability 
for intravenous administration. Other pharmaceutically acceptable surfactants are 
the polyoxyethylene n-alkyl ethers and n-acyl esters, although both these groups of 
surfactant tend to be restricted for topical use.154 Other nonionic surfactants that are 
currently attracting much pharmaceutical interest, although they do not yet have 
acceptability, are the polyglycerol n-acyl esters, the n-alkyl amine N-oxides and 
the w-alkyl polyglycosides (or sugar surfactants). The n-alkyl polyglycosides have 
attracted much pharmaceutical interest, not because of their excellent biodegradability, 
but because they can be manufactured from renewable resources. All of the 
aforementioned surfactants have been used to prepare microemulsions, generally 
as sole surfactant, the only exception being the w-alkyl polyglycosides, which tend 
to require the presence of a cosurfactant. 
Pluronics (or poloxamers) of the type poly(ethylene oxide)-poly(propylene 
oxide)-poly(ethylene oxide) (PEO-PPO-PEO) are another class of pharmaceutically 
138 Lawrence & Warisnoicharoen 
attractive surfactant. Interestingly, most reports detailing the use of polymeric surfactants 
to stabilize a microemulsion describe the preparation of water-in-oil, generally 
in conjunction with a second surfactant.96,181 Siebenbrodt and Keipert162 have 
reported the formation of a triacetin-in-water microemulsion using Pluronic L 64 
as sole surfactant. Lettow et al.,m have used Pluronic PI23 as sole surfactant to 
prepare oil-in-water microemulsions, incorporating a 1:1 oil:P123 weight ratio of 
either 1,3,5-trimethylbenzene or 1,2-dichlorobenzene. 
Finally, the pharmaceutically acceptable zwitterionic lecithin has been extensively 
used as a surfactant, however, with very few exceptions, it is not possible to 
prepare a microemulsion using lecithin as sole surfactant. Generally, lecithin is combined 
with another surfactant such as Tween 80, or a cosurfactant such as ethanol, 
when formulating microemulsions. 
Although ethanol is considered to be pharmaceutically acceptable, typical 
cosurf actants such as propanol and butanol are not. In addition to toxicity issues, the 
use of such cosurfactants, which may possess partial oil and water-solubility, can 
lead to problems with the dilutability of the microemulsion. This is a particular issue 
if the microemulsion is to be administered orally or parenterally. Consequently, a 
number of researchers have explored the use of a second surfactant as cosurfactant 
when formulating a microemulsion. Microemulsions thus prepared tend to be very 
stable against dilution, as the "cosurfactant" generally has little solubility in either 
the oil or aqueous phase. Alternately (pharmaceutically acceptable), short chain 
mono- and di-glycerides have been used in place of a short chain alcohol to successfully 
prepare microemulsions. In a number of instances, short chain fatty acids such 
as sodium caprylate have been used as cosurfactants, primarily for the formation 
of microemulsions for oral delivery; sodium caprylate is known to enhance absorption 
of drugs across the gastrointestinal tract. A number of researchers have also 
used cosolvents such as the polyhydric alcohols, sorbitol, glycerol and propylene 
glycol to aid microemulsion formation. In a number of instances, these materials 
have been described as "cosurfactants", which quite clearly do not sit in the interfacial 
surfactant monolayer. Rather, they tend to exert their effect by altering the 
solvent properties of the polar phase. 
3.2. Oils 
Most reports in the chemical literature detail the preparation of microemulsions 
using aromatic oils such as benzene and short chain alkanes such as hexane. "Pharmaceutical" 
oils, unlike those used in the chemical and agrochemical industries, 
tend to be large in terms of molecular weight and therefore volume, and are relatively 
polar. Both of these properties tend to work against the oil when it comes to 
formulating it in a microemulsion, as it is well established that smaller molecular 
volume oils are easier to solubilize and are solubilized to a greater extent than larger 
Recent Advances in Microemulsions as Drug Delivery Vehicles 139 
oils.2 Although there are reports that in some systems, particularly those containing 
surfactants with long, unsaturated hydrophobes such as polyoxyethylene (10) oleyl 
ether, the largest molecular volume oil is solubilized to a greater extent than some of 
the smaller molecular volume oils.122 The most commonly used "pharmaceutical" 
oils are medium and long chain triglycerides, and esters of fatty acids such as ethyl 
oleate, isopropyl myristate are popular. 
It has become common practice for researchers to screen the solubility of drug 
in the various components of the microemulsion, in order to predict the optimal 
composition of the final formulation. However, extreme care has to be exercised 
when using this approach, as very often, the solubility in the final microemulsion 
formulation does not correlate well with that seen in the various components. 
3.3. Characterization 
It is noticeable that in contrast to their ease of preparation, it is very difficult to 
establish the microstructure of a microemulsion. Yet, such information is important 
as it may influence the drug behavior of the microemulsion in use. For example, it 
is known that the microstructure of the microemulsion may alter the release rate of 
any incorporated solute.1'95 
Currently, a range of physico-chemical techniques are used to characterize 
microemulsions. These techniques are often used in tandem to obtain a better picture 
of the system, as it is unlikely that any one technique alone will give sufficient 
information.144 Scattering techniques (light, neutron and X-ray) and pulsed 
field gradient NMR are generally used to determine the microstructure of the 
microemulsion. One serious limitation with characterizing microemulsions is that 
most techniques rely on the concentration of disperse phase being low enough to 
avoid particle-particle interactions, as an estimated volume fraction of 1 vol% is 
suitable.123 The requirement is a particular problem with microemulsions that contain 
cosurfactants that partition between the oil and water phases, because these 
systems frequently undergo a change upon dilution. 
4. Routes of Administration 
Although most of the original work exploring microemulsions as drug delivery 
vehicles examined their potential for oral drug delivery, microemulsions have now 
been explored as vehicles for most routes of administration. Currently, they are 
probably most widely studied for their potential as transdermal delivery vehicles. 
4.1. Oral 
Microemulsions (and SMEDDS) have been widely studied as oral drug delivery 
vehicles. Indeed, the first commercially available "microemulsion" formulation was 
140 Lawrence & Warisnoicharoen 
a premicroemulsion concentrate of the lipophilic peptide, cyclosporin A. This formulation, 
known commercially as Neoral, was introduced onto the market in the 
late 1980s and immediately attracted much attention, mainly because of the high 
and reproducible bioavailability it produced, but also because developments in 
biotechnology at that time meant that it had never been easier to produce on a large 
scale therapeutically-relevant protein and peptides. Unfortunately, because of their 
physico-chemical properties, in particular their large size and poor stability, proteins 
and peptides are very difficult to formulate. Microemulsions offered an attractive 
solution to this problem, and consequently, most of the original exploratory 
studies on microemulsions as drug delivery vehicles were spent developing oral 
protein/peptide microemulsion formulations. 
4.1.1. Proteins and peptides 
As the majority of therapeutic proteins and peptides are hydrophilic and watersoluble, 
most studies utilizing microemulsions as vehicles for such molecules have 
exploited water-in-oil microemulsions. After cyclosporin A, which is unusually 
highly lipophilic, for a therapeutic peptide, the most widely studied peptide is 
insulin, with much of the early work in this area being performed by Ritschel.152 
For example, Kraeling and Ritschel101 compared the peroral microemulsion formulation 
of insulin and capsule forms and determined that the microemulsion 
formulation increased the bioavailability of the insulin. Recently, more complex 
microemulsion-based systems have been developed in an attempt to improve 
the extent of insulin absorption. For example, a recent study performed by 
Natnasirichaiku et al.186 showed a significant improvement in the oral bioavailability 
of insulin (in diabetic rats) when administered in nanocapsules dispersed 
in a water-in-oil microemulsion. Santiago et al.155 have developed a new, enteric 
oral dosage form of insulin, in which an association of insulin and cyclodextrin 
contained within a microemulsion is processed into granules. In the most recent 
study aimed at developing an oral formulation of insulin, Iek et al.77 used a conventional 
lecithin-based water-in-oil microemulsion formulation prepared from 
21.6 wt% water, 37.6 wt% Labrafil M 1944 CS as oil and stabilized by 40.8 wt% of a 
1:1 weight ratio of lecithin (Phospholipon 90G) and ethanol. In addition to insulin 
(21.6IU/g water), some of the microemulsions contained the enzyme inhibitor 
aprotinin (2500KlU/g water). Although it is the first time that a microemulsion 
formulation has contained both a protein/peptide and an enzyme inhibitor, the 
concept of adding an enzyme inhibitor, to a formulation containing a peptide in an 
attempt to reduce its degradation is not new.188 The plasma glucose and insulin levels 
of the rats after intragastric administration of the formulations to both diabetic 
and non-diabetic rats were significantly different from those obtained after oral 
Recent Advances in Microemulsions as Drug Delivery Vehicles 141 
administration of an aqueous insulin solution. Although the addition of aprotinin 
to the microemulsion containing insulin increased bioavailability when compared 
with those not containing it, the difference was insignificant. 
Other peptides formulated as water-in-oil microemulsions in an attempt to 
improve their oral absorption include RGB peptides,37'38 and more recently, Nacetylglucosaminyl-
N-acetylmuramyl dipeptide (GMDP).119 The poor bioavailability 
of GMDP has been attributed to both its poor stability in the lumen of the 
gastrointestinal tract and its poor intestinal permeability. When GMDP was administered 
intraduodenally in a water-in-medium-chain trigylceride microemulsion, a 
ten-fold increase in bioavailability was observed, i.e. a bioavailability of 80.2% was 
achieved as opposed to 8.4%, seen after administration of an aqueous solution of 
GMDP. This increase in bioavailability is consistent with the work of Constantinides 
et al.37,3S who utilized a similar medium chain triglyceride based microemulsion to 
increase the oral bioavailability of the water-soluble peptide SK&F 106760, after 
intraduodenal administration to rats. 
Ke et al.92 have recently reported an attempt to develop water-in-oil microemulsions 
suitable for the incorporation of therapeutic proteins and peptides using 
a medium chain triglyceride, water and tocopheryl polyethylene glycol 1000 
succinate (TPGS) as the primary surfactant. However, as TPGS could not form 
microemulsions when used as sole surfactant, it was mixed with a second surfactant, 
either Tween 20,40,60 or 80, at a weight ratio in the range of 4:1 to 1:4. A range 
of glycols and polyols were examined as cosurfactants. Although stable, transparent 
microemulsion and gel regions were identified, the extent of these regions was 
influenced by the precise nature and the amount of the secondary surfactant and 
cosurfactant. For example, Tween 80, which is an ester of the unsaturated CI 8 fatty 
acid, oleic acid, was more effective in forming a microemulsion than Tween 60, 
which is an ester of the saturated C18 fatty acid, stearic acid. In this study, although 
the microemulsions were ultimately intended for use as delivery vehicles for protein 
or peptide drugs, they were not examined for this purpose. 
4.1.2. Other hydrophilic molecules 
Other water-soluble therapeutic molecules that have been administered in 
microemulsions include the aminoglycoside antibiotic, gentamicin74 and the biologically 
active polysaccharide, heparin." In common with all aminoglycosides, 
gentamicin is highly polar and is therefore considered unlikely to be absorbed 
from the gastrointestinal tract via simple diffusion. In order to facilitate the transmucosal 
delivery of the drug, Hu et al74 prepared a SMEDDS formulation of gentamicin 
using a range of surfactants. When Labrasol was used as surfactant, a 54.2% 
bioavailability of gentamicin was obtained, compared with only 8.4 and 3.4% when 
142 Lawrence & Warisnoicharoen 
Tween 80 and Transcutol P were respectively used. Labrasol was also found to 
inhibit intestinal secretory transport from the intestinal enterocytes, providing the 
formulation with the additional benefit of inhibiting the efflux of gentamicin out of 
the enterocytes into the GI lumen. 
Due to its low bioavailability, heparin is generally administered by injection. In 
an attempt to formulate an orally active version of heparin, Kim et al." synthesized 
a low molecular weight heparin (LMWH)-deoxycholic acid (DOCA) conjugate 
(termed LMWH-DOCA) and formulated it in a water-in-oil microemulsion using 
as oil, the medium chain trigylceride, tricaprylin, a mixture of Tween 80 and Span 
20 surfactants, LMWH-DOCA and water (volume ratios of 5:3:1:1 respectively). 
Oral administration of LMWH-DOCA in the water-in-tricaprylin microemulsion 
to mice resulted in a bioavailability of 1.5%. Toxicity studies suggested that the 
enhancement in bioavailability, observed with the DOCA-conjugated LMWH, was 
administered in a microemulsion not due any local toxicity such as disruption or 
damaging of the intestinal membrane. 
4.1.3. Hydrophobic drugs 
A number of poorly water-soluble, low molecular weight, lipophilic drugs have 
also been formulated in microemulsions (or SMEEDS) for oral delivery including 
nitrendipine,90 danzol145 halofantrine94 and biphenyl dimethyl dicarboxylate.98 
These studies serve as an illustration of how important it is to understand the 
influence on microemulsion formation of the various formulation components. It 
is worth commenting that the main use of SMEEDS formulations is for the oral 
administration of lipophilic drugs. 
Formulating nitrendipine in a SMEEDS formulation, composed of a 1:1 (w/w) 
mixture of glycerol monocaprylic ester (MCG) and propyleneglycol dicaprylic ester 
(DCPG) and nonionic surfactant (various), was observed to significantly enhance 
its absorption when compared with a suspension or an oil solution,90,91 and served 
to reduce the effect of the presence of food on its absorption. However, the absorption 
profile of nitrendipine was seen to vary with the type of surfactant used; 
absorption was rapid from the Tween 80-stabilized formulation, while the HCO-60- 
based formulation gave a prolonged plasma concentration profile. No absorption of 
nitrendipine was observed from the formulation containing BL-9EX (polyoxyethylene 
alkyl ether, C12E9). Damage to the gastrointestinal mucosa also differed with 
the type of surfactant employed. HCO-60 and Tween 80-based formulations were 
mild to the organs, while BL-9EX-based formulation caused serious damage. 
The study of Porter et al.145 appropriately demonstrates the effect of changing 
the nature of the trigylceride involved in the formulation on drug absorption. 
These workers studied three lipid-based danazol formulations; namely a long-chain 
triglyceride solution (LCT-solution), a SMEDDS based on long (C18) chain lipids 
Recent Advances in Microemulsions as Drug Delivery Vehicles 143 
(LC-SMEDDS) and a SMEEDS formulation containing medium (C8-C10) chain 
lipids (MC-SMEDDS). These formulations were administered to fasted beagle dogs 
and their absorption, compared with that obtained with a micronized danazol formulation 
administered postprandially and in the fasted state. Although both the 
LCT-solution and LC-SMEDDS formulations were found to significantly enhance 
the oral bioavailability of danazol, when compared with fasted administration of the 
micronized formulation, the MC-SMEDDS produced little improvement in danazol 
bioavailability. This result was partly attributed to the fact that upon digestion of 
the medium-chain formulation, significant drug precipitation was observed. 
Khoo et al.94 also considered the effect of formulating halofantrine as a 
pre-microemulsion concentrate in a formulation based on either a medium- or longchain 
triglyceride. Both formulations, which were administered as soft-gelatin capsules, 
contained the same amount of medium or long chain trigylceride and were 
stabilized by the same surfactant/cosurfactant mixture, consisting of Cremophor 
EL and ethanol. Although the plasma levels of the drug were not significantly 
different between the two formulations, the amount of drug absorbed lymphatically 
varied in that 28.3% of the dose administered in the long-chain trigylceride 
formulation was transported lymphatically, as opposed to only 5.0% of the dose 
administered in the medium-chain formulation. 
Kim et al.9S attempted to improve the solubility and bioavailability of biphenyl 
dimethyl dicarboxylate, a drug used in treating liver diseases, by formulating it as 
a premicroemulsion concentrate. In order to optimize drug loading in the formulation, 
these workers screened drug solubility in a range of surfactants and oils, 
and on the basis of these results selected: Tween 80 and Neobee M-5. However, care 
must be taken when using this approach to optimize the formulation with respect to 
drug loading, as it has shown that solubility of drug in the bulk components is not 
a reliable indicator of solubility, in the final microemulsion formulation.120,122 The 
danger of predicting drug solubility in the final formulation, on the basis of bulk 
solubility, can be seen in the study of Kim et al.98 where the solubility of the drug 
in a formulation consisting of a 2:1 weight ratio of Tween 80 to Neobee M-5 was 7 
times that of the formulation containing a Tween 80:Neobee M-5 weight ratio of 1:4, 
despite the solubility of the drug in Neobee M-5 being 10 times that seen in Tween 
80. The final formulation, which consisted of 35 wt% triacetin and 65 wt% Tween 
80 and Neobee M-5 at a weight ratio of 2:1, greatly enhanced the oral bioavailability 
of BDD, possibly due to the increased solubility of the drug and its immediate 
dispersion in the gastrointestinal tract. 
Itoh et al.79 optimized the formulation of the poorly water-soluble 
drug N-4472, N-[2-(3,5-di-tert-butyl-4-hydroxyphenethyl)-4,6-difluorophenyl]-N- 
[4-(Nbenzylpiperidyl)] urea, by complexing it with L-ascorbic acid and incorporating 
the complex into a SMEEDS comprising Gelucire 44/14, HCO-60 and sodium 
dodecyl sulfate. Upon dilution with water, the SMEEDS formulation produce a fine 
144 Lawrence & Warisnoicharoen 
dispersion of 18 nm droplets which were stable over the pH range of 2.0 to 7.0. The 
oral bioavailability of the drug was between 2-4 times that which was obtained 
with an aqueous solution of the complex. 
4.2. Buccal 
To date, very little work has been performed on investigating the use of microemulsions 
as vehicles for buccal delivery. In 1988, Ceschel et a\?x showed that the penetration 
of the essential oil, Salvia sclarea L. through porcine buccal mucosa in vitro 
was increased when formulated as a microemulsion, as opposed to the pure essential 
oil. Scherlund et al.5S investigated the potential of lidocaine and prilocaine 
thermosetting microemulsions and mixed micellar solutions as drug delivery systems 
for anesthesia of the periodonlal pocket. The formulations contained between 
2-10 wt% of a eutectic mixture of lidocaine or prilocaine (melting point 18C), while 
the block copolymer surfactants, Pluronic F127 and F68, were present at between 
13 and 17 wt% for F127, and between 2 and 6 wt% for F68. F127 was chosen, as it is 
known to gel at body temperature and it is important that the formulation is easy 
to apply, remain at the application site, have a fast onset time, be non-irritant, and 
stable under normal storage conditions. The pH of the formulations was varied 
between 5 and 10. Most of the combinations were found to result in clear solutions, 
presumably oil-in-water microemulsions or mixed micellar solutions, depending 
on the pH of the system. At low pH, lidocaine and prilocaine are positively charged, 
and they could be expected to behave largely as water-soluble cationic surfactants, 
hence possibly forming mixed micelles. On the other hand, at high pH, the drug 
substances are poorly soluble and could be expected to act largely as hydrophobic 
solutes and form the core of the microemulsion droplets. 
4.3. Parenteral 
In recent years, considerable emphasis has been given to the development of 
injectable microemulsions (o/w) for the intravenous delivery of drug, in order 
to increase the solubility of the drug39'138'139 to reduce drug toxicity,25'26'126 to 
reduce hypersensitivity,72 and to improve drug solubility and reduce pain upon 
injection.109 A very recent development is the formulation of microemulsions as 
long circulating vehicles, and more recently, as drug tageting agents. In addition, 
water-in-oil microemulsions have been investigated as depot vehicles for the intramuscular 
delivery of drugs.22'64 
The first published study which established the potential of microemulsions 
for use in intravenous delivery was probably that of von Corswant 
et al. in Ref. 39. These researchers prepared a pharmaceutically acceptable, 
bicontinuous microemulsion from a medium-chain triglyceride oil, poly(ethylene 
Recent Advances in Microemulsions as Drug Delivery Vehicles 145 
glycol) 400 and ethanol cosolvents and stabilized by soybean phosphatidylcholine 
and poly(ethylene glycol)(660)-12-hydroxystearate. Prior to administration, the 
microemulsion required dilution with a suitable aqueous phase. Upon dilution, the 
microemulsion formed an oil-in-water microemulsion with droplets of size between 
60 and 200 nm, smaller than the size of the droplets in a commercial intravenous 
emulsion, namely Intralipid. From their animal studies, the authors concluded that 
the microemulsion they developed was suitable for administion by intravenous 
infusion to conscious rats. Unfortunately, although the researchers did determine 
drug solubility in the bicontinuous microemulsions, they did not report this. 
Park and Kim138 also investigated the formulation of poorly water-soluble 
flurbiprofen at ~8 times its aqueous solubility into an oil-in-water microemulsion 
suitable for intravenous administration. The microemulsions were prepared 
using varying weight ratios of oil (ethyl oleate) to surfactant (Tween 20), and contained 
a range of isotonic solutions as the polar (aqueous) phase. Unfortunately, 
insufficient information was supplied regarding the precise compositions of the 
microemulsions, in particular, how much oil and surfactant were present, so as to 
draw conclusions about the formulation; (perhaps surprisingly) the ratio of oil to 
surfactant used did not seem to have any effect on the amount of drug solubilized 
and that the presence of too much drug had a destabilizing effect on the microemulsion. 
Disappointingly, the pharmacokinetic parameters of flurbiprofen, after intravenous 
administration of flurbiprofen-loaded microemulsion to rats, were also 
not significantly different from those of flurbiprofen in phosphate buffered saline 
solution. In a later publication, Park et a/.138 overcame the problem of stability 
seen in their earlier study by replacing the surfactant Tween 20 with lecithin and 
distearoylphosphatidyl- ethanolamine-N-poly(ethyleneglycol) 2000 (DSPE-PEG) 
and using ethanol as a cosolvent. Due to the presence of the long chain polyoxyethylene 
groups on the exterior surface of the microemulsion droplets, it was 
perhaps unsurprising that the biodistribution of flurbiprofen administered in this 
microemulsion was quite different. In particular, reticuloendothelial uptake of flurbiprofen 
decreased, suggesting that it may ultimately be possible to target drugs 
incorporated in this microemulsion to different sites of the body. 
As part of a series of papers, Brime et al.25,26 and Moreno et al.126 prepared a 
novel amphotericin B lecithin-based oil-in-water microemulsion, in an attempt to 
produce a formulation with less toxic effects than the currently available commercial 
formulation, Fungizone. The microemulsion which contained as oil isopropyl 
mystriate and a mixture of either Tween 80 or Brij 96 with lecithin as surfactant. 
In some instances, formulation was lyophilized in an attempt to increase its stability. 
The overall results of the toxicity studies were encouraging as the amphotericin 
B-containing microemulsions exhibited a low toxicity, suggesting a potential 
therapeutic application. 
146 Lawrence & Warisnoicharoen 
Zhang et al.m prepared a lecithin-based SMEDDS formulation of the drug 
norcantharidin. Upon dilution, the release rate of norcantharidin contained in the 
SMEEDS formulation was found to be dependent on the size of the disperse phase 
and the type of lecithin used. Interestingly, although norcantharidin was poorly 
soluble in the ethyl oleate and only slightly soluble in water, microemulsions containing 
ethyl oleate oil exhibited a significant increase in solubilization over the 
corresponding aqueous solution. 
Clonixic acid is currently marketed in salt form because of its poor watersolubility. 
However, the commercial dosage form causes severe pain after intramuscular 
or intravenous injection. To improve the apparent aqueous solubility of 
clonixic acid and to reduce the pain it causes on injection, Lee et al. (2000) incorporated 
3 mg/mL clonixic acid into oil-in-water microemulsions (size 120 nm) prepared 
from pre-microemulsion concentrate of castor oil, and a mixture of Tween 
20 and Tween 85 surfactants (present in a weight ratio of 5:12:18). Although the 
microemulsion formulation significantly reduced the number of rats licking their 
paws as well as the total licking time, suggesting less pain induction by the 
microemulsion formulation; the pharmacokinetic parameters of clonixic acid after 
intravenous administration were not significantly different from those of the commercial 
formulation, lysine clonixinate. The results of the study suggested that a 
microemulsion formulation is an alternative vehicle for clonixic acid. 
Paclilaxel (Taxol) injection is known to cause hypersensitivity reactions. Consequently, 
He et al.72 explored whether it was possible to prepare a non-sensitizing 
paclitaxel microemulsion using egg phosphatidylcholine, Piyronic F68 ancl Cremophor 
EL as surfactants, and ethanol as cosurfactant. Note that there was no 
mention of the presence of a specific oil. The study showed that for an equivalent 
dose, the paclitaxel microemulsion did not cause any hypersensitivity reaction, 
whereas Taxol did. In addition, the bioavailability of the paclitaxel in the new 
microemulsion was significantly higher and the elimination rate slower than that 
achieved with Taxol. The authors suggested that the drug molecules, trapped in the 
oil droplets, diffused into the systemic circulation slowly. Furthermore, the small 
particle size of the droplets (10-50 nm) meant that the microemulsion droplets could 
escape from uptake and phagocytosis of RES. Infact it was previously suggested 
that it should be possible to modify the surface of the microemulsion droplets, with 
polyoxyethylene chains, to significantly improve circulation time.57'118'190 
Kanga et al.S6 have recently explored the possibility of optimizing the release 
of paclitaxel from a SEEDS formulation using the polymer, PLGA. The SEEDS formulation, 
which was a mixture of drug, tetraglycol, Cremophor ELP, and Labrafil 
1944 also contained PLGA of varying molecular weight. The droplet size of the 
microemulsions was in the range of 45-270 nm, with the systems without PLGA 
exhibiting the smaller size. The release rate of paclitaxel decreased in the order of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 147 
PLGA, PLGA 8 K, PLGA 33 K, and PLGA 90 Kg/mol, suggesting that the molecular 
weight of PLGA in microemulsion could control the release rate of paclitaxel from 
4.3.1. Long circulating microemulsions 
Long circulating microemulsions have been suggested as an alternative formulation 
to long circulating vesicles on the basis of their small size, thus avoiding uptake by 
the RES, their stability and their ability to solubilize lipophilic compounds more 
effectively than vesicles, and their ease of preparation. 
Wang et al.,S3 and Junping et al.,m have determined the potential of intravenous 
delivery systems of emulsion/microemulsion systems based on vitamin E, cholesterol 
and PEG2ooo-lipid. In their first study, Wang et a/.,183 prepared emulsions containing 
1 part drug, 3 parts vitamin E, 3 parts cholesterol and 3 parts PEG2000-DSPE 
with the final formulation containing 5mg of drug in 10 mL of saline solution. 
Although the emulsion was reported to form spontaneously on the addition of the 
required amount of saline, the formulation was homogenized to produce a more 
uniform particle size distribution of 123.0 1.2 nm; no information was given as to 
the size of the droplets prior to homogenization. The zeta potential and drug loading 
efficiency of the sub-micron emulsion were -12.67 + 1.35 mv and 96.3 + 0.3. 
Although the size and loading efficiency of the formulation remained uncharged 
when stored at 7 to 8C for a year, ~6.5% decomposition of the drug was observed. 
The plasma area under the curve (AUC) of the drug in the sub-micron emulsion 
was significantly greater than that of free drug. Overall, the drug in the emulsion 
had a lower acute toxicity and greater potential antitumor effects than the free drug, 
suggesting that the formulation is a useful tumor-targeting sub-micron emulsion 
drug delivery system. 
In a follow-up study, Junping et al.84 prepared microemulsions of vincristine 
suitable for injection using vitamin E, PEG2000-DSPE and cholesterol, adding oleic 
acid to it. The weight ratio of components used was I part drug, 5 parts oleic acid, 
5 parts vitamin E, 5 parts cholesterol and 5 parts PEG2000-DSPE. No homogenization 
was used in the preparation of the microemulsion which yielded microemulsion 
droplets of 138.1  1.2 nm, when prepared using saline at pH 7.4. Note that 10 mL 
of microemulsion solution contained 1 mg of drug. The adjustment the pH of the 
aqueous phase pH and the presence of oleic acid was essential for a high drug 
loading (94.3  0.3%), while the vitamin E was required for long-term storage of 
the formulation at 7 to 8C. The formulation was stable, with respect to particle 
size, when stored at 78C in the dark for 1 year, while the loading efficiency of 
drug decreased by approximately 3%, and 7.4% decomposition of the drug was 
observed. The plasma AUC of the vincristine in the microemulsion was significantly 
148 Lawrence & Warisnoicharoen 
greater than that of free drug. As with the previous formulation, the drug in the 
microemulsion exhibited a low acute toxicity and a high potential antitumor effect. 
4.3.2. Targeted delivery 
Shiokawa et al.M recently reported the development of a novel, tumor targeted 
microemulsion formulation suitable for delivery of the lipophilic antitumor antibiotic, 
aclacinomycin A. Tumor targeting was achieved via folate linked to the exterior 
surface of long circulating (pegylated) microemulsions. Folate was selected 
because the folate receptor is abundantly expressed in a large percentage of human 
tumors, but it is only minimally distributed in normal tissues. The basic composition 
of the microemulsion was PEG2ooo-DSPE/cholesterol/vitarnin E/drug 
(present at a 3:3:3:1 weight ratio or 7:48.3:43,3:1.5 molar ratio). In one microemulsion, 
0.24 mol% of folate linked PEG2000-DSPE was present, another contained 0.24 mol% 
of folate linked PEG5000-DSPE. In a third, the folate was linked directly to the 
DSPC and in the final one, no folate was present. The association of the folate- 
PEGsooo-linked microemulsion and folate-PEGaooo-lhiked microemulsion with the 
target cells was 200-and 4-fold higher, whereas their cytotoxicity was 90- and 3.5- 
fold higher than those of nonfolate microemulsion respectively. The folate-PEGsooolinked 
microemulsions showed 2.6-fold higher accumulation in solid tumors 24 hrs 
after i.v. injection and greater tumor growth inhibition than free drug. These findings 
suggest that a folate-linked microemulsion is a feasible means for tumortargeted 
delivery of lipophilic drug. This study shows that folate modification with 
a sufficiently long PEG chain on the exterior of a microemulsions is an effective 
way of targeting the carrier to tumor cells. 
4.4. Topical delivery 
AAA. Dermal and transdermal delivery 
The dermal and transdermal routes of administration offer several advantages compared 
with other routes of administration. However, the poor permeability of the 
stratum corneum often limits the possibilities for choosing the topical administration 
route. Therefore, novel innovative formulations such as microemulsions that 
have the potential to facilitate skin permeation are of great interest. The investigation 
of microemulsions as vehicles for cutaneous drug delivery is increasingly 
common as their potential is realized. Indeed, the cutaneous route is currently the 
most popular route of adminstration for a microemulsion. Microemulsions offer 
significant potentials as transdermal delivery vehicles, since they are robust, frequently 
stable to the addition of significant amounts of soluble enhancers, excipients 
and depending on their molecular architecture. Kreilgaard has reviewed the use of 
Recent Advances in Microemulsions as Drug Delivery Vehicles 149 
microemulsions as cutaneous drug delivery vehicles in 2002. In the present review, 
work prior to 2002 will not be dealt with in any detail. In addition, due to the large 
amount of research in the area, the review is not exhaustive. 
Proteins and peptides 
Recently, the transdermal route has received attention as a promising means to 
enhance the delivery of drug molecules, particularly peptides, across the skin, using 
harsh physical enhancement techniques such as iontophoresis and sonophoresis. 
Very little research has been performed, investigating microemulsions as vehicles 
for peptide delivery. Getie et al.66 examined the skin penetration profiles of 0.75 wt% 
desmopressin acetate released from a water-in-oil microemulsion comprising 5 wt% 
water, 20wt% Tagot 02:Span 80 3:2 and 74.25 wt% isopropyl myristate. However, 
the profile was comparable to that obtained using a standard amphiphilic cream. 
Although the amount of drug that penetrated the upper layers of the skin was 
significantly higher from the cream than from the microemulsion at all time intervals, 
within 6 hrs 6% of the applied dose reached the acceptor compartment from 
the microemulsion instead of 2% from the cream within 300 min, suggesting that 
the water-in-oil microemulsion has potential for the systemic administration of the 
Hydrophilic drugs 
Water-in-oil microemulsions have been used to enhance the penetration of watersoluble 
drugs. For example, Alvarez-Figueroa and Blanco-Mendez9 reported the 
in vitro delivery of water-soluble methotrexate from hydrogels using iontophoresis, 
and passively from oil-in-water and water-in-oil microemulsions prepared using 
either a 3:1 v:v Labrasol: Plurol Isostearique mixture or a 3:1:1.2 v:v:v Tween 80:Span 
80:l,2-octanediol mixture as surfactant/cosurfactant, and either ethyl oleate or isopropyl 
myristate as oil. All microemulsion formulations studied were more effective 
than passive delivery from aqueous solution of the hydrophilic drug, although for 
the microemulsions, delivery was greater from the oil-in-water systems. However, 
delivery from the microemulsions was less than that using iontophoresis, probably 
because of the lower solubility of drug in microemulsions than in simple aqueous 
Escribano et al.53 attempted to improve the transdermal permeation of sodium 
diclofenac. Four formulations were studied. One was an oil-in-water microemulsion 
based on transcutol (19wt%), plurol oleique (19.5 wt%), water (30.6 wt%), 
isostearyl isostearate (10.9 wt%) and Labrasol (19wt%). The other three formulations 
were "co-solvent" systems prepared from various of the ingredients used for 
150 Lawrence & Warisnoicharoen 
the microemulsion formulation. In this study, the microemulsion performed less 
well than the various co-solvent formulations and in a similar manner to an aqueous 
solution of the drug. This observation is perhaps not surprising as various 
enhancers were involved in the microemulsion droplets and were not available to 
improve drug penetration. Also, as it is likely that the drug was predominately in 
the continuous phase of the microemulsion, it is not surprising that the formulation 
behaved in a similar manner to an aqueous solution. 
The in vitro transdermal permeation of the antineoplastic, 5-fluorouracil, 
incorporated at I.25mg/mL in water-in-oil microemulsions prepared using 
AOT/water/isopropylmyristate has been studied by Gupta et al.69 These 
researchers found that as the water content increased from 0.9, 1.8, 2.7 and 3.6% 
w/w, microemulsions prepared with a surfactant to oil ratio of 5:95 showed 1.68, 
2.36, 3.58 and 3.77-fold increases respectively in the skin flux of 5-fluorouracil, 
compared with an aqueous solution of drug. Increasing the surfactant: oil weight 
ratio from 5:95 through 9:91 to 13:87, at fixed water:surfactant content of 15, gave 
3.58-, 5.04- and 6.3-fold enhancements of drug flux. In their study69 used attenuated 
total reflectance-Fourier transform infrared spectroscopy to determine that the 
microemulsions exerted their enhancement by interacting and perturbing the architecture 
of the statun corneum. The extent of this perturbation was dependent upon 
the concentrations of water and AOT in the microemulsion. Preliminary toxicity 
studies suggested that the microemulsions were a suitable vehicle for transdermal 
Amphiphilic drugs 
Jurkovic et al.85 have investigated the formulation of the amphiphilic antioxidant 
ascorbyl palmitate in a microemulsion, with a view to using the formulation as a 
protectant against free radical formation due to UV irradiation. Both oil-in-water 
and water-in-oil microemulsions were prepared using a medium chain triglyceride 
as oil, and PEG-8 caprylic/capric glycerides (Labrasol) and polyglyceryl-6-dioleate, 
(Plurol oleique) as surfactant and cosurfactant. The ascorbyl palmitate was incorporated 
into the microemulsions at various concentrations between 0.5-5.0 wt%. The 
microemulsions were gelled using either xanthan gum (water-in-oil) or Aerosil 200 
(water-in-oil). The effectiveness of the ascorbyl palmitate in the microemulsions 
depended on both the concentration and type of microemulsion. Regardless of 
the type of microemulsion, efficacy was significantly higher at the higher ascorbyl 
palmitate concentrations. Overall, the oil-in-water microemulsions were more 
effective at protecting against UV irradiation, although they delivered ascorbyl 
palmitate to the skin at a slower rate than the water-in-oil microemulsions. 
The effect of formulation composition on the in vitro release rate of the 
amphiphilic drug, diclofenac diethylamine, from a range of microemulsion vehicles 
Recent Advances in Microemulsions as Drug Delivery Vehicles 151 
containing PEG-8 caprylic/capric glycerides (surfactant), polyglyceryl-6 dioleate 
(cosurfactant), isopropyl myristate and water was determined by Djordjevic.49 The 
phase behavior of the microemulsions was determined in the absence of drug. In the 
microemulsions selected for further study, the level of water present ranged from 10 
to 60 wt% while the amount of oil varied from 8 to 46.6 wt%. The physico-chemical 
characterization studies indicated the microstructure to be either bicontinuous or 
non-spherical, and despite its amphiphilic nature, the drug was partitioned mainly 
in the water phase. The non-linearity of the drug release profile from the bicontinuous 
microemulsions was thought to be due to a complex distribution of drug 
within the microemulsion. The flux of the drug increased by >4 times, from a waterin-
oil to an oil-in-water microemulsion, the release of drug from the bicontinuous 
microemulsion, suggesting that the microstructure hampers the release of the drug. 
Hydrophobic drugs 
Dalmora and Oliveria43 and Dalmora et al.,u investigated the release of piroxicam 
encapsulated in /8-CD in cationic oil-in-water microemulsions, in an attempt to 
optimize the drug's delivery. The results demonstrated the potential of the reservoir 
in vivo system following the use of a microemulsion. The high degree of retention 
of the active substance can provide a means for modulating the anti-inflammatory 
effect, by greatly extending the release period relative to those formulations where 
the piroxicam is only dissolved or dispersed in a homogeneous aqueous medium. In 
conclusion, both microemulsions and ^-CD-containing microemulsions can offer 
many promising features for their use as topical vehicles for piroxicam delivery. 
Some of the microemulsions gelled using carbopol 940. 
Paolino et alP7 examined the potential of oil-in-water microemulsions as topical 
drug vehicles for the percutaneous delivery of ketoprofen. Microemulsions were 
prepared using triglycerides as oil, and were stabilized by a mixture of lecithin and 
n-butanol as a surfactant/ co-surfactant system. The percutaneous enhancer, oleic 
acid, was added to some of the microemulsions. Physicochemical characterization 
of the microemulsions yielded a mean droplet size of 35 nm and a negative zeta 
potential of -19.7 mV in the absence of oleic acid and  39.5 mV in its presence. 
The ketoprofen-loaded microemulsions showed an enhanced permeation through 
excised human skin with respect to conventional formulations, although no significant 
percutaneous enhancer effect was observed in the presence of oleic acid. 
Microemulsions showed a good human skin tolerability on volunteers. 
Shukla et al.165 have investigated the potential of oil-in-water (o/w) microemulsions 
as vehicles for the dermal delivery of a eutectic mixture of lidocaine 
(lignocaine) and prilocaine, which acted as the oil phase. The microemulsion was 
stabilized by a blend of a 2:3 ratio Tween 80 and Poloxamer 331, a mixture of water 
152 Lawrence & Warisnoicharoen 
and propylene glycol were used as the hydrophilic phase. These microemulsions 
were able to solubilize up to 20 wt% of the eutectic mixture. 
In an attempt to enhance the transdermal delivery of the poorly water soluble 
drug, triptolide, and to reduce the toxicity problems associated with its usage, a 
water-in-oil microemulsion was compared with that of solid lipid nanopartides.124 
The microemulsion which was formulated using 40wt% isopropyl myristate, 
50 wt% Tween-80:l,2-propylene glycol (5:1, v/v) and water and contained 0.025 wt% 
of triptolide, gave a steady-state flux (for over 12 hours) and a permeability coefficient 
of triptolide of 6.4  0.7 mg/cm2 per h and 0.0256  0.002 cm/h; a value which 
was approximately double that of the solid liquid nanoparticles and 7 times higher 
than that of triptolide solution of the same concentration. In another study, Chen 
etal.33 also studied the incorporation of the drug, into a similar microemulsion using 
oleic acid as oil. Oleic acid was added because it is a known penetration enhancer, 
although there was no evidence of it acting as such in the present formulation. The 
addition, however, of 1 wt% menthol to the formulation slightly increased penetration 
from 1.58  0.04 to 2.08  0.06 \ig/cm2 per h (p < 0.05). Encouragingly, no 
obvious skin irritation was observed for the formulation studied, suggesting that 
microemulsions are promising vehicles for the transdermal delivery of triptolide. 
Ross et al.153 examined the transdermal penetration, across full thickness hairless 
mouse skin, of the insect repellant, N,N-diethyl-m-toluamide (DEET), contained 
in either a 1:1 v / v ethanohwater solution (containing 20 wt% DEET) or one 
of two commercially available microemulsion formulations (3M Ultrathon Insect 
Repellant (containing 31.6 wt% DEET; 3M, St. Paul, MN), and Sawyer Controlled 
Release DEET Formula (19.0%; Sawyer Products, Safety Harbor, FL). Both formulations 
were of interest because they were marketed as retarding the absorption of 
DEET due to being microemulsions. All of the DEET preparations exhibited considerable 
penetration, e.g., the ethanolic DEET formulation had a time to detection 
of approximately 30 min with steady stale at 85 min. The penetration obtained with 
the Sawyer was no different from that obtained from the ethanolic solution. The 
other microemulsion formulation (3M) demonstrated a different profile; despite 
being a higher concentration of DEET (30wt% versus 20wt%) and a comparable 
time to detection (40 min), the time to reach steady state was delayed, although 
there was still substantial absorption at steady state. 
Sintov and Shapiro168 prepared a high surfactant lidocaine microemulsion, containing 
as surfactant a mixture of glyceryl oleate and either PEG-40 stearate or 
PEG-40 hydrogenated castor oil, isopropyl myristate as oil, tetraglycol as cosurfactant, 
water, and up to 10wt% of drug, although 2.5 wt% was generally used. 
The microstructure of the microemulsion went from oil-in-water, through bicontinuous 
to water-in-oil. The penetration of the drug from the various formulations 
showed that the surfactant mixture containing PEG-40 stearate was best, while the 
Recent Advances in Microemulsions as Drug Delivery Vehicles 153 
water and surfactant/cosurfactant concentration was also important. Significantly, 
the lag time for penetration was reduced, suggesting that these microemulsions 
loaded with drug would provide rapid local analgesia. 
Priano et tzl.U7 investigated the delivery from a water-in-oil microemulsion, of 
apomorphine present as ion-pair complex with octanoate to increase its lipophilicity 
and to diminish its dissociation. As the drug was present at a high concentration, 
the dispersed phase acted as a reservoir, making it possible to maintain an almost 
constant concentration in the continuous phase and therefore achieving pseudozero-
order release kinetics. The composition of the microemulsion was complex, 
containing 18.2 wt% water, 42.1 wt% of oily phase of isopropyl miristate-decanol 
1:1.5 v/v, 3.9 wt% R-apomorphine hydrochloride, 7.3 wt% Epikuron 200, 7.1 wt% 
benzyl alcohol, 4.6 wt% octanoic acid 3.5 wt% sodium octanoate, 5.7 wt% sodium 
taurocholate, 7.6 wt% 1,2-propanediol. The microemulsion was thickened by the 
addition of 5.9 wt% Aerosil 2000. The microemulsion was able to provide in vitro, 
through hairless mouse skin, a flux of 88g/h per cm2 for 24hrs, with a kinetic 
release of pseudo-zero-order, and was chosen for in vivo study; all the components 
were biocompatible and safe. The flux gave a first approximation of the feasibility 
of the transdermal administration in man. 
The pain and discomfort caused by the injection of local anesthetics has stimulated 
research into developing topical anesthetics. However, the currently available 
formulations, such as Ametopgel, (4 wt% amethocaine base preparation) have a 
number of disadvantages, in particular a long delay of typically 40-60 min between 
application and anesthetic effect and the requirement for a plastic occlusive dressing. 
Arevalo et alP have recently developed a decane-in-water microemulsion stabilized 
by lauromacrogol 300 and containing 4 wt% of amethocaine in an attempt 
to achieve faster drug permeation, thus reducing the time to reach optimum anesthetic 
effect. The amethocaine microemulsion proved to be a promising fast-acting 
analgesic in experimental preclinical studies. 
Mixtures of hydrophilic and hydrophobic drugs 
Although microemulsions have long been suggested as suitable formulations for 
the co-adminstration of drugs of very varying physico-chemical properties, it is only 
very recently that anyone has reported doing so. Lee et al.lw have developed a novel 
microemulsion enhancer formulation for the co-administration of hydrophilic (lidocaine 
HC1, diltiazem HC1) and lipophilic (lidocaine free base, estradiol) drugs. The 
microemulsions composed of isopropyl myristate and water, and were stabilized by 
the nonionic surfactant, Tween 80. Transdermal enhancers such as w-methyl pyrrolidone 
(NMP) and oleyl alcohol were incorporated into all systems without apparent 
disruption of the system. Unfortunately, the authors did not give the precise, 
154 Lawrence & Warisnoicharoen 
composition of the microemulsions tested; it was only mentioned that they contained 
a 1:1 v:v mixture of water and ethanol, isopropylmyristate as oil and Tween 
80 as surfactant, and were either oil-in-water or water-in-oil. Interestingly, regardless 
of the physico-chemical nature of the drug, the oil-in-water microemulsions 
provided significantly better flux for all drugs studied (p < 0.025). Enhancement 
of drug permeability from the oil-in-water systems was 17-fold for lidocaine base, 
30-fold for lidocaine HC1,58-fold for estradiol, and 520-fold for diltiazem HC1. Significantly, 
the simultaneous delivery of estradiol with diltiazem hydrochloride did 
not affect the transport of either drug (p > 0.5). 
Traditionally, vaccines have been administered by injection using needles, although 
the concept of topical immunization through intact skin has attracted much attention. 
Cui et al.42 recently hypothesized that a fluorocarbon-based microemulsion 
system could be one possible way to deliver plasmid DNA across the skin. 
Cui et al.42 screened a range of fluorosurfactants for their ability to form ethanolin-
perfluorooctyl bromide microemulsions. Note that the authors provided no 
evidence of a microemulsion being formed. The stability of plasmid DNA in the 
formulations was also examined. From the surfactant screen, the commercially 
available Zonyl FSN-100, an ethoxylated nonionic fluorosurfactant, was selected 
for further study. Significant enhancements in luciferase expression and antibody 
and T-helper type-1 based immune responses, relative to those of "naked" pDNA 
in saline or ethanol, were observed after topical application of plasmid DNA in 
ethanol-in-perfluorooctyl bromide microemulsion system. From these studies, it 
can be concluded that fluorocarbon-based microemulsions are suitable for DNA 
vaccine delivery, although the mechanism(s) of the immune response induction is 
not known. It is possible that the transport of the molecules across the skin is via the 
hair-follicles, because DNA is too large and highly charged to cross intact stratum 
4.5. Ophthalmic 
Conventional ophthalmic dosage forms tend to be either simple solutions of watersoluble 
drugs or suspension or ointment formulations of water-insoluble drugs. 
Unfortunately, as these delivery vehicles generally result in poor levels of drug 
absorption across the cornea, most of the applied drug does not reach its intended 
site of action. However, because of the relative safety and convenience of topical 
application in ophthalmology, as well as the relatively low risk (compared 
with other routes of administration) of systemic side-effects, topical administration 
Recent Advances in Microemulsions as Drug Delivery Vehicles 1 55 
of ophthalmic agents is the preferred route of delivery. Microemulsions and submicroemulsions 
should offer a possible solution to the problem of poor delivery 
to the cornea, by sustaining the release of the drug, as well as by providing a 
higher penetration of drug into the deeper layers of the eye. In addition, they offer 
the potential of increasing the solubility of the drug in the ophthalmic delivery 
Gallarate et al.6} were probably the first to examine the potential of microemulsions 
as vehicles for ophthalmic delivery. Since then, a number of groups have 
successfully demonstrated the ability of microemulsions (sub-microemulsions) to 
prolong the ocular delivery of drug. In their study, Gallarte et al.a were able to 
further prolong the release of timolol by forming an ion pair with octanoic acid. 
Garty and Lusky63 demonstrated that the delivery of pilocarpine from an oil-inwater 
microemulsion was delayed to such an extent that the instillations of the 
microemulsion formulation twice daily were equivalent to four times daily the 
applications of conventional eye drops. A similar result was reported by Muchtar 
et alP who determined in vitro that the corneal penetration of indomethacin formulated 
in a sub-microemulsion was more than three times that obtained using 
commercially available drops. A number of researcher have investigated the potential 
of positively charged microemulsions to retain the delivery vehicle in the eye, 
thereby sustaining delivery23,52 
To date, a range of drugs have been formulated in a microemulsion in an attempt 
to sustain release including adaprolol maleate,11'125 timolol,61 levobunolol,62 
chloramphenicol162 tepoxalin,54 piroxicam,100 delta-8-tetrahydrocannabinol,129 
pilocarpine,21'52'63'71,133 indomethacin,130 antibodies20 and dietary iso-flavonoids 
and flavonoids.83 In general, these studies showed that it was possible to delay the 
effect of drug incorporated in a microemulsion, thereby improving bioavailability. 
The proposed mechanism of the delayed action is that microemulsion droplets are 
not eliminated by the lachrymal drainage, thereby acting as drug reservoirs. The 
first studies conducted on man with microemulsions containing adaprolol maleate 
and pilocarpine, confirmed the results of the earlier studies performed mainly using 
rabbits.18,178 Vandamme178 has recently reviewed the use of microemulsions as ocular 
delivery system, and thus only studies since then will be considered in the 
present review. 
Fialho and da Silva-Cunha58 recently investigated the long term application 
of a microemulsion system in rabbits intended for the topical ocular administration 
of dexamethasone. The formulation contained 5 wt% isopropyl myristate as 
oil, 15 wt% Cremophor EL as surfactant, and a polar phase of water and 15 wt% 
propylene glycol, with dexamethasone present at a concentration of 0.1 wt%. 
Significantly, ocular irritation tests in rabbits suggested that the microemulsion did 
not provide significant alteration to eyelids, conjunctiva, cornea and iris over a Fe 
156 Lawrence & Warisnoicharoen 
3-month period. In addition, the formulation exhibited greater penetration of dexamethasone 
in the anterior segment of the eye and longer release of the drug when 
compared with a conventional preparation. The area under the curve obtained 
for the microemulsion system was more than two-fold that of the conventional 
preparation (p < 0.05). 
Gulsen and Chauhan68 have recently developed a disposable soft contact lens of 
a drug-containing microemulsion dispersed in a poly 2-hydroxyethyl methacrylate 
(HEMA) hydrogel, suitable for ophthalmic delivery, in an attempt to reduce drug 
loss and side-effects. Upon insertion into the eye, the lens will slowly release the 
drug into the pre lens (the film between the air and the lens) and the post-lens (the 
film between the cornea and the lens) tear films, thus providing a sustained delivery 
of drug. Assuming the size and drug loading of the microemulsions is low, the lenses 
should be transparent. It was found using these microemulsion-containing lenses, 
with and without a stabilizing silica shell, that drug could be released for a period 
of >8 days. By altering droplet size and loading, it is possible to tailor release. 
4.6. Vaginal 
In their 2001 review, D'Cruz and Uckun proposed that microemulsion gel formulations 
had great potential as intra vaginal/ rectal drug delivery vehicles for lipophilic 
drugs, such as microbicides, steroids, and hormones, because of their high drug 
solubilization capacity, increased absorption, and improved clinical potency, as 
long as a non toxic formulation could be prepared. In their review, D'Cruz and 
Uckun reported the formulation of two microemulsion-based gels using commonally 
available pharmaceutical excipients. Repeated intravaginal applications of formulations 
to rabbits and mice were found to be safe and did not cause local, 
systemic, or reproductive toxicity. D'Cruz and Uckun investigated the potential 
of the microemulsion-based gels as delivery vehicles of two lipophilic drugs, WHI- 
05 and WHI-07, which exhibit potent anti-HIV and contraceptive activity. As AIDS 
is spread largely through sexual intercourse, the development of a dual action 
vaginal spermicidal microbicide to curb mucosal viral transmission, as well as to 
provide fertility control would have a tremendous impact world wide. D'Cruz 
and Uckun46"48 investigated the formulation of 2 wt% of the lipophilic drugs in a 
microemulsion-based gel, composed of Phospholipon 90G and Captex 300 as the 
oil phase, with Pluronic F68 and Cremophor EL as surfactants, and seaspan carragennan 
and Xantral as gelling agents. The microemulsions were gelled to obtain the 
necessary viscosity for the gel-microemulsion formulation. Under the conditions 
of their intended use, intravaginal application of the gel-microemulsions containing 
2 wt% of drug in a rabbit model resulted in marked contraceptive activity, as 
well as exhibiting a lack of toxicity. Therefore, as a result of its dual anti-HIV and 
Recent Advances in Microemulsions as Drug Delivery Vehicles 157 
spermicidal activities, the drug-containing gels shows unique clinical potential as 
a vaginal prophylactic contraceptive for women who are at a high risk of acquiring 
HIV by heterosexual transmission. 
4.7. Nasal 
Nasal route has been demonstrated as being a possible alternative to the intravenous 
route for the systemic delivery of drugs. In addition to rapid absorption and 
avoidance of hepatic first-pass metabolism, the nasal route allows the preferential 
delivery of drug to the brain via the olfactory region, and is therefore a promising 
approach for the rapid-onset delivery of CMS medications. The solution-like feature 
of microemulsions could provide advantages over emulsions in terms of the 
sprayability, dose uniformity and formulation physical stability. 
Li et al.lu developed a diazepam-containing ethyl laurate-in-water microemulsion, 
stabilized by Tween 80 and containing propylene glycol and ethanol as cosolvents 
for the rapid-onset intranasal delivery of diazepam. A single isotropic region, 
which was considered to be a bicontinuous microemulsion, was seen at high surfactant 
concentrations but at various Tween 80: propylene glycol: ethanol ratios. 
Increasing Tween 80 concentration increased the microemulsion area, microemulsion 
viscosity, and the amount of water and oil solubilized. In contrast, increasing 
ethanol concentration produced the opposite effect. A microemulsion consisting of 
15 wt% ethyl laurate, 15 wt% water and 70 wt% Tween 80:propylene glycohethanol 
at a 1:1:1 weight ratio contained 41 mg/mL of the poorly-water soluble diazepam. 
The nasal absorption of diazepam from the formulation was fairly rapid with a maximum 
drug plasma concentration being obtained within 2 to 3 min, while bioavailability 
at 2hrs post-administration was ~50% of that obtained with intravenous 
Zhang et al.192 attempted to prepare an oil-in-water microemulsion, containing a 
high concentration of nimodipine, suitable for brain uptake via the intranasal route 
of delivery. Three microemulsion systems stabilized by either Cremophor RH 40 or 
Labrasol, and containing a variety of oils, namely isopropyl myristate, Labrafil M 
1944CS and Maisine 35-1, were developed and characterized. The nasal absorption 
of the drug from the three microemulsions was studied in rats. The formulation composed 
of 8 wt% Labrafil M 1944CS, 30 wt% Cremophor RH 40/ethanol (3:1 weight 
ratio) and water solubilized up to 6.4 mg/mL of drug and exhibited no ciliotoxicity. 
After intranasal administration, the peak plasma concentration was obtained 
of 1 hr, while the absolute bioavailability was ~32%. Significantly, uptake of the 
drug in the olfactory bulb after nasal administration was three times that which 
was obtained from intravenous injection. In addition, the ratios of the AUC in brain 
tissues and cerebrospinal fluid to that in plasma obtained after nasal administration 
1 58 Lawrence & Warisnoicharoen 
were significantly higher than those seen after administration. In conclusion, the 
microemulsion system appears to be a promising approach for the intranasal delivery 
of nimodipine. 
Richter and Keipert51 investigated the in vitro permeability of the highly 
lipophilic material, androstenedione, across excised bovine nasal mucosa, porcine 
cornea and an artificial cellulose membrane. In order to control release, the 
two microemulsion formulations studied contained either hydroxypropyl-yScyclodextrin 
or propylene glycol. Both microemulsions were prepared from 5 wt% 
isopropyl myristate, 20 wt% Cremophor EL and water. The permeation of the drug 
through the three tissues was influenced by the microemulsion. For example, the 
apparent permeability coefficients (Papp) of the drug from the microemulsions 
across nasal mucosa did not differ from the Papp of the drug contained in solution. 
In the case of the other two membranes, release from both of the microemulsion formulations 
exhibited extended time lags, so no Papp could be calculated. It seems that 
the composition of the microemulsion had a greater impact on the Papp of cornea 
than on the Papp of the other tissues. The structure of the different membranes is 
probably responsible for the observed differences in permeation. 
4.8. Pulmonary 
Emulsions and (to a far lesser extent) microemulsions have been investigated as 
vehicles for pulmonary delivery. By far, the most widely studied systems are those 
containing fluorocarbon oil and are stabilized by a (predominately) fluorinated 
surfactant. Fluorocarbon oils are of pharmaceutical interest because of their biological 
inertness and their high (and unique) ability to dissolve gas, which means 
they can support the exchange of the respiratory gases in the lungs. In addition, 
a fluorocarbon oil, namely perfluorooctylbromide, is in Phase 11:111 clinical trials 
in the United States, for the treatment of acute respiratory distress by liquid ventilation. 
It should be noted that en-large hydrocarbon surfactants are ineffective 
solubilizers in fluorocarbon-based systems. Instead, fluorocarbon surfactants are 
required. To date, fluorocarbon-based (micro)emulsions have been investigated 
for use as oil-in-water systems for in vivo oxygen delivery (blood substitutes), 
targeted systems for diagnosis and therapy, and water-in-fluorocarbon systems 
for pulmonary drug delivery.40'102 Water-in-perfluorooctylbromide microemulsions 
have been shown to deliver homogeneous and reproducible doses of a tracer (caffeine) 
using metered-dose inhalers (pMDI) pressurized with hydrofluoroalkanes 
Lecithin-based reverse microemulsions have also been investigated as a means 
of pulmonary drug delivery.170'171 In these studies, dimethylethyleneglycol (DMEG) 
and hexane were used as models for the propellants, dimethyl ether (DME) and 
Recent Advances in Microemulsions as Drug Delivery Vehicles 159 
propane respectively. A combination of equilibrium analysis and component diffusion 
rate determination (by pulsed-field gradient [PFG]-NMR) and iodine solubilization 
experiments were used to confirm the formation of a microemulsion. 
Water soluble solutes, including selected peptides and fluorescently labeled polya
6-[N-(2-hydroxyethyl) D,L-aspartamide] were dissolved in the microemulsions in 
a lecithin- and water-dependent manner. Experiments with DME/lecithin demonstrated 
microemulsion characteristics similar to those in the model propellant and 
produced a droplet size and a fine particle fraction suitable for pulmonary drug 
Patel et al.uo have prepared water-in-hydrofluorocarbon (specifically 134a) 
microemulsions using a combination of fluorinated polyoxyethylene ether surfactants 
and a short chain hydrocarbon alcohol such as ethanol. In the absence 
of a hydrocarbon alcohol, only cosolvent systems, but not microemulsions, were 
formed. Due to the high molecular weight of the fluorocarbon surfactant, large 
concentrations of fluorocarbon surfactant are required to solubilize relatively small 
amounts of water compared with comparable hydrocarbon-based surfactants. This 
has obvious implications for the pharmaceutical application of such systems. 
To date, very little on the potential of oil-in-water microemulsions for pulmonary 
drug delivery has been investigated, yet they are attractive vehicles because 
of their ability to solubilize high amounts of drug.157 
4.8.1. Antibacterials 
Al-Adham et al.6 demonstrated that microemulsion formulations have a significant 
antimicrobial action against planktonic populations of both Pseudomonas aeruginosa 
and Staphylococcus aureus (i.e. greater than a 6 log cycle loss in viability 
over a period as short as 60s). Transmission electron microscopy studies indicated 
that this activity may in part be due to significant losses in outer membrane 
structural integrity. Nevertheless, these results have implications for the potential 
use of microemulsions as antimicrobial agents against this normally intransigent 
More recently, the same group6 have determined the antibiofilm activity of 
an oil-in-water microemulsion, prepared from 15wt% Tween 80, 6wt% pentanol 
and 3wt% ethyl oleate, by incubating the microemulsion with an established 
biofilm culture of Ps. aeruginosa PA01 for a period of 4hrs. The planktonic MIC 
of sodium pyrithione and the planktonic and biofilm MICs of cetrimide were 
used as positive controls and a biofilm was exposed to a volume of normal sterile 
saline as a treatment (negative) control. The results showed that exposure to 
the microemulsion resulted in a three log-cycle reduction in biofilm viability, as 
compared to a one long-cycle reduction in viability observed with the positive 
1 60 Lawrence & Warisnoicharoen 
control treatments, suggesting that microemulsions are highly effective antibiofilm 
5. Conclusion 
As can be seen, microemulsions are attractive d r u g delivery vehicles that offer much 
scope for improving drug delivery. Although microemulsions have been seriously 
studied as a delivery vehicle in the last >20 years, there are few microemulsion 
products currently on the market. Comparing microemulsions with vesicular drug 
delivery systems, it is pertinent to note that it took >25 years before vesicles were 
commercially exploited as drug delivery vehicles, and this was with the immense 
research effort expended in their study. Microemulsions have by contrast been much 
less widely studied. It is only a matter of time before more microemulsion-based 
formulations appear on the market. 
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Lipoproteins as Pharmaceutical Carriers 
Suwen Liu, Shining Wang and D. Robert Lu 
1. Introduction 
Large protein structures (in nanometer range) may be utilized as pharmaceutical 
carriers of drugs and DNA for targeted and other specialized delivery in biological 
systems. Lipoproteins are such structures which function as natural biological carriers 
and transport various types of lipids in blood circulation. There are many 
studies suggesting that lipoproteins can serve as efficient carriers for anticancer 
drugs, gene or other type of compounds.1-4 Previous results showed that hydrophobic 
cytotoxic drugs could be incorporated into lipoproteins, without changing the 
integrity of native lipoprotein structure. Lipoproteins as drug carriers offer several 
advantages.5-6 Firstly, they are endogenous components and do not trigger 
immunological response. They have a relatively long half-life in the circulation. Secondly, 
they have small particle size in the nanometer range, allowing the diffusion 
from vascular to extravascular compartments. Thirdly, lipoproteins can potentially 
serve as the carriers for targeted drug delivery through specific cellular receptors. 
For example, low density lipoprotein (LDL)-drug complexes may target cancer 
cells which, in many cases, have higher LDL-receptor expression than normal cells. 
Fourthly, the lipid core of lipoprotein provides a suitable compartment for carrying 
hydrophobic drugs. 
As a result of these advantages, lipoproteins have received wide attentions in 
recent years in the development of drug-targeting strategies to use them as specialized 
delivery vehicles. This review intends to provide an overview of the development 
and the specialized utilization of lipoproteins for drug delivery purpose. After 
1 74 Liu, Wang & Lu 
briefly introducing the structure and the basic biological functions of lipoproteins, 
we will focus on four classes of lipoproteins, namely, chylomicron, very low-density 
lipoprotein (VLDL), low-density lipoprotein (LDL), and high-density lipoprotein 
(HDL), as the carriers for various drug compounds. Cholesterol-rich emulsions 
(LDE) and artificial lipoproteins as drug carriers will also be discussed. 
2. The Structure of Lipoproteins 
Lipoproteins, as implied by their names, are biological protein-lipid complexes. 
Lipoproteins serve the functions of carrying hydrophobic substances in blood circulation 
and transporting them to various biological sites through the protein-receptor 
interactions.6,7 The size of lipoproteins is in the nanometer range and they have a 
spherical shape with complex physicochemical properties. Figure 1 illustrates the 
general structure of lipoprotein. The hydrophobic core contains water-insoluble 
substances and is surrounded by a polar shell. The polar shell consists of phospholipids, 
unesterified cholesterol and different types of apolipoproteins, which 
bind to various cellular receptors for specific biological functions. Therefore, based 
on their physicochemical properties, lipoproteins are nanoemulsions with targeting 
functions provided by the apolipoproteins. Owing to the unique structure of 
lipoproteins, they can serve a two-mode function of solubilizing hydrophobic substances, 
including triglycerides and cholesteryl esters, within the nanoemulsion 
core and allow themselves to float in blood circulation. 
Lipoproteins can be classified into five major classes, based on their densities 
from gradient ultracentrifugation experiments. The lipoprotein classification 
includes chylomicron, very low-density lipoprotein (VLDL), intermediate-density 
lipoprotein (IDL), low-density lipoprotein (LDL), and high-density lipoprotein 
(HDL). These classes of lipoproteins have different sizes, different protein to lipid 
ratios and different types of apolipoproteins. In general, chylomicrons act on transporting 
dietary triacylglycerols and cholesterol to the adipose tissue and liver, following 
the absorption of dietary hydrophobic substances from the intestines. Very 
Fig. 1. General structure of lipoproteins. 
Lipoproteins as Pharmaceutical Carriers 1 75 
Table 1 Physicochemical properties of lipoproteins. 
Lipoprotein Transport Route Size(nm) Protein (%) Total lipids (%) 
Chylomicron Intestines to Liver 75-1200 1.5-2.5 97-99 
VLDL Liver to tissues 30-80 5-10 90-95 
IDL Liver to tissues 25-35 15-20 80-85 
LDL Liver to tissues 18-25 20-25 75-80 
HDL Tissues to liver 5-12 40-55 45-60 
low density lipoprotein, intermediate density lipoprotein and low density lipoprotein 
work at different stages to transport triacylglycerols and cholesterol from the 
liver to various tissues. High density lipoprotein brings endogenous cholesterol 
from the tissues back to the liver. The general physicochemical properties of lipoproteins 
can be seen in Table 1. 
3. Chylomicron as Pharmaceutical Carrier 
Chylomicrons are assembled in the intestine from the absorbed dietary lipids 
and transported by lymphatic system. Although most of the drugs administered 
orally are absorbed directly into the portal blood to reach the systemic circulation, 
an alternative absorption route through the intestinal lymphatics may be available 
for hydrophobic drugs. It is estimated that a high hydrophobicity (log o / w 
partition co-efficient > 5) of drug molecules is required for intestinal lymphatic 
transport.8 Chylomicrons can thus potentially serve as an important natural carrier 
for hydrophobic drugs to transport through lymphatic system.9 It is known 
that targeted drug delivery through the lymphatics is important for anti-viral drug 
molecules for the protection of B- and T-lymphocytes, which maintain relatively 
higher concentrations through the lymphatics than the systemic circulation. Chylomicrons 
have a much larger size than other lipoproteins, and thus can carry more 
drug molecules from the absorption site. With the presence of food, chylomicrons 
are the predominant lipoprotein produced by the small intestine to carry dietary 
lipids efficiently because of its large size. 
Various types of bioactive molecules have been incorporated into reconstituted 
chylomicron structure for delivery purposes. In gene delivery, Hara et a/.10,11 developed 
reconstituted chylomicron which incorporated a hydrophobic DNA complex 
and used it as an in vivo gene transfer vector. They found that the DNA-incorporated 
chylomicrons induced a high gene expression in mouse liver after the reconstituted 
chylomicron was administered through portal vain injection. Furthermore, 
it was also reported that artificial, protein-free lipid emulsions could be utilized to 
model the metabolism of lymph chylomicron in rats, not only in the initial partial 
176 Liu, Wang & Lu 
hydrolysis by lipoprotein lipase, but also in the delivery of a remnant-like particle 
to the liver.12 As a targeted therapeutic approach to hepatitis B, anti-viral iododeocyuridine 
was incorporated into recombinant chylomicrons, resulting in the drug 
molecules being selectively targeted to the liver parenchymal cells.13 It has been 
suggested that chylomicron can serve as a special carrier for liver cell targeting.14 
Due to the targetability, this approach could be further developed as an effective 
therapy for hepatitis B patients. 
4. VLDL as Pharmaceutical Carrier 
VLDL particles have a size range of 30-80 nm. They are assembled in the endoplasmic 
reticulum (ER) and matured in Golgi apparatus of hepatocytes before 
secretion.15 After entering into the plasma, VLDL particles are catabolized by a 
series of biochemical actions, including apolipoprotein exchange of apoC-I, apoCII, 
apoC-III, and apoE; lipolysis by triglyceride lipase; and cell-surface receptormediated 
uptake. As lipolysis proceeds, VLDL particles become smaller and are 
eventually converted to IDL. Some of the IDL particles are rapidly taken up by hepatocytes 
via a receptor-mediated mechanism while others undergo further hydrolysis 
before being converted to LDL. The catabolism route of VLDL suggests the 
possibility of using VLDL as a drug carrier for targeted delivery. ApoE is a protein 
ligand present on the surface of VLDL and it is well known that the receptor of 
apoE is overexpressed on some types of cancer cells. Therefore, VLDL can potentially 
serve as an antineoplastic drug carrier. 
As a drug carrier, VLDL is an interesting candidate because it contains a relatively 
small amount of proteins (about 5-10 % protein) and a large amount of triglycerides 
(about 50-65% within the emulsion core) which can be used to solubilize 
hydrophobic substances sufficiently. By mimicking the compositions and structure 
of VLDL, Shawer et al. developed a VLDL-resembling phospholipid nanoemulsion 
system that carried a new anti-tumor boron compound for targeted delivery to cancer 
cells.16 The nanoemulsion demonstrated sufficient capability to solubilize the 
hydrophobic compound. The structure of the phospholipid nanoemulsion was verified 
based on the changes in the molecular surface area and the molecular volume 
of each component of the nanoemulsion when the particle size is changed (from different 
size fractions). If certain molecules are located at the core of nanoemulsion, 
their numbers per overall volume should not be changed when the particle size 
is increased. If certain molecules are located at the surface of nanoemulsion, their 
numbers per overall volume should decrease when particle size is increased. This is 
because the overall surface area decreases when particle size is increased. Similar to 
the natural lipoproteins, it was demonstrated that phospholipid was predominately 
Lipoproteins as Pharmaceutical Carriers 177 
located at the surface and the hydrophobic substances, triolein and cholesteryl 
oleate, were mainly located in the core of the phospholipid nanoemulsion. 
Recently, a similar nanoemulsion formulation was used to encapsulated 
quantum dots (QD) as a new bioimaging carrier.17 Quantum dots (QDs) are 
semiconductor nanocrystals that are emerging as unique fluorescence probes in 
biomedicine.18-21 When manufactured, most of the quantum dots have organic ligand 
coating on their surface and are extremely hydrophobic. The research goal was 
to encapsulate QDs in phospholipid nanoemulsion and to examine the physical 
stability, size distribution and their interactions with cancer cells. It was found that 
CdSe QDs can be efficiently encapsulated in the phospholipid nanoemulsion. The 
QD-encapsulated phospholipid nanoemulsion are stable and interact well with cultured 
cells to deliver the QDs inside the cells for fluorescence imaging.17 In other 
studies, it has been demonstrated that cytotoxic drugs such as 5-fluorouracil (5-FU), 
5-iododeoxyuridine (IudR), doxorubicin (Dox), and vindesine can be effectively 
incorporated into VLDL, and the resultant complexes showed effective cytotoxicity 
to human carcinoma cells.22 
5. LDL as Pharmaceutical Carrier 
LDL (18-25 nm) is not directly synthesized in human body. Instead, most of them 
are formed through the VLDL pathway. LDL is the major circulatory lipoprotein 
for the transport of cholesterol and cholesteryl esters, and it can be internalized by 
cells via LDL receptor-mediated endocytosis. The internalization process of LDL 
has been well characterized and the understanding of the mechanism can potentially 
help the designing of the drug targeting strategy through the LDL receptor 
(Fig. 2). The binding of dephosphorylated adaptor protein to the plasma membrane 
LDL Receptors 
(. ( . X l B l O l f c . H K . * ^ . , . ^ 
Cell , HMGCoA 
\ LDL Receptors 
mug > ^-.t, ^ v f i ' * oo -* 
LDL Binding  Internalization Drug Release ^Regulation 
Fig. 2. LDL receptor pathway and targeted drug delivery. 
1 78 Liu, Wang & Lu 
initiates the formation of coated pits which are covered by the protein clathrin. The 
receptors from the surrounding regions of the plasma membrane shift towards the 
binding site for internalization. Apolipoproteins including apo B-100 and apo E 
are recognized and bound by the LDL receptor on the cell surface to form a complex 
which is internalized into the coated pits. After internalization of the LDL, 
the coated pits are pinched off and within a very short time, they shed off their 
clathrin coating. The internalized LDL particle is transferred to endocytotic vesicles 
or endosomes. Due to the acidic pH within the endosomes, LDL dissociates 
from its receptor. This is followed by the fusion of the endosomes with lysosomes 
which contain hydrolases. The protein component of LDL is broken into free amino 
acids, while the cholesteryl ester component is cleaved by lysosomal lipase. The free 
cholesterol is released and incorporated into the cell membrane. Excess cholesterol 
is re-esterified by the action of acyl-CoA:cholesterol acyltransferase (ACAT). 
Among various lipoproteins, LDL has been widely studied as a drug carrier for 
targeted and other specialized deliveries, because many types of cancer cells show 
elevated expression of LDL receptors than the corresponding normal cells.23-26 
Comparing with chylomicron, VLDL, and IDL, LDL also has a longer serum halflife 
of 2-4 days,27 making it a desirable drug carrier. Low density lipoprotein was 
found to be suitable as carriers for cytotoxic drugs to target cancer cells. LDLdrug 
complexes can be formed through various processes without changing the 
lipoprotein integrity.28-31 
5.1. LDL as anticancer drug carriers 
Doxorubicin (Dox) is widely used in treating different tumors. Its main side effects 
are cadiotoxicity and multidrug resistance, especially during prolonged treatment 
in the patients. LDL has been studied as a target carrier for Dox in nude mice, bearing 
human hepatoma HepG2 cells.32 Both in vitro and in vivo studies indicated that when 
Dox was incorporated into LDL, the multidrug resistance could be circumvented 
and the cardiotoxicity could be reduced as well.33 Kader and Pater22 used VLDL, 
LDL and HDL as carriers to deliver four cytotoxic drugs, 5-fluorouracil (5-FU), 
5-iododeoxyuridine (IUdR), doxorubicin (Dox) and vindesine. They found that 
significant drug loading was achieved in all three classes of lipoproteins, consistent 
with the sizes and hydrophobicity of the drug. Experiments were carried out to 
examine the changes in drug cytotoxicity against HeLa cervical and MCF-7 breast 
carcinoma cells, after the incorporation into lipoprotein. The results demonstrated 
that VLDL-drug complex did not affect their IC50 on both HeLa and MCF-7 cell 
lines, when compared with free drugs. However, the IC50 values of LDL- and HDLdrug 
complexes were significantly lower compared with free drugs. Their studies 
further indicated that drugs were incorporated into lipoproteins without disrupting 
Lipoproteins as Pharmaceutical Carriers 1 79 
their integrity; drugs remained in their stable forms inside lipoproteins; and human 
LDL and HDL could be particularly useful in the delivery of antineoplastic drugs. 
5.2. LDL as carriers for other types ofbioactive compounds 
Although LDL has been widely studied as a carrier to deliver anticancer compounds, 
it may also be useful to deliver other types of bioactive compounds. 
LDL may serve as a carrier for site-specific delivery of drugs to atherosclerotic 
lesions.34 When dexamethasone palmitate (DP), a steroidal anti-inflammatory drug, 
was incorporated in LDL, an inhibitory effect of this complex on foam cell formations 
was demonstrated. The study indicated that LDL could potentially carry 
DP to atherosclerotic lesions.34 Fluorophore-labeled LDL was also used for optical 
imaging in tumors diagnosis. For example, carbocyanine dyes can be used 
as near infrared (NIR) optical imaging probes with long wavelength absorption, 
high extinction coefficients and high fluorescence quantum yield. In vitro confocal 
microscopic study and ex vivo low-temperature fluorescent scanning demonstrated 
that carbocynine-labled LDL probes, Dil-LDL, could be selectively delivered to 
B16/HepG2 tumor cells and the corresponding animal tumors via the LDL receptor 
pathway.35 It was also proposed that Dil is located and oriented in the phospholipid 
monolayer when it binds to LDL. 
5.3. LDL for gene delivery 
LDL has also been investigated as gene delivery carriers. Comparing with viral 
gene-delivery vectors and some other types of non-viral gene delivery vectors, 
the LDL system shows certain advantages in transfection efficiency and safety 
considerations.5 Several LDL based gene delivery systems have been reported. 
Kim's group developed a terplex system which comprises LDL, lipidized poly(Llysine) 
and plasmid DNA. The complex had a diameter of about 100 nm. The studies 
showed high efficiency to deliver plasmid DNA to smooth muscle cells and fibroblast 
cells.36,37 In addition, a novel LDL-DNA complex was formulated by Khan 
et al.38 LDL was cationized using carbodiimide and the modified lipoprotein complex 
significantly increased the DNA binding capacity with improved stability. The 
novel delivery system also demonstrated the ability to target cells through LDL 
6. HDL as Pharmaceutical Carriers 
Among various lipoproteins, HDL has the smallest size with a diameter of 5-12 nm. 
It shares common structural characteristics with other lipoproteins. However, its 
180 Liu, Wang & Lu 
polar shell contributes more than 80% of the total mass. Newly synthesized HDL 
hardly contains any cholesteryl ester molecules. Cholesteryl esters are gradually 
added to the particles by lecithin via enzymatic reaction: cholesterol acyltransferase 
(LCAT), which is a 59-kD glycoprotein associated with HDL. The interaction of 
HDL with cells appears similar to that of LDL.39 Although the function of HDL in 
the human body is not well-defined, it generally transports excess cholesterol and 
cholesteryl esters from various tissue cells back to the liver. Comparing with other 
types of lipoproteins, small size and fast internalization by tumor cells are the major 
advantages of utilizing HDL for drug delivery and targeting. 
HDL has mainly been utilized for the delivery of water insoluble anticancer 
drugs through the targeting function.40-41 When the anticancer drug, Taxol, was 
incorporated into HDL, stable complexes were formed and they were examined for 
cancer-cell targeting.41 Reconstituted HDL was explored as a drug carrier system for 
a lipophilic prodrug, IDU-OI2.42 The studies indicated that the lipophilic prodrug 
could be efficiently incorporated into reconstituted HDL particles. This approach 
may also be useful to encapsulate other lipophilic derivatives of water-soluble 
drugs. The utilization of HDL for drug targeting may lead to a more effective therapy 
for infectious diseases, such as hepatitis B, since the HDL-drug complexes were 
demonstrated to be selectively taken by parenchymal liver cells.42 Comparing with 
free drugs in cytotoxicity assays, the IC values of HDL-drug complexes were significantly 
decreased, about 2.5 to 23-fold lower.22 Interestingly, it was observed that 
HDL-drug complex specifically increased the cytotoxicity to carcinoma cells. Earlier 
studies showed that HDL could increase the sensitivity of HeLa cells to the 
cytotoxic effects of Dox.43 Similar to LDL-drug complex, the lipoprotein receptor 
pathway appears to be involved in the interactions between HDL-drug complex 
and cancer cells. 
7. Cholesterol-rich Emulsions (LDE) as Pharmaceutical 
LDE is a lipid based formulation, an emulsion with a lipid structure resembling LDL 
particle and it is made without protein incorporation. Essentially, it is composed of 
a cholesteryl ester core surrounded by a monolayer of phospholipids. Comparing 
with native LDL, LDE is removed from the blood circulation more rapidly.44 It 
appears possible that LDE can acquire apoE and other apolipoproteins from native 
lipoproteins in plasma. ApoE can be recognized by LDL receptors, thus allowing the 
binding of LDE to the receptors. However, it is known that LDE binds to receptors 
through apoE, but not through apoBlOO. The interaction between apoE and the 
receptor appears stronger than that of apoBlOO.45 
Lipoproteins as Pharmaceutical Carriers 181 
LDE is considered as a potential carrier for anticancer drugs to deliver 
chemotherapeutic agents to neoplastic cells. Although there is no protein in the 
LDE formulations, previous studies showed that the LDL receptor could still play 
an important role in the cellular uptake of these lipid complexes.46-56 LDE binds 
to low-density lipoprotein receptors which are upregulated in cancer cells, leading 
to a higher concentration in neoplastic tissues.24-57 LDE-carmustine complex was 
studied with a neoplastic cell line and its biodistribution was studied in mice. An 
exploratory clinical study was also conducted. The result showed that the uptake of 
LDE-carmustine complex by tumor was several fold greater than the uptake by the 
corresponding normal tissue. The association of carmustine with LDE preserves the 
tumor-cytotoxicity of carmustine with reduced side effects.58 Preliminary clinical 
study59 was also carried out using LDE-carmustine complex to treat patients with 
advanced cancers. The results demonstrated that the systemic toxicity of the drug 
was significantly reduced. 
Rodrigues et ah investigated the formulation of LDE containing antineoplastic 
compound paclitaxel.55 The experiments revealed a 75% incorporation efficiency 
and the stable complex of the drug molecules incorporated in LDE emulsion. Its 
LD50 was ten-fold greater than that of a commercial formulation of paclitaxel. It was 
suggested by the authors that the cellular uptake and the cytotoxic activity of LDEpaclitaxel 
complex might be mediated by the LDL receptors due to the cholesterol 
moiety in the LDE formulation.55 
In addition to LDE, artificial lipoproteins have been constructed. Several 
research groups have developed various types of artificial lipoproteins.44-60-62 Most 
of them constructed the artificial lipoproteins by incorporating natural apoB protein 
into lipid microemulsion for the purpose of examining the lipoprotein metabolism. 
Artificial lipoproteins containing poly-lysine has also been investigated as the DNA 
carrier for cellular transfection, with the potential to reduce the cytotoxicity and to 
improve the transfection efficiency.63-64 
8. Concluding Remark 
Lipoproteins are natural nanostructures in biological systems. They have unique 
physicochemical properties which may be utilized as pharmaceutical carriers for 
drug compounds and other bioactive substances. Owing to the structural diversity 
of lipoproteins, including chylomicron, VLDL, LDL and HDL, various specialized 
delivery systems may be developed to fully utilize their delivery potentials. New 
nanostructures, such as LDE and artificial lipoproteins, can also be constructed to 
mimic the structure of natural lipoproteins. As these new nanostructures are built 
from scratch, they may be more efficient in encapsulating drug and other bioactive 
molecules, and more effective for specialize drug delivery. 
1 82 Liu, Wang & Lu 
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186 Liu, Wang & Lu 
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culture using an artificial lipoprotein carrier system. Pharm Res 21:676-683. 
Solid Lipid Nanoparticles 
as Drug Carriers 
Karsten Mader 
1. Introduction: History and Concept of SLN 
Nanosized drug delivery systems have been developed to overcome one or several 
of the following problems: (i) low and highly variable drug concentrations 
after peroral administration due to poor absorption, rapid metabolism and elimination 
(ii) poor drug solubility which excludes i.v. injection of an aqueous drug 
solution (iii) drug distribution to other tissues combined with high toxicity (e.g. 
cancer drugs). Several systems, including micelles, liposomes, polymer nanoparticles, 
nanoemulsions and nanocapsules have been developed. During the last few 
years, solid lipid nanodispersions (SLN) have attracted increased attention. It is the 
aim of this chapter to discuss the general features of these systems with respect to 
manufacturing and performance. 
In the past, solid lipids have been mainly used for rectal and dermal applications. 
In the beginning of the 80s, Speiser and coworkers developed solid lipid 
microparticles (by spray drying)1 and "Nanopellets for peroral administration".2 
These Nanopellets were produced by dispersion of melted lipids with high speed 
mixers or ultrasound. The manufacturing process was unable to reduce all particles 
to the submicron size. A considerable amount of microparticles was present 
in the samples. This might not be a serious problem for peroral administration, 
but it excludes an intravenous injection. "Lipospheres", described by Domb, are 
188 Mader 
close related systems.3-5 They are also produced by means of high shear mixing or 
ultrasound and also often contain considerable amounts of microparticles. 
The quality of the SLN has been significant improved by the use of high pressure 
homogenization (HPH) in the early 90s.6-8 Higher shear forces and a better distribution 
of the energy force more effective particle disruption, compared with high shear 
mixing and ultrasound. Dispersions obtained by this HPH are called Solid Lipid 
Nanoparticles (SLN). Most SLN dispersions produced by high pressure homogenization 
(HPH) are characterized by an average particle size below 500 nm and 
a low microparticle content. Other production procedures are based on the use of 
organic solvents HPH/solvent evaporation9 or on dilution of microemulsions.10'11 
The ease and efficacy of manufacturing lead to an increased interest in SLN. 
Furthermore, it has been claimed that SLN combine the advantages yet without 
inheriting the disadvantages of other colloidal carriers.12,13 Proposed advantages 
 Possibility of controlled drug release and drug targeting 
 Increased drug stability 
 High drug pay load 
 Feasibility to incorporate lipophilic and hydrophilic drugs 
 No biotoxicity of the carrier 
 Avoidance of organic solvents 
 No problems with respect to large scale production and sterilization. 
However, during the last years, some of these claims have been questioned and 
it became evident that SLN are rather complex systems which possess not only 
advantages but also serious limitations. 
2. Solid Lipid Nanoparticles (SLN) Ingredients 
and Production 
2.1. General ingredien ts 
General ingredients include solid lipid(s), emulsifier(s) and water. The term lipid 
is used generally in a very broad sense and includes triglycerides (e.g. tristearine, 
hard fat), partial glycerides (e.g. Imwitor), pegylated lipids, fatty acids (stearic acid), 
steroids (e.g. cholesterol) and waxes (e.g. cetylpalmitate). All classes of emulsifiers 
(with respect to charge and molecular weight) have been used to stabilize the lipid 
dispersion. The most frequently used compounds include different kinds of poloxamer, 
polysorbates, lecithin and bile acids. It has been found that the combination 
of emulsifiers might prevent particle agglomeration more efficiently. 
Solid Lipid Nanoparticles as Drug Carriers 189 
Unfortunately, poor attention has been given by most investigators to the 
physicochemical properties of the lipid. Fatty acids, partial glycerides and other 
polar lipids are able to interact with water to much a greater extent, compared 
with a long chain triglyceride (e.g. they might form liquid crystalline phases). Polar 
lipids will have much more interaction with stabilizers (e.g. formation of mixed 
micelles), while more lipophilic lipids will show phase segregation. The author 
strongly suggests to follow the proposal by Small and to classify lipids according 
to their interactions with water.14 
2.2. SLN preparation 
2.2.1. High shear homogenization and ultrasound 
High shear homogenization and ultrasound are dispersion techniques which were 
initially used for the production of solid lipid nanodispersions.1-3 Both methods are 
widespread and easy to handle. However, dispersion quality is often poor due to the 
presence of microparticles. Furthermore, metal contamination has to be considered 
if ultrasound is used. 
Ahlin et al. used a rotor-stator homogenizer to produce SLN from different 
lipids, including trimyristin, tripalmitin, tristearin, partial glycerides 
(WitepsolW35, WitepsolH35) and glycerol tribehenate (Compritol888) by meltemulsification.
15 They investigated the influence of different process parameters, 
including emulsification time, stirring rate and cooling conditions on the particle 
size and the zeta potential. Poloxamer 188 was used as steric stabilizer (0,5%w/w). 
For WitepsolW35 dispersions, the following parameters were found to produce 
the best SLN quality: stirring 8min at 20000rpm, the optimum cooling conditions 
lOmin at 5000 rpm at room temperature. In contrary, the best conditions 
for Dynasan 116 dispersions were 10 min emulsification at 25 000 rpm and 5 min of 
cooling at 5000 rpm in cool water (T = 16C). An increased stirring rate did not significantly 
decrease the particle size, but improved the polydispersity index slightly. 
No general rule can be derived from differences in the established optimum emulsification 
and cooling conditions. In most cases, average particle sizes in the range 
of 100-200 nm were obtained in this study. 
2.3. High pressure homogenization (HPH) 
HPH has emerged as a very reliable and probably the most powerful technique 
for the preparation of SLN. HPH has been used for many years for the production 
of nanoemulsions for parenteral nutrition. In most cases, scaling up represents 
zero or limited problems. High pressure homogenizers push a liquid with high 
pressure (100-2000 bar) through a narrow gap (in the range of few microns). The 
190 Mader 
fluid accelerates on a very short distance to very high velocities. The high shear 
stress disrupts the particles down to the submicron range. Typical lipid contents 
are in the range of 5 to 10%. Even higher lipid concentrations (up to 40%) have been 
homogenized to lipid nanodispersions.16 
Two general approaches of the homogenization step, the hot and the cold 
homogenization techniques, can be used for the production of SLN.17,18 In both 
cases, a preparatory step involves the drug incorporation into the bulk lipid by 
dissolving the drug in the lipid melt. 
2.4. Hot homogenization 
The hot homogenization is carried out at temperatures above the melting point of 
lipid. Therefore, it is in fact the homogenization of an emulsion. A preemulsion of 
the drug loaded lipid melt and the aqueous emulsifier phase (same temperature) is 
obtained by high-shear mixing device (Ultraturrax). The quality of the preemulsion 
is very important for the final product quality. In general, higher temperatures 
result in lower particle sizes due to the decrease of the viscosity of the inner phase.19 
However, high temperatures may also increase the degradation rate of the drug and 
the carrier. The homogenization step can be repeated several times. It should be kept 
in mind however, that HPH increases the temperature of the sample (approximately 
10C for 500 bar20). In most cases, 3 to 5 homogenization cycles at 500 to 1500 bar are 
sufficient. Increasing the homogenization pressure or the number of cycles often 
results in an increase of the particle size due to particle coalescence, which occurs 
as a result of the high kinetic energy of the particles.21 
It is important to note that the primary product of the hot homogenization is a 
nanoemulsion due to the liquid state of the lipid. Solid particles are expected to be 
formed by the following cooling of the sample to room temperature, or to temperatures 
below. Due to the small particle size and the presence of emulsifiers, lipid 
crystallization may be highly retarded and the sample may remain as a supercooled 
melt for several months.22 
2.5. Cold homogeniza Hon 
Cold homogenization has been developed to overcome the following three problems 
of the hot homogenization technique: 
(1) Temperature induced drug degradation 
(2) Drug distribution into the aqueous phase during homogenization 
(3) Complexity of the crystallization step of the nanoemulsion, leading to several 
modifications and/or supercooled melts 
The first preparatory step for cold homogenization is the same as in the hot homogenization 
procedure and includes the solubilization of the drug in the melt of the 
Solid Lipid Nanoparticles as Drug Carriers 191 
bulk lipid. However, the following steps differ. The drug containing melt is rapidly 
cooled. The high cooling rate favors a homogenous distribution of the drug within 
the lipid matrix. The solid, drug containing lipid is milled to microparticles. Typical 
particle sizes obtained by means of ball or mortar milling are in the range of 
50 to 100 microns. Low temperatures increase the fragility of the lipid, and therefore 
favor particle disruption. The solid lipid microparticles are suspended in a 
chilled emulsifier solution. The preemulsion is subjected to HPH at or below room 
temperature. An effective temperature control and regulation is needed in order to 
ensure the unmolten state of the lipid due to the increase in temperature during 
homogenization.20 In general, compared with hot homogenization, larger particle 
sizes and a broader size distribution are observed in cold homogenized samples.23 
A modified version of this technique has been recently published by the group of 
Miiller-Goymann. They dispersed a solid 1:1 lecithin-hardfat mixture (described as 
solid reversed micelles) in Tween containing water using high pressure homogenization.
2.5.1. SLN prepared by solvent emulsification/evaporation 
The solvent emulsification/evaporation processes adapts techniques which have 
been previously used for the production of polymeric micro- and nanoparticles. 
The solid lipid is dissolved in a water-immiscible organic solvent (e.g. cyclohexane, 
or chloroform) that is emulsified in an aqueous phase. Upon evaporation of the solvent, 
a nanoparticle dispersion is formed by precipitation of the lipid in the aqueous 
medium. Westesen prepared nanoparticles of tripalmitate by dissolving the triglyceride 
in chloroform.25 This solution was emulsified into an aqueous phase by high 
pressure homogenization. The organic solvent was removed from the emulsion by 
evaporation under reduced pressure. The mean particle size ranges from approximately 
30 to lOOnm depending on the lecithin/co-surfactant blend. Particles with 
very small diameters (30 nm) were obtained by using bile salts as co-surfactants. 
Comparable small particle size distributions were not achievable by melt emulsification 
of similar composition. The mean particle size depends on the concentration 
of the lipid in the organic phase. Very small particles could only be obtained with 
low fat loads (5 w%) related to the organic solvent. With increasing lipid content, the 
efficacy of the homogenization declines due to the higher viscosity of the dispersed 
2.5.2. SLN preparations by solvent injection 
The solvent injection method has been developed by Fessi to produce polymer 
nanoparticles.26 Nanoparticles were only produced with solvents which distribute 
very rapidly into the aqueous phase (e.g. ethanol, acetone, DMSO), while larger 
192 Mader 
particle sizes were obtained with more lipophilic solvents. According to Fessi, the 
particle size is critically determined by the velocity of the distribution processes and 
only water miscible solvents can be used. The solvent injection method can also be 
used for the production of solid lipid nanoparticles.27'28 However, the method is 
limited to lipids which dissolve in the polar organic solvent. Advantages of the 
method are the avoidance of elevated temperatures and high shear stress. However, 
the lipid concentration in the primary suspension will be less compared with 
High-Pressure-Homogenization. Furthermore, the use of organic solvents clearly 
represents a drawback of the method. 
2.5.3. SLN preparations by dilution of microemulsions 
or liquid crystalline phases 
SLN preparation techniques which are based on the dilution of microemulsions 
have been developed by Gasco and coworkers. Unfortunately, there is no common 
agreement within the scientific community about the definition of a microemulsion. 
One part of the scientific community understands under microemulsions 
high fluctuating systems which can be regarded as a critical solution, and therefore 
do not contain an inner and outer phase. This model has been confirmed 
by self-diffusion NMR studies of Lindman.29 In contrast, Gasco and other scientists 
understand microemulsions as two systems composed of an inner and 
outer phase (e.g. O/W-microemulsions). They are made by stirring an optical 
transparent mixture at 65-70C, typically composed of a low melting lipid fatty 
acid (e.g. stearic acid), emulsifier (e.g. polysorbate 20, polysorbate 60, soy phosphatidylcholin, 
taurodeoxycholic acid sodium salt), co-emulsifiers (e.g. Butanol, 
Na-monooctylphosphate), and water. The hot microemulsion is dispersed in cold 
water (2-3C) under stirring. Typical volume ratios of the hot microemulsion to 
the cold water are in the range of 1:25 to 1:50. The dilution process is critically 
determined by the composition of the microemulsion. According to the literature, 
the droplet structure is already contained in the microemulsion, and therefore, no 
energy is required to achieve submicron particle sizes.30,31 The temperature gradient 
and the pH-value determine the product quality in addition to the composition 
of the microemulsion. High temperature gradients facilitate rapid lipid crystallization 
and prevent aggregation.32'33 Due to the dilution step, lipid contents which are 
achievable are considerably lower, compared with the HPH based formulations. 
Another disadvantage includes the use of organic solvents. 
Recent work describes a similar approach to produce SLN. A hot liquid crystalline 
phase (instead of a microemulsion) is diluted in cold water to yield a solid 
lipid nanodispersion.34 This approach avoids the use of high pressure homogenization 
and organic solvents, and therefore represents an interesting opportunity. 
Solid Lipid Nanoparticles as Drug Carriers 193 
2.6. Further processing 
2.6A. Sterilization 
Sterility is required for parenteral formulations. Dry or wet heat, filtration, 
y-irradiation, chemical sterilization and aseptic production are general, opportunities 
to achieve sterility. The sterilization should not change the properties of the 
sample with respect to physical and chemical stability and the drug release kinetics. 
Sterilization by heat is a reliable procedure which is most commonly used. It was 
also applied for Liposomes.35,36 Steam sterilization will cause the formation of an 
oil in water emulsion, due to the melting of the lipid particles. The formation of SLN 
requires recrystallization of the lipids. Concerns are related to temperature induced 
changes of the physical and chemical stability. The correct choice of the emulsifier 
is of significant importance for the physical stability of the sample at high temperatures. 
Increased temperatures will affect the mobility and the hydrophilicity of 
all emulsifiers, but to a different extent. Schwarz found that Lecithin is preferable 
to Poloxamer for steam sterilization, as only a minor increase in the particle size 
and the number of microparticles was observed after steam sterilization.37'38 An 
increase in particle size for Poloxamer 188 stabilized Compritol-SLN was observed 
after steam sterilization. It was found that a decrease of the sterilization temperature 
from 121C to 110C can reduce sterilization induced particle aggregation to a 
large extent. This destabilization can be attributed to the decreased steric destabilization 
of the Poloxamer. It is well known for PEG-based emulsifiers that increased 
temperatures lead to dehydration of the ethylenoxide chains, pointing to a decrease 
of the thickness of the protecting layer. It has been demonstrated by 1H-NMR spectroscopy 
on Poloxamer stabilized lipid nanoparticles, that even a moderate temperature 
increase from RT to 37C decreases the mobility of the ethylenoxide chains 
on the particle surface.39 Results of Freitas et dl. indicate that the lowering of the 
lipid content (to 2%), and the surface modification of the glass vials and nitrogen 
purging might prevent the particle growth to a large extent and avoid gelation.40 
Further studies of Cavalli et al.4* and Heiati42 demonstrate the possibility of steam 
sterilization of drug loaded SLN. 
Filtration sterilization of dispersed systems requires very high pressure and is 
not applicable to particles >0, 2 /nm. As most SLN particles are close to this size, 
filtration is of no practical use, due to the clocking of the filters. Few studies investigated 
the possibility of y-sterilization. It must be kept in mind that free radicals 
are formed during y-sterilization in all samples, due to the high energy of the yrays. 
These radicals may recombine with no modification of the sample or undergo 
secondary reactions which might lead to chemical modifications of the sample. 
The degree of sample degradation depends on the general chemical reactivity and 
the molecular mobility and the presence of oxygen. It is therefore not surprising 
194 Mader 
that chemical changes of the lipid bilayer components of liposomes were observed 
after y-irradiation.43 Schwarz investigated the impact of different sterilization techniques 
[steam sterilization at 121C (15min) and 110C (15min); y-sterilization] on 
SLN characteristics.37'38 In comparison to lecithin stabilized systems, Poloxamer 
stabilized SLN were less stable than steam sterilization. However, this difference 
was not detected for y-sterilized samples. Compared with steam sterilization at 
121 C, the increase in particle size after y-irradiation was lower, but comparable to 
that at 110C. 
Unfortunately, most investigators did not search for steam sterilization or irradiation 
induced chemical degradation. It should be kept in mind that degradation 
does not always cause increased particle sizes. In contrast, the formation of species 
like lysophosphatides or free fatty acids could even preserve small particle sizes, 
but might cause toxicological problems. Further studies with more focus on chemical 
degradation products are clearly necessary to permit valid statements of the 
possibilities of SLN sterilization. 
2.6.2. Drying by lyophilization, nitrogen purging and spray drying 
SLN are thermodynamic unstable systems, and therefore, particle growth has to be 
minimized. Furthermore, SLN ingredients and incorporated drugs are often unstable, 
hydrolyzing or oxidizing. The transformation of the aqueous SLN-suspension 
in a dry, redispersible powder is therefore often a necessary step to ensure storage 
stability of the samples. Lyophilization is widely used and is a promising way 
to increase chemical and physical SLN stability over extended periods of time. 
Lyophilization also offers principle possibilities for SLN incorporation into pellets, 
tablets or capsules. 
Two additional transformations are necessary which might be the source of 
additional stability problems. The first transformation, from aqueous dispersion to 
powder, involves the freezing of the sample and the evaporation of water under vacuum. 
Freezing of the sample might cause stability problems due to the freezing out 
effect which results in the changes of the osmolarity and the pH. The second transformation, 
resolubilization, involves situations at least in its initial stages which 
favor particle aggregation (i.e. low water and high particle content, high osmotic 
The protective effect of the surfactant can be compromised by lyophilization.44 
It has been found that the lipid content of the SLN dispersion should not exceed 
5%, so as to prevent an increase in the particle size. Direct contact of lipid particles 
are decreased in diluted samples. Furthermore, diluted SLN dispersions 
will also have higher sublimation velocities and a higher specific surface area.45 
The addition of cryoprotectors (e.g. Sorbitol, Mannose, Trehalose, Glucose, and 
Solid Lipid Nanoparticles as Drug Carriers 195 
Polyvinylpyrrolidon) will be necessary to decrease SLN aggregation and to obtain 
a better redispersion of the dry product. Schwarz et al. investigated the lyophilization 
of SLN in detail.46 Best results were obtained with the cryoprotectors, Glucose, 
Mannose, Maltose and Trehalose, in the concentration range between 10% and 
15%. The observations come into line with the results of the studies on liposome 
lyophilization, which indicated that Trehalose was the most sufficient substance 
to prevent liposome fusion and the leakage of the incorporated drug.47 Encouraging 
results obtained with unloaded SLN cannot predict the quality of drug loaded 
lyophilizates. Even low concentrations of 1% Tetracain or Etomidat caused a significant 
increase in particle size, excluding an intravenous administration.46 
Westesen investigated the lyophilization of tripalmitate-SLN using glucose, 
sucrose, maltose and trehalose as cryoprotective agents.48 Handshaking of redispersed 
samples was an insufficient method, but bath sonification produced better 
results. Average particle sizes of all lyophilized samples with cryoprotective agents 
were 1.5 to 2.4 times higher than the original dispersions. One year storage caused 
increased particle sizes of 4 to 6.5 times compared with the original dispersion. 
In contrast to the lyophilizates, the aqueous dispersions of tyloxapol/phospholid 
stabilized tripalmitate SLN exhibited remarkable storage stability. The instability 
of the SLN lyophilizates can be explained by the sintering of the particles. TEM pictures 
of tripalmitate SLN show an anisometrical, platelet-like shape of the particles. 
Lyophilization changes the properties of the surfactant layer due to the removal of 
water, and increases the particle concentration which favors particle aggregation. 
Increased particle sizes after lyophilization (2.1 to 4.9 times) were also reported by 
Cavalli.41 Heiati compared the influence of four cryoprotectors (i.e. trehalose, glucose, 
lactose and mannitol) on the particle size of azidothymidine palmitate loaded 
SLN lyophilizates.42 In agreement to other reports, Trehalose was found to be the 
most effective cryoprotectant. The freezing procedure will affect the crystal structure 
and the properties of the lyophilizate. Literature data suggest that the freezing 
process needs to be optimized to a particular sample size. Schwarz recommended 
rapid freezing in liquid nitrogen.46 In contrast, other researchers observed the best 
results after a slow freezing process.49 Again, best results were obtained with samples 
of low lipid content and with the cryoprotector trehalose. Slow freezing in a 
deep freeze (70C) was superior to rapid cooling in liquid nitrogen. Furthermore, 
introduction of an additional thermal treatment of the frozen SLN dispersion (2 hr at 
22C; followed by 2 hr temperature decrease to  40C) was found to improve the 
quality of the lyophilizate. Lately, lyophilization has been used to stabilize retinoic 
acid loaded SLN.50 
An interesting alternative to lyophilization has been recently suggested by 
Gasco's group. Drying with a nitrogen stream at low temperatures of 3 to 10C 
has been found to be superior.51 Compared with lyophilization, the advantages of 
196 Mader 
this process are the avoidance of freezing and the energy efficiency resulting from 
the higher vapor pressure of water. 
Spray drying has been scarcely for SLN drying, although it is cheaper compared 
with lyophilization. Freitas obtained a redispersable powder with this method, 
which meets the general requirements of i.v.-injections, with regard to the particle 
size and the selection of the ingredients.52 Spray drying might potentially cause 
particle aggregation due to high temperatures, shear forces and partial melting 
of the particles. Freitas recommends the use of lipids with high melting points 
>70C to avoid sticking and aggregation problems. Furthermore, the addition of 
carbohydrates and low lipid contents favor the preservation of the colloidal particle 
size in spray drying. 
3. SLN Structure and Characterization 
The characterization of SLN is a necessity and a great challenge. Lipid characterization 
itself is not trivial as the statement by Laggner shows53: "Lipids and fats, as soft 
condensed material in general, are very complex systems, which not only in their 
static structures but also with respect to their kinetics of supramolecular formation, 
Hysteresis phenomena or supercooling can gravely complicate the task of defining 
the underlying structures and boundaries in a phase diagram". This is especially 
true for lipids in the colloidal size range. Therefore, possible artifacts caused by sample 
preparation (removal of emulsifier from particle surface by dilution, induction 
of crystallization processes, changes of lipid modifications) should be kept in mind. 
For example, the contact of the SLN dispersion with new surfaces (e.g. a syringe 
needle) might induce lipid crystallization or modification, and sometimes result in 
the spontaneous transformation of the low viscous SLN-dispersion into a viscous 
gel. The most important parameters of SLN include particle size and shape, the 
kind of lipid modification and the degree of crystallization, and the surface charge. 
Photon correlation spectroscopy (PCS) and Laser Diffraction (LD) are the most 
powerful techniques for routine measurements of particle size. It should be kept in 
mind that both methods are not "measuring" particle sizes. Rather, they detect 
light scattering effects which are used to calculate particle sizes. For example, 
uncertainties may result from nonspherical particle shapes. Platelet structures commonly 
occur during lipid crystallization54 and are very often described in the SLN 
literature.55-59 The influence of the particle shape on the measured size is discussed 
by Sjostrom.55 Further difficulties arise both in PCS and LD measurements for samples 
which contain several populations of different size. Therefore, additional techniques 
might be useful. For example, light microscopy is recommended although 
it is not sensitive to the nanometer size range. It gives a fast indication about the 
Solid Lipid Nanoparticles as Drug Carriers 197 
presence and the character of microparticles. Electron Microscopy provides, in contrast 
to PCS and LD, direct information on the particle shape.57'58 Atomic force 
microscopy (AFM) has attracted increasing attention. A cautionary note applies to 
the use of AFM in the field of nanoparticles, as an immobilization of the SLN by 
solvent removal is required to assess their shape by the AFM tip. This procedure is 
likely to cause substantial changes of the molecular structure of the particles. Zur 
Miihlen demonstrated the ability of AFM to image the morphological structure of 
SLN.60 The sizes of the visualized particles are of the same magnitude, compared 
with the results of PCS measurements. The AFM investigations revealed the disklike 
structure of the particles. Dingier investigated cetylpalmitate SLN (stabilized 
by polyglycerol methylglucose distearate, Tego Care 450) by electron microscopy 
and AFM and found an almost spherical form of the particles.61 The usefulness of 
cross flow Field-Flow-Fractionation (FFF) for the characterization of colloidal lipid 
nanodispersions has been recently demonstrated.58 Lipid nanodispersions with 
constant lipid content, but different ratios of liquid and solid lipids did show similar 
particle sizes in dynamic light scattering. However, retention times in FFF were 
remarkably dissimilar due to the different particle shapes (i.e. spheres vs. platelets). 
Anisotropic particles such as platelets will be constrained by the cross flow much 
more heavily compared with the spheres of similar size. The very high anisometry 
of the SLN particles has been confirmed by electron microscopy, where very thin 
particles of 15 nm thickness and the length of several hundred nanometers became 
The measurement of the zeta potential allows predictions about the storage 
stability of colloidal dispersions.62 In general, particle aggregation is less likely 
to occur for charged particles (i.e. high zeta potential) due to electric repulsion. 
However, this rule cannot strictly apply to systems which contain steric stabilizers, 
because the adsorption of steric stabilizer will decrease the zeta potential due to the 
shift in the shear plane of the particle. 
Particle size analysis is just one aspect of SLN quality. The same attention has to 
be paid on the characterization of lipid crystallinity and modification, because these 
parameters are strongly correlated with drug incorporation and release rates. Thermodynamic 
stability and lipid packing density increase, and drug incorporation 
rates decrease in the following order: 
supercooled melt < a-modification < B'-modification < 6-modification 
In general, it has been found that melting and crystallization processes of 
nanoscaled material can differ considerable from that of the bulk material.63 The 
thermodynamic properties of material having small nanometer dimensions can be 
considerably different, compared with the material in bulk form (e.g. the reduction 
198 Mader 
of melting point). This occurs because of the tremendous influence of the surface 
This statement is also valid for SLN, where lipid crystallization and modification 
changes might be highly retarded,64 due to the small size of the particles and 
the presence of emulsifiers. Moreover, crystallization might not occur at all and 
has been shown that samples which were previously described as SLN (solid lipid 
particles) were in fact supercooled melts (liquid lipid droplets).65 The impact of 
the emulsifier on SLN lipid crystallization has been shown by Bunjes.66 The same 
group demonstrated also a size dependent melting of SLN.67 
Differential Scanning Calorimetry (DSC) and X-ray scattering are most commonly 
applied to asses the status of the lipid. DSC uses the fact that different lipid 
modifications possess different melting points and melting enthalpies. By means 
of X-ray scattering, it is possible to assess the length of the long and short spacings 
of the lipid lattice. It is highly recommended to measure the SLN dispersion 
themselves, because solvent removal will lead to modification changes. Sensitivity 
problems and long measurement times of convential X-ray sources might be 
overcome by synchrotron irradiation.64 In addition, this method permits to conduct 
time resolved experiments and allows the detection of intermediate states 
of colloidal systems which will be non detectable by convential X-ray methods.53 
Recent work shows that SLN might form superstructures by parallel alignment of 
SLN platelets. These reversible particle self-assemblies were observed by Illing 
et al. in tripalmitin dispersions when the lipid concentration exceeds 40mg/g. 
Higher lipid concentrations did enhance particle self-assembly. The tendency to 
form self-assemblies has been found to depend on the particle shape, the lipid 
and the surfactant concentration.68 Infrared and Raman Spectroscopy are useful 
tools to investigate structural properties of lipids and they might give complentary 
information to X-ray and DSC.54 Raman measurements on SLN show that the 
arrangement of lipid chains of SLN dispersions changes with storage.69 
Rheometry might be particularly useful for the characterization of the viscoelastic 
properties of SLN dispersions. The rheological properties are important with 
respect to the dermatological use of SLN, but they also provide useful information 
about the structural features of SLN dispersions and their storage dependency. 
Studies of Lippacher show that the SLN dispersion posses higher elastic properties 
than emulsions of comparable lipid content.70-72 Furthermore, a sharp increase of 
the elastic module is observed at a certain lipid content. This point indicates the 
transformation from a low viscous lipid dispersion to an elastic system, with a 
continuous network of lipid nanocrystals. Illing and Unruh did compare the rheological 
properties of trimyristic, tripalmitic and tristearic SLN suspensions. The 
results indicate that the viscosity of triglyceride suspensions increases with the 
lipid chain length and an increased anisotropy of the particles.73 Souto et al. used 
Solid Lipid Nanoparticles as Drug Carriers 199 
rheology to study the influence of SLN addition on the rheological properties of 
The co-existence of additional colloidal structures (micelles, liposomes, mixed 
micelles, nanodispersed liquid crystalline phases, supercooled melts, drugnanoparticles) 
has to be taken into account for all SLN dispersions. Unfortunately, 
many investigators neglect this aspect, although the total amount of surface active 
compounds is often comparable to the total amount of the lipid. The characterization 
and quantification are serious challenges due to the similarities in size. In addition, 
the sample preparation will modify the equilibrium of the complex colloidal 
system. Dilution of the original SLN dispersion with water might cause the removal 
of surfactant molecules from the particle surface and induce further changes such 
as crystallization or the changes of the lipid modifications. It is therefore highly 
desirable to use methods which are sensitive to the simultaneous detection of different 
colloidal species, which do not require preparatory steps such as Raman, 
NMR and ESR spectroscopy. 
NMR active nuclei of interest are 1H, 13C, 19F and 35P. Due to the different chemical 
shifts, it is possible to attribute the NMR signals to particular molecules or their 
segments. For example, lipid methyl protons give signals at 0.9 ppm, while protons 
of the polyethylenglycole chains give signals at 3.7 ppm. Simple ^-spectroscopy 
permits an easy and rapid detection of supercooled melts, due to the low linewidths 
of the lipid protons69,75-77. This method is based on the different proton relaxation 
times in the liquid and semisolid/solid state. Protons in the liquid state give sharp 
signals with high signal amplitudes, while semisolid/solid protons give very broad 
or invisible NMR signals under these circumstances. NMR has been used to characterize 
calixarene SLN78 and hybrid lipid particles (NLC), which are composed 
of liquid and solid lipids.59 Protons from solid lipids are not detected by standard 
NMR, but they can be visualized by solid state NMR. A drawback of solid 
state NMR is the rapid spinning of the sample that might cause artifacts. A recent 
paper describes the use of this method to monitor the distribution of Q10 in lipid 
matrices.79 Unfortunately, the authors did use "drying of the sample to constant 
weight" as a preparatory step, which will cause significant changes of the sample 
ESR requires the addition of paramagnetic spin probes to investigate SLN dispersions. 
A large variety of spin probes is commercially available. The corresponding 
ESR spectra give information about the microviscosity and micropolarity. ESR 
permits the direct, repeatable and non-invasive characterization of the distribution 
of the spin probe between the aqueous and the lipid phase.80 Experimental results 
demonstrate that storage induced crystallization of SLN leads to an expulsion of 
the probe out of the lipid into the aqueous phase.81 Furthermore, using an ascorbic 
acid reduction assay, it is possible to monitor the time scale of the exchange between 
200 Mader 
the aqueous and the lipid phase.59 The transfer rates of molecules between SLN and 
liposomes or cells have been determined by ESR.82 
4. The "Frozen Emulsion Model" and Alternative SLN Models 
Lipid nanoemulsions are composed of a liquid oily core and a surfactant layer 
(lecithin). They are widely used for the parenteral delivery of poorly soluble 
drugs.83-85 The original idea of SLN was to achieve a controlled release of incorporated 
drugs by increasing the viscosity of the lipid matrix. Therefore it is not 
surprising that in original model, SLN is being described as "frozen emulsions" 
(see Fig. 1, left and middle).8687 However, lipids are known to crystallize very frequently 
in anisotropic platelet shapes54 and anisotropic. Sjostrom et al. described 
in 1995 that the particle shape of Cholesterylacetate SLN did strongly depend on 
the emulsifier.55 Platelet shaped particles have been detected for lecithin stabilized 
particles, while PEG-20-sorbitanmonolaurate stabilized particles preserved their 
spherical shape. Anisotropic particles have been found in numerous other SLN 
dispersions.56-59 Based on the experimental results, a platelet shaped SLN model 
can be proposed as an alternative (see Fig. 1, right). 
In the year 2000, Westesen questioned the frozen emulsion droplet model with 
the following statement88: 
"Careful physicochemical characterization has demonstrated that these lipid-based 
nanosuspensions (solid lipid nanoparticles) are not just emulsions with solidified 
During the development process of these systems, interesting phenomena have 
been observed, such as gel formation on solidification and upon storage, unexpected 
dynamics of polymorphic transitions, extensive annealing of nanocrystals 
over significant periods of time, stepwise melting of particle fractions in the 
Nanoemulsion SLN: "Frozen emulsion droplet" SLN: Platelet shaped particles o o  
Core: liquid lipid (oil) S Core: solid lipid H Shell: stabilizer 
Fig. 1. General structure of a nanoemulsion (left), and proposed models for SLN: Frozen 
emulsion droplet model (middle) and platelet shaped SLN model (right). 
Solid Lipid Nanoparticles as Drug Carriers 201 
lower-nanometer-size range, drug expulsion from the carrier particles on crystallization 
and upon storage, and extensive supercooling." 
Her comment highlights the complex behavior and changes of SLN dispersions. 
In addition, the presence of competing colloidal structures (e.g. micelles, 
liposomes, mixed micelles, nanodispersed liquid crystalline phases, supercooled 
melts and drug-nanoparticles) should be considered. Additional colloids might 
have an impact on very different aspects, including the correct measurement of 
particle size, drug incorporation and toxicity. A recent study shows that the cell 
toxicity of the SLN dispersion was reduced by dialysis due to the removal of water 
soluble components.89 
5. Nanostructured Lipid Carriers (NLC) 
Nanostructured lipid carriers (NLC) have been recently proposed as a new SLN 
generation with improved characteristics.90 The general idea behind the system is 
to improve the poor drug loading capacity of SLN by "mixing solid lipids with 
spatially incompatible lipids leading to special structures of the lipid matrix",91 
while still preserving controlled release features of the particles. Three different 
types of NLC have been proposed (NLC I: The imperfect structured type, NLC 
II: The structureless type and NLC III: The multiple type). Unfortunately, these 
structural proposals have not been supported by experimental data. They assume 
a spherical shape and they are not compatible with lipid platelet structures. 
For example, NLC III structures should contain small oily droplets in a solid 
lipid sphere (Fig. 2, left). Detailed analytical examination of NLC systems by Jores 
et al. demonstrate that "nanospoon" structures are formed, in which the liquid oil 
adheres on the solid surface of a lipid platelet (Fig. 2, right). 
Jores et al. did conclude that "Neither SLN nor NLC lipid nanoparticles showed 
any advantage with respect to incorporation rate or retarded accessibility to the 
drug, compared with conventional nanoemulsions. The experimental data concludes 
that NLCs are not spherical solid lipid particles with embedded liquid 
liquid lipid (oil) J A solid lipid  stabilizer 
Fig. 2. Proposed NLC III structure (modified after91) and experimental determined 
"nanospoon" structure described by Jores et al. (side view of particle).58'59 
202 Mader 
droplets, but rather, they are solid platelets with oil present between the solid 
platelet and the surfactant layer". Very similar structures have been found on Q10 
loaded SLN by Bunjes et til.92 
6. Drug Localization and Release 
Proposed advantages of SLN, compared with nanoemulsions, include increased 
protection capacity against drug degradation and controlled release possibilities 
due to the solid lipid matrix. The general low capacity of crystalline structures to 
accommodate foreign molecules is a strong argument against the proposed rewards. 
It is therefore necessary to distinguish between drug association and drug incorporation. 
Drug association means that the drug is associated with the lipid, but it 
might be localized in the surfactant layer or between the solid lipid and the surfactant 
layer (similar to the oil in Fig. 2, right). Drug incorporation would mean 
the distribution of the drug within the lipid matrix. Another limiting aspect comes 
from the fact that the platelet structure of SLN, which is found in many systems, 
leads to a tremendous increase in surface area and the shortening of the diffusion 
lengths. Furthermore, additional colloid structures present in the sample are 
alternatives for drug localization the SLN for drug incorporation as it was pointed 
out by Westesen88: "The estimation of drug distribution is difficult for dispersions 
consisting of more than one type of colloidal particle. Depending on the type of 
stabilizer and on the concentration ratio of stabilizer to matrix material significant 
numbers of particles such as liposomes and/or (mixed) micelles may coexist with 
the expected type of particles". 
The detailed investigation of drug localization is very difficult and only a few 
studies exist. Parelectric spectroscopy has been used to investigate the localization 
of glucocorticoids. The results indicate that the drug molecules are attached to the 
particle surface, but not incorporated into the lipid matrix. With Betamethasonvalerate, 
the loading capacity of the particle surface was clearly below the usual concentration 
of 0.1%.93 Lukowski used Energy Dispersive X-ray Analysis and found 
that the drug Triamcinolone, Dexamethasone and Chloramphenicol are partially 
stored at the surface of the individual nanoparticles.94 
The importance of the emulsifier is reflected in a study from Danish scientists.95 
They produced gamma-cyhalothrin (GCH) loaded lipid micro- and nanoparticles. 
GCH had only limited solubility in the solid lipid and was expulsed during storage. 
The appearance of GCH crystals was strongly dependent from the solubility 
of the GCH in the emulsifier solutions. Emulsifier with high GCH solubility provoked 
rapid crystal growth. This observation is in accordance with a mechanism of 
crystal growth according to Ostwald ripening. Slovenian scientist found that ascorbylpalmitate 
was more resistant against oxidation in non-hydrogenated soybean 
Solid Lipid Nanoparticles as Drug Carriers 203 
lecithin liposomes, compared with SLN.96 It shows that liposomes might have a 
higher protection capacity compared with SLN. 
Fluorescence and ESR studies have been used by Jores et al. to monitor the 
microenvironment and the mobility of model drugs. The results indicate that even 
highly lipophilic compounds are pushed into a polar environment during lipid 
crystallization. Therefore, the incorporation capacity of SLN is very poor for most 
molecules.69 A nitroxide reduction assay gave results in accordance with the results 
of the distribution. Compared with nanoemulsions, nitroxides were more accessible 
in SLN and NLC to ascorbic acid, localized in the aqueous environment. Therefore, 
nanoemulsions were more protective than SLN and NLC systems. 
Drug release from SLN and NLC could be either controlled by the diffusion of 
the drug or the erosion of the matrix. The original idea was to achieve a controlled 
release of SLN due to the slowing down of drug diffusion to the particle surface. This 
idea is, however, questionable due to drug expulsion during lipid crystallization. 
In addition, very short diffusion lengths in nanoscaled delivery systems lead to 
short diffusion times, even in highly viscous or solid matrices. In most cases, the 
delivery of the drug will be controlled by the slow dissolution rate in the aqueous 
environment. Drug release rate will be highly dependent on the presence of further 
solubilizing colloids (e.g. micelles), which are able to work as a shuttle for the drug 
and the presence or absence of a suitable acceptor compartment. Many investigators 
studied only the release in buffer media. A controlled release pattern under such 
conditions is not surprising, as it is caused by low solubilization kinetics due to 
the poor solubility of the drug. In vivo, acceptor compartments will be present 
(e.g. lipoproteins, membranes) and will speed up release processes significantly. 
Whenever possible, drug loaded SLN should be compared with nanosuspensions 
to separate the general features of the drug and the influence of the lipid matrix. 
Results by Kristl et al. indicate that lipophilic nitroxides diffuse between SLN 
and liposomes. The diffusion kinetics was strongly dependent on the nitroxide 
structure. In contrast, uptake of nitroxides in cells was similar between lipophilic 
nitroxides, suggesting endocytosis as the main mechanism.82 The detailed mechanisms 
of drug release in vivo are poorly understood. In vitro data by Olbrich demonstrate 
that SLN are degraded by lipases.97,98 Degradation by lipase depends on the 
lipid and strongly on the surfactant. Steric stabilization (e.g. by poloxamer) of SLN 
and NLC are less accessible because lipase needs an interface for activation. It is 
also known that highly crystalline lipids are poorly degraded by lipase. 
7. Administration Routes and In Vivo Data 
SLN and NLC can be administrated at different routes, including peroral, dermal, 
intravenously and pulmonal. Peroral administration of SLN could enhance the drug 
204 Mader 
absorption and modify the absorption kinetics. Despite the fact that in most of the 
SLN, the drug will be associated but not incorporated in the lipid, SLN might have 
advantages due to enhanced lymphatic uptake, enhanced bioadhesion or increased 
drug solubilization by SLN lipolysis products such as fatty acids and monoglycerides. 
A serious challenge represents the preservation of the colloidal particle in 
the stomach, where low pH values and high ionic strengths favor agglomeration 
and particle growth. Zimmermann and Muller studied the stability of different 
SLN formulations in artificial gastric juice." The main findings of this study are 
that (i) some SLN dispersions preserve their particle size under acidic conditions, 
and (ii) there is no general lipid and surfactant which are superior to others. The 
particular interactions between lipid and stabilizer are determining the robustness 
of the formulation. Therefore, the suitable combination of ingredients has to be 
determined on a case by case basis. 
Several animal studies show increased absorption of poorly soluble drugs. The 
efficacy of orally administrated Triptolide free drug and Triptolide loaded SLN 
have compared in the carrageenan-induced rat paw edema by Mei et al.wo Their 
results suggest that SLN can enhance the anti inflammatory activity of triptolide 
and decrease triptolide-induced hepatotoxicity. The usefulness of SLN to increase 
the absorption of the poorly soluble drug all-trans retinoic acid has been shown 
by Hu et al. on rats.101 Gascos group investigated the uptake and distribution of 
Tobramycin loaded SLN in rats.102'103 They observed an increased uptake into the 
lymph, which causes prolonged drug residence times in the body of the animals. 
Furthermore, AUC and clearance rates did depend on the drug load. The same 
group described also enhanced absorption of Idarubicin-loaded solid lipid nanoparticles 
(IDA-SLN), in comparison to the drug solution. Furthermore, the authors 
described that SLN were able to pass the blood-brain barrier and concluded that 
duodenal administration of IDA-SLN modifies the pharmacokinetics and tissue 
distribution of idarubicin.104 
Parenteral administration of SLN is of great interest too. To avoid the rapid 
uptake of the SLN by the RES system, stealth SLN particles have been developed 
by the adoption of the stealth concept from liposomes and polymer nanoparticles. 
Reports indicate that Doxorubicin loaded stealth SLN circulate for long period of 
time in the blood and change the tissue distribution.105 Therefore, SLN could be 
alternatives to marketed stealth-liposomes, which can decrease the heart toxicity 
of this drug due to changed biodistribution. Long circulation times have also been 
observed for Poloxamer stabilized SLN with Paclitaxel.106 
The dermal application is of particular interest and it might become the main 
application of SLN.107 SLN pose occlusive properties which are related to the solid 
structure of the lipid.108 Human in vivo results of the group of Muller demonstrate 
that SLN can improve skin hydration and viscoelasticity.109 SLN have also 
Solid Lipid Nanoparticles as Drug Carriers 205 
UV protection capacity due to their reflection of UV light.110 Furthermore, data by 
Schafer-Korting suggest SLN can be used to decrease drug side effects due to SLN 
mediated drug targeting to particular skin layers.111 
Further reports describe additional applications of SLN as well as gene 
delivery,112 delivery to the eye,113 pulmonary delivery,114 and drug targeting of 
anticancer drugs.115 Studies of the different groups also propose the use of SLN for 
brain targeting to deliver MRI contrast agents116 or antitumour drugs.117'118 
8. Summary and Outlook 
SLN and NLC are now investigated by many scientists worldwide. In contradiction 
to early proposals, they certainly do not combine all the advantages of the other 
colloidal drug carriers and avoid the disadvantages of them. SLN are complex colloidal 
dispersions, not just "frozen emulsions". SLN dispersions are very susceptible 
to the sample history and storage conditions. Disadvantages of SLN include 
gel formation on solidification and upon storage, unexpected dynamics of polymorphic 
transitions, extensive annealing of nanocrystals over significant periods 
of time, stepwise melting of particle fractions in the lower-nanometer-size range, 
drug expulsion from the carrier particles on crystallization and upon storage, and 
extensive supercooling. The anisotropic shape of many SLN dispersions increases 
the surface area significantly, decreases the diffusion lengths to the surface and 
changes the rheological behavior dramatically (e.g. gel formation). Furthermore, 
the presence of alternative colloidal structures (micelles, liposomes) has to be considered 
to contribute to drug localization. In most cases, the drug will be associated 
with the lipid and not incorporated. Studies demonstrate that SLN and NLC might 
have no advantages compared with submicron emulsions, in regard to protection 
from the aqueous environment. 
On the other side, animal data suggest that SLN can change the pharmacokinetics 
and the toxicity of drugs. In many cases, drug incorporation might not be 
required and drug association with the lipid can be sufficient for lymphatic uptake. 
Clearly, more detailed studies are necessary to get a deeper understanding of the 
in vivo fate of these carriers. Whenever possible, SLN and NLC systems should 
be compared directly with alternative colloidal carriers (e.g. liposomes, nanoemulsions, 
nanosuspensions) to evaluate their true potential. 
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administration to rabbits of non-stealth and stealth doxorubicin-loaded solid 
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distribution of doxorubicin in brain and other tissues. / Drug Targ 10:327-335. 
Lipidic Core Nanocapsules as New 
Drug Delivery Systems 
Patrick Saulnier and Jean-Pierre Benoit 
A new generation of controlled size Lipidic core NanoCapsules (LNC) is presented 
with respect to their simple formulation, interfacial characteristics, pharmacokinetic 
and biodistribution properties. We describe their ability to load and release 
hydrophobic drugs. 
1. Introduction 
The ultimate goal of therapeutics is to deliver any drug at the right time in a safe 
and reproducible manner to a specific target at the required level. A great deal 
of effort is currently made to develop novel drug delivery systems that are able 
to fulfil these specifications. Among them, nanoscale drug carriers appear to be 
promising candidates. Colloidal carriers are particularly useful because they can 
provide protection of a drug from degradation in biological fluids and promote 
its penetration into cells. However, because the body is so well equipped to reject 
any intruding object, for the materials to stand any chance of success within this 
hostile yet sensitive environment, they must be chosen very carefully. In particular, 
attention has to be turned to the composition of the surface of colloidal drug 
carriers.1 Indeed, their clearance rate from the circulatory system is determined by 
their uptake by the mononuclear phagocytic system (MPS), which in turn depends 
on their physico chemical surface characteristics. In order to enhance circulation 
time, steric protection of various nanoparticulate drug carriers can be achieved by 
the presence of hydrophilic and flexible polymers to their surface. In the search 
214 Saulnier & Benoit 
for injectable, biocompatible and long-circulating systems, many colloidal systems 
have been evaluated. 
Different kinds of vectors can be used. For example, molecular vectors where 
the drug is complexed or associated to a transport molecule are currently used. 
Many vectors are also constituted by viruses or hybrid viruses, following the modification 
of their genomes in order to avoid the possibility of replication. In this 
way, they are used as gene delivery systems. However, we will focus on non viral 
vectors in this chapter. They are always formulated using soft physico chemical 
methods, by taking advantage of macromolecular self-assembly properties at the 
colloidal state in order to produce well-controlled particles. The number of required 
biological and physico chemical properties of these systems is high in order to formulate 
operant vectors. One of the most important specifications of these systems 
is the biocompatibility and biodegradability of each component that needs to be 
chosen carefully from a restricted list of molecules. Secondly, they need to be well 
constructed in terms of size and interf acial properties, in order to constitute stealthy 
systems that will not be phagocyted by the MPS and consequently will have the 
longest residence time in blood. 
We should not forget that such vectors exist biologically. Low density lipoproteins 
(LDL) are interesting systems possessing many of the required specifications. 
Unfortunately, their extraction, purification or reconstitution is still a challenge 
with strong physico chemical problems to solve. No convenient common solvent 
of proteins and lipids exists in order to reconstitute a similar supra-molecular framework. 
Consequently, we have to keep in mind a formulation of nanoparticles with 
biomimetic properties to those related to LDL as close as possible. 
We would now like to describe a novel class of nanoparticles (Lipidic core 
NanoCapsules:LNC) formulated without organic solvents with biocompatible and 
biodegradable molecules.2 We will see that after modification of the composition, 
we can control their size without difficulty in the 10-200 nm range, with a 
monomodal and narrow size distribution. 
Initially, we suggest describing the LNC formulation following some particular 
auto-organizational properties of Poly Ethylene Glycol (PEG)-like surfactants, 
induced by several emulsion-phase inversions in which they are incorporated. We 
will particularly emphasize the different physical methods that determine the characterization 
of the final structure of LNC, as well as their stability in suspensions. 
Then, we will describe strong correlations between their stealthy properties in blood 
and structural characteristics, mainly size and interfacial properties. In specific, we 
have evaluated the activation of the complement system in an original in vitro 
model. These nanocapsules are devoted to the encapsulation of drugs that need to 
be dispersed in their oily core. As a proof that the concept works, we will describe 
the ability of LNC to encapsulate and release simple lipophilic molecules, ibuprofene 
and amiodarone, in the last paragraph. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 215 
2. Lipidic Nanocapsule Formulation and Structure 
2.1. Process 
The first step of the process consists of the formulation of a stable emulsion characterized 
by its oily phase (O), aqueous phase (W) and finally its surfactants 
mixture (S). 
Due to the complexity of the mixture, the brand names will be used throughout 
the following text. It is important to note that no organic solvent or mediumchain 
alcohols are used in the formulation. All these molecules are known to be 
biocompatible and biodegradable. This indicates that the lack of residual toxicity 
can guarantee the safe use of LNC for human administration. Solutol is mainly 
comprised of 12-hydroxystearate of PEG 660 that corresponds to a hydrophilic surfactant 
(HLB = 11). The lecithin used is a mixture of hydrophobic phospholipids. 
The main compounds of each phase are reported in Table 1. 
The beginning of the formulation (see Fig. 1) corresponds to a magnetic stirring 
of all the components for which the proportions will be defined later, with a 
gradual rise in temperature from room temperature to 80C at a rate of 4C/min, 
leading to an W/O emulsion characterized by low conductivity. The system is 
Table 1 Compounds used in the LNC formulation. 
S  Solutol HS-15:12-hydroxysterarate of PEG 660 and PEG 660 (low content) 
 Lipoid: lecithin 
O  Labrafac: triglycerides (C8-C10) 
W  Purified water 
Fig. 1. Emulsion-phase inversion induced by temperature changes and the principle of 
LNC formulation. 
216 Saulnier & Benoft 
cooled from 80 to 55C (4C/min), leading to an O/W emulsion characterized by 
its high conductivity. Between these two kinds of emulsion, a transition zone called 
the Phase Inversion Zone (PIZ) is defined where the system is known to be in 
bicontinuous states.23 
In order to provide appropriate and optimal interfacial properties to the wateroil 
interfaces, the formulation typically requires three temperature cycles across the 
PIZ. The system is stopped at a temperature corresponding to the beginning of the 
PIZ, just before performing a final, fast-cooling dilution process in cold water (2C). 
This second step of the formulation leads to LNC in suspension in an aqueous phase. 
The interfacial rheology method developed in several papers demonstrates 
that the interfacial association of all the implicated molecules of the process is different 
from other commoner systems.4 Cohesion energy at the interface, as well 
as the interaction of the interfacial molecules with the adjacent phases, reaches a 
minimum for the concentrations used. We think that this particularity can explain 
why the system can be broken down in an ideal way during final dilution. The 
surfactants involved in the stabilization of the bicontinuous systems can easily 
leave the microemulsion in order to constitute the colloidal structures (LNC). 
It might be noted that temperatures corresponding to the PIZ are much too high 
to decline this method to the simple encapsulation of thermo-sensitive molecules. 
Fortunately, we have shown that the electrolyte concentration (NaCl) strongly influences 
the location of PIZ on the temperature scale. When we increase the electrolyte 
concentration, we decrease the PIZ temperature to reach acceptable levels. 
2.2. Influence of the medium composition 
Obviously, the presence or not of LNC strongly depends on the composition of 
the system reported in Fig. 2(a) as a pseudo-ternary diagram.5 Each point corresponds 
to strictly similar formulation processes and the entire diagram describes 
the appropriate feasibility zone. 
It should be noticed that the optimal formulation corresponds to w /w concentration 
of around 20% for the oil phase, 60% for the water phase and 20% for 
Solutol. In the zone corresponding to the LNC formulation, a statistical model 
is applied in order to approximate the influence of the composition on the size 
distribution measured by the dynamic light scattering method. 
Polynomial interpolations between well-controlled points are performed. The 
corresponding results are reported in Fig. 2(b) where different iso-size curves are 
presented. The same procedure was applied to the size variation coefficients. These 
two curve beams are powerful tools, allowing an optimized formulation to be 
found, once a given and reproducible size distribution is elaborated just by tuning 
the composition. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 21 7 
(a) (b) 
Fig. 2. Feasibility diagram of LNC. a: zone of favorable formulation; b: iso-size curves in 
the favorable zone. 
Fig. 3. Schematic representation of LNC. 
It is important to note that LNC have demonstrated very good freeze-drying 
and stability characteristics in storage conditions for several months, as determined 
by DSC measurements,6 confirming the structure presented in Fig. 3. 
LNCs are constituted of a lipidic core surrounded by a surfactant shell, where 
lecithin is located in the inner part of the shell and the Solutol in the outer part. 
2.3. Structure and purification of the LNC by dialysis 
Considering that in the biological environment of the blood stream, the particles 
interact strongly with various interfaces, one possible model for studying the interfacial 
behavior of these particles is their spreading at the air-water interface. Classically, 
the Langmuir balance was used to describe interfaces composed by simple 
21 8 Saulnier & Beno?t 
mixtures. The basic technique was the measurement of the surface pressure (7r)-area 
(A) isotherm, by determining the decrease in surface tension as a function of the 
area available for each molecule on the aqueous sub phase. This included the study 
of the monolayer formation, the compressibility of the interface, the mutual interactions 
of molecules in the monolayer, but also interactions with the sub-phase 
molecules across interfacial rheological measurements.7 Following this, these suspension 
spreading results were compared with zeta potential measurements. These 
studies8,9 clearly indicate that the mother suspension, just after dilution in cold 
water, is composed of 
 Stable nanocapsules as described before; these objects diffuse strongly in the 
aqueous phase after spreading at the air/water interface. 
 Unstable nanocapsules with similar size, but with a lower amount of phospholipids 
(Lipoid) in the inner part of their shell. These capsules are not sufficiently 
robust to support the interfacial energies during spreading. Consequently, the 
components or fragments of the initial particles can be detected at the air-water 
 Free PEG (minor component of the Solutol) released from the outer part of 
the shell. 
It is obvious that the excess of PEG, as well as an important fraction of the 
unstable particles could be limited by dialysis. We will see in the next chapter an 
original investigation of these dialysis effects. 
2.4. Imagery techniques 
AFM images [Fig. 4(a)] were obtained after spreading the initial suspension of 
50 nm (10 nm) LNC on a fresh mica plate, and then allowing a complete evaporation 
of the water at room temperature. A contact mode was applied with a contact 
force of 10 nN, as well as a non contact mode without modification of the related 
images. The particle shape looked like a cylinder, 2nm high and 275 nm wide, 
corresponding to a total volume similar to a 60 nm sphere. We demonstrate the 
deformation of LNC after water evaporation, but without fusion of the particles, 
something that often occurs with liposomes. 
Classical TEM images were taken of the covered copper grids, following staining 
with a 2% phosphotungstic acid aqueous solution. It is noted on Fig. 4(b) that 
the lateral diameters are relatively polydispersed in a 20-70 nm range. 
Fig. 4(c) corresponds to a cryo-TEM image (kindly provided by Olivier Lambert, 
IECB-UBS UMR CNRS 5471) where individualized LNC are detectable. It is 
important to note that this image was performed after a dialysis, followed by an 
appropriate dilution of the mother suspension. 
Lipidic Core Nanocapsules as New Drug Delivery Systems 219 
(b) TEM (c) Cryo-TEM 
<>,J 'HUE ^ > 
* "s ' * * 
*o * If 
Fig. 4. Visualization of LNC by (a) AFM, (b) TEM and (c) cryo-TEM. 
3. Electrical and Biological Properties 
3.1. Electro kinetic comportment 
The stable Lipid NanoCapsules (LNC) contain pegylated 12-hydroxy stearate, as 
well as free PEG in the outer part of the shell, which can be an important biological 
specification that we will describe latter. The distribution of PEG chains at the 
surface was determined by their electrokinetic properties. Thus, electrophoretic 
mobility was measured as a function of ionic strength and pH, for particles differing 
in sizes, dialysis effects, and the presence or not of lecithin in their shell. The study 
enabled us to find the isoelectric point (IEP) as well as the charge density (ZN) in 
relation to the dipolar distribution in the polyelectrolyte accessible layer (thickness 
1 A), by using soft particle electrophoresis analysis10 (see Fig. 5). 
This study showed that LNC presented electrophoretic properties conferred by 
PEG groups at the surface constituting dipoles that are able to interact with counter 
ions (H+, Na+) or water dipoles. 
The levels of IEP, ZN and 1/1 changed after dialysis, due to the removal of 
molecules that were poorly linked (mainly free PEG) at the outer part of the surface, 
allowing accessibility to the inner adjacent part of the shell. 
Water shell 
Fig. 5. Accessible layer to counter ions characterized by its thickness (1 A ) and its dipolar 
charge density (ZN). 
(a) AFM 
* 1 
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it i 
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-ViCS. . ' . i f , . 
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220 Saulnier & Beno?t 
100 nm LNC presented the best-organized and the accessible part of the shell, 
compared with other sizes of LNC, before and after dialysis. Lecithin was found to 
be present in the inner part of the polyelectrolyte layer and was found to play a role 
in the disorganization of the outer part. Dialyzing LNC formulated with lecithin 
led to stable and well structured nanocapsules, ready for an in vivo use as a drug 
delivery system.11 
3.2. Evaluation of complement system activation 
Generally, after intravenous administration, nanoparticles (NP) are rapidly 
removed from the blood stream because they are recognized by cells of the MPS such 
as Kiipffer cells in the liver, or spleen and bone-marrow macrophages. However, a 
brush of PEG chains grafted on the surface is known to decrease the recognition of 
nanoparticles by the immune system after intravenous administration.12 One has 
demonstrated that a strong correlation prevails between the complement activation 
and the stealthy properties of LNC. 
Therefore, these properties were evaluated by measuring the degree of complement 
activation11 [CH50 technique and crossed Immunoelectrophoresis (C3 cleavage)] 
and the level of macrophage uptake, in relation to the organization of PEG 
chains, according to the electrokinetic properties of the LNC surface. These experiments 
were performed on 20, 50 and 100 nm LNC before and after dialysis. The 
CH50 technique is presented in Fig. 6. 
Nanoparticles are dispersed in human serum with sensitized erythrocytes. 
After incubation, lysis is evaluated by a classical spectrophotometric method. The 
measured absorbance is related to the consumption of complement proteins by 
The main conclusions are that whatever the in vitro test, all LNC were not recognized 
by the non specific components of the immune system. It was probably due 
to the strong density of PEG chains at their surface. Furthermore, dialysis maintains 
a sufficiently high density of PEG and had no incidence on the complement 
4. Pharmacokinetic Studies and Biodistribution 
At first, the biodistribution of radiolabeled nanocapsules was studied by scintigraphy 
and y counting, after intravenous administration in rat whereby the 99mTc-oxine 
was incorporated in the lipid core and 125I labelled the shell of the nanocapsules.13 
Dynamic scintigraphic acquisition was carried out 3 hrs after administration and y 
activity in blood and tissues was followed for more than 24 hrs (see Fig. 7). 
An early half-disappearance time of about 47  6 min was found for 125I and 
41  11 min for 99mTc. These ranges of residence times were interesting for specific 
Lipidic Core Nanocapsules as New Drug Delivery Systems 221 
Lysis of 
^B Sheep erythrocyte 
 Complement proteins 
M Amibody anti-sheep eryihrocyie No lysis of 
Fig. 6. CH50 method for the evaluation of complement system activation. 
200 300 
Time (min) 
500 600 
Fig. 7. Evolution of radioactivity blood repartition after the intravenous administration of 
LNC expressed as a percentage of the injected dose. 
222 Saulnier & Benoit 
site delivery. Meanwhile, it appears that the length of the PEG chain (in this case, 
15 ethylene oxyde groups per molecule) should be increased to extend the vascular 
residence time. 
Recently, it has been shown that adding different DSPE-PEG to the system 
enhances the t1/2 values to several hours, depending on the concentration and the 
PEG length.14 t1/2 (half-life), MRT (Mean Residence Time) in blood and AUC (Area 
Under Curve) were evaluated by using [3H]-cholesteryl hexadecyl ether mixed with 
the lipid and the surfactant at the beginning of the formulation. 
The main conclusion was that the LNC formulated in this study compared 
advantageously with other nanoparticulate systems, particularly for their residence 
time in blood. Nanocapsule uptake by the different organs of rat was evaluated 
24 hrs after intravenous administration. It was shown that LNC deposited mainly 
in the liver and the spleen, but also in the heart, and the results were comparable 
to a liposome reference. 
5. Drug Encapsulation and Release 
5.1. Ibuprofene 
LNC were characterized for their suitability as an ibuprofene delivery device for 
pain treatments.15 After in vitro investigations, ibuprofene- loaded LNC were evaluated 
after intravenous and oral administration in rats. For each system, the carrier 
was evaluated through its potential antinociceptive efficiency. 
We present in Fig. 8, the release of ibuprofene in a phosphate buffer after 
its incorporation in LNC during formulation. For each case, LNC provide high 
ibuprofene loadings (95%). The main feature is an initial burst followed by a 
CP" 80 
W atex beads 
l : < 
r> % V 
- J1 
-          V - : 
  .   ; ' . .   , , .  
: WOK 
Fig. 8. Sizing studies of submicron bubbles. 
distribution is less than one micron and the mean size is less than 1 micron in 
The images below depict a photomicrograph of MRX-815 bubbles alongside a 
photomicrograph of one-micron size latex microbeads. The bubbles are one micron 
in diameter and smaller. We found in our lab that the smallest bubbles are not 
well shown on the light microscopy due to limitations of the imaging technique. 
The sizing profile shows that there are bubbles up to approximately two microns 
in diameter, but more than 70% of the bubbles are smaller than one micron in 
Investigators have demonstrated that ultrasound can be used to generate cavitation 
in an aqueous medium.18 Cavitation research has led to studies involving 
ultrasound-mediated clot lysis at a variety of frequencies.19"23 Furthermore, 
microbubbles and submicron-sized bubbles provide a nucleus at which cavitation 
can occur, thereby lowering the ultrasound energy requirements.24 
While intravenous administration with local application of ultrasound appears 
to be effective for sonothrombolysis in both pre-clinical and clinical models, applications 
using an infusion catheter are also being investigated.25 It is believed 
that submicron-sized bubbles and ultrasound-mediated cavitation are able to 
affect the thrombus architecture by increasing permeability through the thrombus 
matrix, thereby improving accessibility and the penetration of thrombolytic 
enzymes to more efficiently lyse clots. Studies by Francis et a/.,26'27 demonstrated 
that ultrasound alone increased the spacing between fibrin strands in 
clots, presumably improving the penetration of lytic enzymes, such as t-PA, into 
the clot. 
By way of explanation, when bubbles are insonified, these bubbles can oscillate 
in response to the acoustic pressure wave. If driven with a sufficient acoustic 
pressure, the rapid expansion and contraction of the bubble will result in local 
234 Ungeretal. 
velocities at the bubble surface on the order of hundreds of meters per second. 
If the expansion of the bubble is large enough, the bubble will become unstable, 
resulting in the destruction of the bubble into smaller fragments.28 The rapid 
oscillation of the bubble in response to an acoustic pulse is referred to as "cavitation". 
Bubbles undergoing this violent expansion and contraction produce liquid 
jets, local shock waves, and free radicals. Although the exact mechanism is still 
being studied, the effect of cavitating bubbles has been demonstrated to have several 
effects on the surrounding tissues, including the poration of cell membranes 
resulting in enhanced membrane permeability (sonoporation) or the disruption 
of local thrombus. Thus, the combination of ultrasound with microbubbles has 
potential applications in blood clot dispersion and local drug delivery to treat 
cardiovascular disease, cancer, and diseases of the central nervous system. The 
figure below shows individual images from ultra-high speed videomicroscopy of 
a single bubble. The bubble is shown in the resting state on the far left hand side 
of the figure. The bubble expands after the application of the ultrasound pulse, 
then collapses and fragments. The daughter bubbles expand and collapse again, 
leaving behind small nano-sized fragments.29 Localized activation of bubbles with 
ultrasound can be used for a number of different medical applications including 
Whereas in diagnostic ultrasound contrast imaging where there is an r6 dependence 
between size and ultrasound reflection for therapy, it is advantageous to 
have much smaller bubbles. As shown in Fig. 10, when bubbles are cavitated by 
ultrasound, they may undergo a relatively greater increase in the expansion ratio 
ri/ror where r^ is the maximum size for the radius of the bubble after insonation, and 
r0  the initial resting radius.31 The relative expansion with insonation is greatest 
for the smallest diameter submicron-sized bubbles. This conceivably results in a 
more effective cavitational force, and hence more efficient lysis of thrombi. 
Another effect of ultrasound on microbubbles which has the potential to be utilized 
therapeutically is the use of acoustic radiation force to selectively concentrate 
microbubbles at a target site.32'33,34 Microbubbles driven with ultrasound, experience 
radiation force in the direction of ultrasound wave propagation.35 Pulses of 
t % '2 IV \ ,' V y. 
Fig. 9. In the images above, a single 3 (im bubble is shown (far left) in the resting state. 
Insonation with a single pulse of ultrasound energy causes the bubble to expand, collapse, 
and fragment, yielding nanometer-sized fragments. As the bubbles expand and collapse, they 
generate a local Shockwave that can be used therapeutically. Reproduced with permission 
from Chomas et (A., Appl Phys Lett, 2000.30 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 235 
Fig. 10. The relationship between nanobubbles' size at resting state and expansion ratio 
under insonation. Reproduced with permission of D. Patel et a\., IEEE Ultrasonics, Ferroelectrics, 
and Frequency Control. In press. 
many cycles can deflect resonant microbubbles over distances in the order of millimeters. 
Thus, it may be possible to bring microbubbles circulating in the blood 
pool into contact with targeting sites on a blood vessel wall, in a region selected 
by the positioning of the ultrasound beam. This effect has been demonstrated to 
increase the retention of microbubbles at a target site over an order of magnitude.3 6 
In addition to favorable acoustic characteristics, submicron-sized bubbles have 
other potential advantages for therapy, compared with larger-sized microbubbles. 
The smaller bubbles may penetrate a clot more easily and may have better biodistribution 
characteristics for targeting. 
The pictorial representation below (Fig. 11) is the hypothetical mechanism of 
action for MRX-815 bubbles flowing through the vasculature in association with 
Fig. 11. It is hypothesized that when submicron-sized bubbles are injected systemically, 
some will aggregate on the thrombus, and due to their small size, work into the clot. When 
the bubbles cavitate, the kinetic energy disperses the clot, both from its periphery, and due 
to the fact that bubbles are able to penetrate the clot from within. 
236 Ungeretal. 
a thrombus. Ultrasound could cause cavitation of the bubbles, transferring their 
dispersive energy to the clot and dispersing the clot safely and painlessly. Particle 
sizing studies of the effluent from in vitro studies of SMB-assisted sonothrombolysis 
have shown that the particles are submicron in size.37 
The figures below show the experimental set-up used in our lab for a flow 
through phantom for testing sonothrombolysis, and then treatment of a clot in 
the phantom. The clot was exposed to 1 MHz ultrasound and tissue plasminogen 
activator (t-PA), followed by an infusion of MRX-815 microbubbles. As shown in 
the figures, after 40 min of treatment there is near complete resolution of the clot. 
The graph below shows the results from a series of clots exposed to t-PA, 
t-PA + ultrasound and t-PA + ultrasound + MRX-815 bubbles in our lab. Note 
that the greatest reduction of thrombi was in the group exposed to bubbles. 
, - 
Fig. 12. Above a schematic of the experimental set-up: (A) the clot pre-treatment, (B) after 
32 min of treatment, (C) after 40 min of treatment. The clot was 96% dissolved. 
2 50.00 
Z 40.00 
Saline US t-PA t-PA, SMB, SMB, 
Fig. 13. SMB = Bubbles. 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 237 
4. Clinical Studies 
Vascular thrombosis is a major cause of death in industrialized countries, responsible 
for myocardial infarction, stroke and peripheral arterial occlusions.38 In 
addition, deep vein thrombosis (DVT), which afflicts one in twenty Americans during 
their lifetime,39'40 may also be an application for sonothrombolysis. 
ImaRx completed a Phase I/II clinical trial in thrombosed dialysis grafts for 
the purpose of preliminary feasibility and safety for sonothrombolysis treatment 
of clotted grafts. Initial studies in thrombosed dialysis grafts provided a venue to 
evaluate the principle of sonothrombolysis in vascular thrombosis. As such, clinical 
trial efforts will move forward to address the treatment of stroke, peripheral arterial 
occlusions (PAO) and deep vein thrombosis (DVT). 
Below are examples shown from clinical trials for sonothrombolysis in dialysis 
grafts and DVT. The examples are not an indication that all sonothrombolysis 
treatments will have similar outcomes. 
Images from a venogram in a patient with DVT showed that the patient was 
administered bubbles via infusion catheter into the popliteal vein over a period of 
1 hr, while ultrasound was applied across the skin. No thrombolytic drug such as 
t-PA was administered. Clinically, this particular patient had marked reduction in 
pain post-treatment with sonothrombolysis. 
Stroke is the third most common cause of death, after heart disease and cancer 
in North America. It incurs far more expenses than any other diseases due to its 
long term disability.41 In the US, stroke accounts for over $50 billion each year 
to the health care system.42 The only approved pharmacologic therapy to help 
restore blood flow in stroke patients is t-PA (Activase). Less than 5% of patients 
are treated with t-PA due to concerns over bleeding and the risk relative to the 
benefit.43 Encouraging results have been obtained, however, in human studies with 
ultrasound and t-PA, and most recently, with ultrasound + t-PA + microbubbles. 
Fig. 14. The j:iyio.i a:n on the left is of a clotted dialysis graft. Very little contrast enters the 
graft as it is filled with clot. The image on the right, post-bubble treatment, shows complete 
opacification of the graft due to successful dissolution of thrombosis by sonothrombolysis. 
238 Ungeretal. 
Fig. 15. On the pre-treatment image (left), there is complete occlusion of the superficial 
femoral vein (SFV). Collateral veins are seen carrying the blood flow that would normally 
be carried by the SFV. Post-treatment, there is good flow in the SFV and much less flow is 
seen in the collateral vessels due to the increased flow in the SFV. 
Dr. Andrei Alexandrov from the University of Texas in Houston led a study of 
ultrasound + t-PA in acute ischemic stroke.44,45 In this study, 126 patients were randomized 
prospectively to receive either a 1 hr infusion of t-PA at a dose of 0.9 mg/kg 
alone, or t-PA plus 2 hrs of continuous trans-cranial Doppler (TCD) ultrasound 
applied through the temporal window where the skull is thinnest and most easily 
penetrated by ultrasound. Of the 63 patients treated with t-PA alone, there was a 
13% recanalization rate of the intra-cranial circulation at 2 hrs.46,47 In the same number 
of patients receiving t-PA + ultrasound, there was a highly significant increase 
in recanalization to 38% at two hours, indicating that ultrasound-mediated therapy 
aided in thrombus dispersion. 
Dr. Carlos Molina, from Barcelona, Spain, conducted a similar study but with 
microbubbles.48 The addition of microbubbles enhances the cavitational nuclei 
with a decrease in power requirements. Dr. Molina's study demonstrated that 
the recanalization rate increased impressively to 55%.49 In this study, Dr. Molina 
administered three doses of Levovist, a microbubble agent comprised of air-filled 
galactose microparticles. Dr. Molina's pioneering work has demonstrated the utility 
of using bubbles in conjunction with ultrasound to improve the clinical outcome 
of acute stroke. ImaRx is currently moving MRX-815 into stroke treatment 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 239 
SMB-assisted sonothrombolysis therapy could move beyond the current clinical 
regimens by eliminating the thrombolytic agent. Pre-clinical trials in both canine 
and porcine models have been encouraging.50'51-52 Human studies will be conducted 
to determine if lipid-coated bubbles will improve recanalization rates in patients 
treated with this new ultrasound-mediated paradigm. 
5. Blood Brain Barrier 
Poor transport into the CNS is an obstacle to effectively treat diseases including 
brain tumors, Alzheimer's and other neuro-degenerative diseases. There are two 
principal barriers to drug transport into the CNS: (a) the blood brain barrier (BBB) 
and (b) the ABC transporters, ABCC1 and ABCB1. 
Unlike the rest of the body, the capillary foot processes of the cerebral endothelial 
cells are tight, preventing peptides and macromolecules from leaking through 
to the brain.53 Although the BBB may be permeable to selected ions and small 
molecules, ABCB1, also known as the P-glycoprotein, acts to remove the molecules 
by a drug-efflux system before they enter the brain. Several different strategies have 
been developed to overcome these limitations.54 One approach to drug delivery to 
the brain is by the transient opening of the BBB. 
Hypertonic solutions containing mannitol, which act by shrinking the endothelial 
cells when co-administered with drugs, have been shown to result in enhanced 
cerebral drug uptake.55,56 However, to cause minimum side effects, it is essential for 
the therapy to be regional and localized. Recently, Hynynen et al.57 have shown that 
the BBB can be transiently opened using ultrasound and microbubbles (Illustrated 
in Fig. 16). When bubbles were administered intravenously and focused ultrasound 
was applied across the intact skull, the BBB could be reversibly opened, permitting 
passage of hydrophilic low molecular weight molecules such as gadolinium-DTPA, 
and macromolecules such as fluorescently labeled albumin (Fig. 17) into the CNS.58 
The permeability resolved over a period of hours without damage to the neurons. 
Similar studies have been performed in a porcine model showing that nonfocused 
ultrasound with microbubbles can be used to open the BBB.59 Figure 18 
shows increased dye deposition in the cerebral tissue. 
Introduction of microbubbles as the cavitation nucleus prior to the application 
of ultrasound, lowered the energy needed to open the BBB, thereby lowering 
the bioeffects of ultrasound.60 Using this technique, large biomolecules such as 
horseradish peroxidase (a 40 kDa protein) have been shown to pass through the 
BBB with minimal damage to the brain tissues.61 
It can be envisaged that drugs (small or macromolecules) bound to the 
microbubbles would function as a more efficient drug delivery vehicle, since these 
240 Unger et al. 
/ , 
0 <(S 
/ / 
' < * 
Fig. 16. Cartoon representation of hypothesized ultrasound mediated drug delivery to the 
brain. (A) Cerebral capillaries with tight endothelial junctions prevent passage of molecules 
(including microbubbles and nanoparticles) into the brain. (B) Ultrasound is applied to 
the skull through the temporal window where the skull is thinnest (inset), cavitating the 
microbubbles and opening up the endothelial junctions. (C) Therapeutic agents may now 
pass through the opened junctions. 
4.7 MPa 
2.3 MPa 
3.3 MPa 
1.0 MPa 
Fig. 17. Tl-weighted MR images of rabbit brain after treatment shows contrast enhancement 
at 4 locations (arrows), coronal image across focal plane. Reproduced with permission 
from Hynynen et ah, Radiology. 
would provide the cavitation nuclei and the drug payload in one entity, circumventing 
the co-administration of drug and microbubble. In such instances, the drug 
could be (a) bound to the lipid membrane (hydrophobic drugs), (b) bound to the 
charged lipids on the surface (gene delivery), or (c) buried in the interior in an oily 
layer of a droplet (hydrophobic drugs) (Fig. 19). Furthermore, (d) these drug loaded 
bubbles or droplets may have the potential to be targeted to a specific site in the 
brain by surface ligands. 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 241 
ug/g tissue 
Untreated Ultrasound Untreated Ultrasound + MB 
Fig. 18. Control pigs and pigs treated with ultrasound alone showed no difference in Evan's 
blue uptake. There was a significant difference in uptake when microbubbles were used in 
conjunction with ultrasound. Adapted from Porter et ah,} Am Soc Echocardiogr. 
Fig. 19. Different ways that bubbles or droplets may be able to transport drugs. Drugs may 
be (a) bound or embedded in the lipid membrane, (b) bound to the surface charges of the 
phospholipid membrance (c) buried in the oil in a droplet (d) targeting ligands can be incorporated 
onto the membrance. 
This technology of activation with ultrasound and microbubbles has the potential 
to also be used in the drug discovery process. By exposing cultured neurons 
to drugs, ultrasound and bubbles, high concentrations of the drug may be able to 
deliver to the cells without damaging them. This can potentially be used to screen 
neurons for new therapeutic compounds. 
242 Unger et al. 
Potential CNS diseases amenable to treatment with 
submicron bubble delivery and classes of drugs 
Disease Drugs 
Alzheimer's Disease and 
other neurodegenerative 
disease, seizures and 
psychiatric disorders 
Primary and Secondary 
(metastases) Brain Tumors 
Stroke, brain ischemia 
Infection, e.g., AIDS 
Low molecular weight therapeutics with poor 
delivery to CNS, proteins, gene-based therapeutics. 
Low molecular weight therapeutics with poor 
delivery to CNS, proteins, genetic drugs. Radiation 
Cavitation nuclei to augment sonothrombolysis, 
either with or without use of thrombolytic 
agent. Delivery of oxygen with microbubbles. 
Improvement of cerebral perfusion with 
microbubble-enhanced sonication. Delivery of 
anti-oxidants and growth factors. 
Delivery of anti-infectives, anti-retrovirals to 
6. Drug Delivery 
In the foregoing sections, we discussed activating the bubbles or using them 
in conjunction with ultrasound-mediated processes (e.g. microbubble mediated 
sonothrombolysis to enhance the local activity of the drug such as t-PA), or that 
the availability of a drug may be increased, e.g. by opening the blood brain barrier. 
In this section, we will discuss evaluating drug-carrying microbubbles for drug 
6.1. Targeted bubbles 
As preliminary studies to demonstrate feasibility of using targeted bubbles as 
potential drug delivery agents, two different targeted bubbles were prepared 
using a mixture of DPPC, DPPE-PEG5000 and DPPA, as well as different oils 
and perfluorocarbons using a mixture of DDFP and n-perfluorohexane. In one 
study, a bioconjugate ligand targeted to the am, An integrin was synthesized by 
solid phase peptide methodology.62 Briefly, the bioconjugate, lipids, biocompatible 
drug, perfluoropropane were combined into a mixture and bubbles prepared 
by shaking the vials at approximately 4200 rpm. The size of the targeted bubbles 
Lipid-Coated Submicron-Sized Particles as Drug Carriers 243 
Fig. 20. Intravital microscopy demonstrating adherence of targeted microbubble to thrombus. 
Picture on the right is a graphic representation outlining the location of bound microbubbles 
on thrombus. Reproduced with permission from Schumann et ah, Investi Radiol. 
was approximately 2 fim, as measured by light obscuration measurements on a 
Particle Sizing Systems Model 470 sizer (Particle Sizing Systems, Santa Barbara, 
Calif.). Bubbles were injected into a mouse model where thrombi were previously 
formed in the cremasteric arterioles and venules. Fluorescent imaging revealed 
binding of the targeted bubbles to the thrombi in both arterioles and venules. 
Figure 20 demonstrates the utility of a targeted bubble. 
Similarly, targeted bubbles were used in a HUVEC cell culture model. Briefly, 
bubbles with a targeting ligand directed to a^ft receptors on HUVEC cells were 
 1 /xm), which allows them to 
be administered intravenously without any risk of embolization. According to the 
process and the composition used in the preparation of nanoparticles, nanospheres 
256 Gref & Couvreur 
Fig. 1. (A) Schematic representation of the nanocapsule structure; (B) Morphological 
appearance of a nanocapsule with an oily core (transmission electron microscopy after freeze 
or nanocapsules can be obtained. Nanospheres are matrix systems in which the drug 
is dispersed within the polymer throughout the particle. Contrarily, nanocapsules 
are vesicular or "reservoir" (heterogenous) systems, in which the drug is essentially 
confined to a cavity surrounded by a tiny polymeric membrane (Fig. 1). As in the 
case of nanospheres, depending on their physicochemical properties and composition, 
the drug may adsorb onto the surface as well as being included in the central 
core of nanocapsules. Therefore, drug localization is an important parameter in the 
characterization of nanocapsule preparations. 
The nanocapsule core may be acqueous or composed of a lipophilic solvent, 
usually an oil. In order to achieve good drug loading, the core materials are chosen 
among the good solvents for the drug.1 Expected advantages of confining the drug 
within a central cavity are: (a) burst effect may be avoided; (b) the drug is not 
in direct contact with tissues and therefore irritation at the site of administration 
could be reduced, and (c) the drug may be better protected from degradation both 
during storage and after administration. One of the advantages of nanocapsules 
over nanospheres is their low polymer content and a high loading capacity for 
lipophilic drugs. 
Nanocapsules can either be obtained by interfacial polymerization of 
monomers or from preformed polymers. In the former, the molar mass of the coating 
polymer will depend on the preparation conditions and even on the drug used, 
whereas in the latter, it is determined at the outset. Polymerization of monomers 
may lead to a covalent linkage between the polymer and the drug. To date, all the 
methodologies described for preparing nanocapsules involve the preparation of 
emulsions. Oil-in-water (O/W) emulsions lead to the formation of nanocapsules 
with an oily core, suspended in water. Water-in-oil (W/O) emulsions lead to the 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 257 
obtention of nanocapsules with an acqueous core, suspended in oil. More recently, 
nanocapsules with an acqueous core suspended in an acqueous medium were also 
Nanocapsule technology and their pharmaceutical applications will be further 
discussed according to the method of obtaining the polymeric wall (polymerization 
in situ or preformed polymer) and whether the core is acqueous or oily. 
2. Preparation 
2.1. Nanocapsules obtained by interfacial polymerization 
The advantage of obtaining nanocapsules by interfacial polymerization is that the 
polymer is formed in situ, allowing the polymer membrane to follow the contours of 
the inner phase of an O/W or W/O emulsion, thus entrapping drugs with high loadings. 
However, because reactive monomers are used, unwanted chemical reactions 
may occur between the drug and the monomer, before or during the polymerization 
The preparation of nanocapsules by polymerization requires a fast polymerization 
of the monomers at the interface between the organic and the acqueous phase of 
the emulsions. Alkylcyanoacrylates, which polymerize within seconds, have been 
proposed for the preparation of both oil- and water-containing nanocapsules. Their 
polymerization is initiated by hydroxyl ions either from the equilibrium dissociation 
of water or by nucleophilic groups of any compound in the polymerization 
2.1.1. Oil-containing nanocapsules 
The oil-containing nanocapsules are suitable for the encapsulation of the lipophilic 
and oil-soluble compounds. They are generally obtained by interfacial polymerization 
of alkylcyanoacrylates, after preparing a very fine oil-in-water emulsion 
with an additional water-miscible organic solvent such as ethanol or acetone.3'4 
These solvents serve as vehicles for the monomers, and also help to disperse the 
oil as very small droplets in the acqueous phase, which contains a hydrophilic 
surfactant. Indeed, as pointed out by Gallardo et al.,5 the organic solvents must 
be completely water-miscible, so that the formation of small enough oil droplets 
occurs spontaneously, while the solvent is diffusing towards the acqueous phase 
and the water is diffusing toward the organic phase. Meanwhile, the polymerization 
of the monomer induced by the contact with hydroxyl ions from the water phase 
must be swift to allow efficient formation of the polymer envelope around the oil 
droplet, thus achieving effective encapsulation of drugs. Generally, particles with 
258 Gref & Couvreur 
sizes ranging between 250 and 300 nm, depending on the experimental conditions, 
were obtained.5,6 
In a general procedure of nanocapsule preparation, the oil, the monomer, and 
the biologically active compound are dissolved together in the water-miscible 
organic solvent to prepare the organic phase.3-9 This organic phase is then injected 
via a cannula, under strong stirring, into the acqueous phase containing water and 
a hydrophilic surfactant. The nanocapsules are formed to give a milky suspension 
immediately. The organic phase is then removed under reduced pressure and the 
nanocapsules are purified by ultracentrifugation. Depending on the density of the 
oil forming the core, nanocapsules will concentrate either as a pellet at the bottom 
of the ultracentrifuge tubes or as a floating layer at the top of the tubes. 
A wide range of oils is suitable for the preparation of nanocapsules, including 
vegetable or mineral oils and pure compounds such as ethyl oleate and benzyl 
benzoate. The criteria for selection are the absence of toxicity, lack of affinity for 
the coating polymer, the absence of risk of degradation of the polymer, and a high 
capacity to dissolve the drug that is entrapped. Generally, Miglyol is used to 
form the core of the nanocapsules.3-7,9,10 Lipiodol and benzyl benzoate have also 
been successfully used to form nanocapsules.4 Soluble surfactants were chosen 
among Poloxamers,3-9 Triton X1009 and Tween 80.9 In some cases, nanospheres 
formation together with nanocapsules were observed. Aprotic, fully water-soluble 
solvents such as acetone and acetonitrile lead to high-quality nanocapsule preparations, 
whereas protic water-miscible solvents including ethanol, n-butanol, 
and isopropanol promoted the formation of nanospheres during nanocapsule 
preparation.5,9 It has been hypothesized that alcohols potentially initiate the polymerization 
reaction of alkylcyanoacrylates to form polymer nuclei or preformed 
polymers that may precipitate as nanospheres, when the organic phase is added to 
the acqueous phase.5 Lowering the pH in the organic phase was shown to inhibit 
polymerization in this medium.6 
Oil-containing nanocapsules have been used to encapsulate several types 
of biologically active compounds including both lipophilic molecules such as 
carbamazepine, indomethacin, lomustine, ethosuccimide, phenytoin,1,10-14 and 
hydrophilic drugs such as peptides.15-18 The lipophilic drugs were solubilized in the 
organic phase and were encapsulated during the preparation of the nanocapsules, 
usually using ethanol as the water-miscible organic solvent.4,17 The encapsulation 
efficiency of lipophilic drugs was found to be related to their solubility in the encapsulated 
oil.1 Quite surprisingly, hydrophilic compounds such as peptides have also 
been successfully encapsulated in oil-containing nanocapsules. Indeed, these highly 
water-soluble compounds do not tend to dissolve in oil. It has been suggested that 
the extremely rapid polymerization of the alkylcyanoacrylate occurring at the surface 
of the oil droplet limits the diffusion of the peptide towards the acqueous 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 259 
phase, therefore leading to its entrapment in nanocapsules.15 Another explanation 
is that surfactants may form inverse micelles in the oily phase, allowing some dissolution 
of hydrophilic compounds in this phase. Interestingly, in contrast to what 
has been observed with poly(alkylcyanoacrylate) nanospheres,19 peptides do not 
react chemically with the alkylcyanoacrylate monomer during the preparation of 
nanocapsules when ethanol is used. The presence of a large excess of alcohol seems 
to prevent the hydroxyl and amino groups of the peptides from reacting with the 
monomer, thus retaining the biological activity of the entrapped peptides.16-18,20'21 
For example, encapsulated insulin was still recognized by the insulin receptor of 
hepatocytes after nanoencapsulation.15,22 
2.1.2. Nanocapsules containing an acqueous core 
Nanocapsules with an acqueous core are a recent technology developed for the 
efficient encapsulation of water-soluble compounds, which are generally difficult 
to include within nanospheres. They were obtained by interfacial polymerization, 
where the alkylcyanoacrylates monomers were added to a W/O emulsion.23 
Anionic polymerization of the cyanoacrylate in the oily phase was initiated at the 
interface by nucleophiles such as hydroxyl ions in the acqueous phase, leading to 
the formation of nanocapsules with an acqueous core. In a typical procedure (Fig. 2), 
an acqueous phase at pH 7.4, consisted of ethanol and water, was prepared.23 This 
solution was emulsified in an organic phase containing Miglyol and Montane 80. 
The slow addition (4hrs) of the isobutylcyanoacrylate monomer in the organic 
phase under mechanical stirring allowed the polymerization to occur. This typical 
procedure leads to water droplets that are surrounded by a polymer core. The 
* " \ 9^ ' = Monomer 
COOR oo 
Fig. 2. Schematic representation of the interfacial polymerization of cyanoacrylic 
monomers leading to the formation of nanocapsules with an acqueous core. 
260 Gref & Couvreur 
resuspension of the nanocapsules with a mean diameter approximately 350 nm in 
a water phase has been achieved by the ultracentrifugation of the oily suspension, 
with an excess of demineralized water containing a surfactant. After removal of the 
upper oily phase, the nanocapsules pellet was resuspended in water. 
These nanocapsules are very useful for the encapsulation of hydrophilic compounds 
such as oligonucleotides and peptides. In this case, these macromolecules 
are dissolved in the acqueous phase before the interfacial polymerization process 
takes place. For example, encapsulation efficiencies of 50% with an oligothymidylate 
(phosphodiester) and of 81% with a full phosphorothioate oligonucleotide 
(directed against EWS Fli-chimeric RNA) were obtained.23'24 These entrapment differences 
were attributed to possible interactions of the oligonucleotides with the 
oily phase, Montane 80, or to the possible location of the oligonucleotide at the 
water-oil interface which could become saturated.24 
The localization of the oligonucleotide (within the acqueous core or adsorbed 
on the surface) has been investigated through fluorescence quenching experiments 
using fluorescein-labeled oligonucleotide and potassium iodine as an external 
quencher.23 It has been shown that fluorescent oligonucleotides were located in the 
acqueous core of the nanocapsules, surrounded by a polymeric wall, inaccessible to 
the quencher. On the contrary, when the fluorescent-oligonucleotides were free in 
solution, the fluorophores were highly accessible and strong quenching occurred. 
Similar quenching could be obtained with nanoencapsulated oligonucleotides only 
after the hydrolysis of the polymer wall, thus releasing the oligonucleotides. 
Zeta potential experiments have confirmed the localization of oligonucleotide 
in the acqueous core of the capsule.25 Moreover, nanoencapsulated oligonucleotides 
were protected against degradation by serum nucleases.25,26 Phosphorothioate 
oligonucleotides directed against EWS Fli-1 chimeric RNA encapsulated within 
poly(alkylcyanoacrylate) nanocapsules were tested in vivo for their efficacy against 
the experimental Ewing sarcoma in mice after intratumoral administration.24 Intratumoral 
injection of antisense-loaded nanocapsules led to a significant inhibition of 
tumor growth, whereas no antisense effect could be detected with the free oligonucleotide. 
These results were explained on the basis of a good protection of the 
oligonucleotide in the nanocapsules, which may act as a controled release system 
of oligonucleotide within the tumor. 
Salmon calcitonin was also successfully entrapped within poly (butylcyanoacrylate) 
nanocapsules of 300 nm in diameter.27 When the diameter was 
reduced to 50 nm, the encapsulation efficiencies decreased from 50 to 30%. After 
storage at room temperature or at 4C, the nanocapsules retained their size for at 
least 34 months. The encapsulated calcitonin remained stable at 4C for one year. 
Polyalkylcyanoacrylate nanocapsules were also prepared by interfacial polymerization, 
using a microemulsion instead of an emulsion as the template. 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 261 
Microemulsions are spontaneously forming, thermodynamically stable dispersed 
systems having a uniform droplet size of less than 200 nm. As such, they represent 
an interesting system that may be exploited for the preparation of nanocapsules 
too. Practically, a pseudo-ternary phase diagram of a mixture of medium 
chain glycerides (caprylic/capric triglycerides and mono-, diglycerides), a mixture 
of surfactants (polysorbate 80 and sorbitan monooleate) and water was constructed. 
Microemulsion domains were characterized by conductivity and viscosity 
to select systems suitable for the interfacial polymerization of ethyl-2-cyanoacrylate. 
Nanocapsules of 150 nm were obtained in those conditions and they were found to 
be able to encapsulate significant amounts of insulin.28 Size of the capsules may be 
controled, depending on different formulation variables.29 Factors influencing the 
encapsulation of hydrophilic compounds have been identified too.30 
2.2. Nanocapsules obtained from preformed polymers 
The preparation of nanocapsules from preformed polymers avoids some drawbacks 
of the interfacial polymerization process, such as the lack of control of the 
polymer molar masses and polydispersity, the presence of residual monomer in 
the preparation, and the possibility of drug inactivation.31 An interfacial deposition 
process to prepare nanocapsules, also known as nanoprecipitation, has been 
developed.32,33 In this simple and reproducible method, a water-miscible organic 
phase such as an alcohol or a ketone containing oil (with or without lipophilic 
surfactant) is mixed with an acqueous phase containing a hydrophilic surfactant. 
The preformed polymer, insoluble in both the oily and the acqueous phase, 
is solubilized in the organic phase. After the addition of the organic phase to 
the acqueous phase, the polymer diffuses with the organic solvent towards the 
acqueous phase and is stranded at the interface between oil and water. The driving 
force for nanocapsule formation is the rapid diffusion of the organic solvent 
in the acqueous phase, inducing interfacial nanoprecipitation of the polymer surrounding 
the droplets of the oily phase. Synthetic polymers such as poly(D,Llactide), 
poly(e-caprolactone) and poly(alkylcyanoacrylate) are most frequently 
employed for nanocapsule formation.32 Arabic gum, gelatin, ethylcellulose or 
hydroxypropylmethylcellulose phthalate were also successfully used.32 The size 
of nanocapsules is usually found between 100 and 500 nm, and it depends on 
several factors, namely, the chemical nature and the concentration of the polymer 
and the encapsulated drug, the amount of surfactants, the ratio of organic 
solvent to water, the concentration of oil in the organic solution, and the speed 
of diffusion of the organic phase in the acqueous phase. In general, the lower the 
interfacial tension and the viscosity of the oil, the smaller the nanocapsules are 
262 Gref & Couvreur 
Both lipophilic and hydrophilic surfactants are used in the preparation of 
nanocapsules by this technique. However, not all the surfactants that are technically 
suitable are acceptable for parenteral administration; as such, the choice 
has to be made with the administration route in mind. Generally, the lipophilic 
surfactant is a natural lecithin of relatively low phosphatidylcholine content, 
whereas the hydrophilic one is ionic (i.e. lauryl sulphate, quaternary ammonium), 
or more commonly nonionic (i.e. poly(oxyethylene)-poly(propropylene) 
Poly(ethylene glycol)-coated nanocapsules were also prepared by nanoprecipitation, 
using preformed diblock poly(lactide)-poly(ethylene glycol) copolymers 
or blends of these copolymers with the homopolymer poly(lactide.)35-38 
However, the most physically stable nanocapsules were those prepared with 
poly(lactide)-poly(ethylene glycol) copolymer alone. RU 58668, a promising pure 
antiestrogen, was entrapped into poly(ethylene glycol)-coated nanospheres and 
into nanocapsules with a similar coating.37 A series of preformed diblock polyesterpolyethylene 
glycol) copolymers were used for the design of these nanoparticles, 
both the molar masses of the poly(ethylene glycol) blocks and the nature 
of the hydrophobic polyester blocks being varied. Nanospheres which had a 
smaller size (~110nm), compared with nanocapsules (~250nm), were however 
able to incorporate larger amounts of the antioestrogen than the nanocapsules 
In an alternative method named solvent displacement method, an O/W emulsion 
was formed.39 The organic phase contained the polymer, the oil and the drug, 
and the acqueous solution contained a stabilizing agent. In this procedure, the 
organic solvent was displaced into the external phase by the addition of an excess of 
water. This technique has several advantages such as the small quantities of solvents 
used, the good control of the size of the nanocapsules (80-900 nm), and the control 
of the thickness of the polymeric wall by monitoring the polymer concentrations.40 
However, large amounts of water have to be removed at the end of the 
Two formulation processes which bring lipids into play should also be mentioned. 
The first methodology is based on the inversion phase of an emulsion 
to prepare original lipidic nanocapsules. These capsules, interestingly obtained 
as a suspension in saline water, were constituted by medium chain triglycerides 
and hydrophilic /lipophilic surfactants. According to the authors, the formulation 
method has been developed to avoid the use of organic solvent or the high quantity 
of surfactants and co-surfactants, due to the potential toxicity of their residues after 
human administration. Their original structure was found to be a hybrid between 
polymeric nanocapsules and liposomes as their oily core is being surrounded by a 
tensioactive rigid membrane.41-43 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 263 
In another process, cisplatin lipid-based nanocapsules have been prepared by 
the repeated freezing and thawing of an equimolar dispersion of phosphatidylserine 
(PS) and phosphatidylcholine (PC) in a concentrated acqueous solution of 
cisplatin. Here, the molecular architecture of these novel nanostructures was elucidated 
by solid-state NMR techniques.15N NMR and 2H NMR spectra of nanocapsules 
containing 15N- and 2H-labeled cisplatin respectively, demonstrated that the 
core of the nanocapsules consists of solid cisplatin devoid of free water. Magicangle 
spinning 15N NMR showed that approximately 90% of the cisplatin in the 
core is present as the dichloro species. The remaining 10% was accounted for by 
a newly discovered dinuclear Pt compound that was identified as the positively 
charged chloride-bridged dimer of cisplatin. NMR techniques, sensitive to lipid 
organization 31P NMR and 2H NMR, revealed that the cisplatin core is coated by 
phospholipids in a bilayer configuration and that the interaction between solid 
core and bilayer coat exerts a strong ordering effect on the phospholipid molecules. 
Compared with phospholipids in liposomal membranes, the motion of the phospholipid 
headgroups is restricted and the ordering of the acyl chains is increased, 
particularly in PS.44 Analysis of the mechanism of the nanocapsule formation suggests 
that the method may be generalized to include other drugs showing low water 
solubility and lipophilicity.45 
3. Characterization 
Size evaluation of nanocapsules is most frequently done by photon correlation 
spectroscopy, transmission electron microscopy, and scanning electron microscopy, 
without or after freeze-fracture.33,39,46 At present, transmission electron microscopy 
performed after freeze-fracture has given the most useful information about 
nanocapsule structure, highlighting the polymer envelope and the inner cavity, 
and allowing the wall thickness to be estimated.1'7,47 Thus, polymer coatings were 
estimated to be around 5 ran, depending on the monomer concentration.47 Freezefracture 
(Fig. 1) has also allowed the visualization of different possible organizations 
of lipophilic surfactant, which can form vesicles, micelles, bilayers, or monolayers, 
depending on its concentration.33 The spherical shape of the nanocapsules was 
confirmed by atomic force microscopy.39 Most images of nanocapsules have been 
obtained by transmission electron microscopy performed on negatively stained 
preparations, allowing to gain information about nanocapsule morphology and 
integrity1,47 (Fig. 3A). Nanocapsules embedded in a suitable resin were cut into 
thin slices.48 They were observed using electron microscopy, the contrast being 
created by encapsulation of a colloidal gold-labeled molecule during nanocapsule 
preparation. In this manner, both polymer envelope and the internal cavity were 
distinguished easily (Fig. 3B). 
264 Gref& Couvreur 
50nm 100 nm 100 nm 
B 3 l 
Fig. 3. (A) Morphological appearance of polydactic acid-co-glycolic) nanocapsules using 
the transmission electron microscopy. (B) Labeling insulin with gold allows to distinguish the 
localization of this molecule into the internal core of poly(isobutyl cyanoacrylate) nanocapsules; 
Transmission Electron Microscopy. 
Zeta potential measurements are also very useful for the chraracterization of 
the nanocapsules. Surfactants and polymer are the major components that can affect 
this parameter. Many polymers such as poly (D,Llactide), poly(e-caprolactone) and 
lecithins impart a negative charge to the surface, whereas nonionic surfactants such 
as Poloxamer tend to reduce the absolute value of zeta potential.34 Calvo et alP 
described nanocapsules coated with positively charged polysaccharide chitosan. 
Their surface charge depended mainly on the viscosity of the chitosan solution used 
for coating. Positive values up to 46 mV were also observed with diethylaminoethyldextran 
coated nanocapsules.8 Generally, Zeta potential values above 30 mV (positive 
or negative values) lead to more stable nanocapsule suspensions, because 
repulsion between the particles prevented their aggregation. In contrast to observations 
with nanospheres, the negative Zeta potential of the nanocapsules was 
not completely masked by the presence of neutral poly(ethylene glycol) chains at 
the surface.63 This was due to the presence of lecithin in the polyethylene glycol) 
"brush", which remained necessary for nanocapsule stability. It was further highlighted 
that the presence of such a "brush" could reduce complement activation, 
an important step in the recognition of particles by macrophages.50'51 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 265 
Centrifugation in a density gradient was used to confirm the existence of 
nanocapsules by comparing with the colloidal carriers prepared without polymer 
or oil. For example, isopycnic centrifugation in a density gradient of Percoll 
was used in the case of nanocapsules with a Miglyol core and a coating of 
poly (alky lcyanoacry late) or poly(D,L lactide).39 The density of the nanocapsules 
was found to be intermediate between that of nanospheres and that of emulsions. 
These studies also demonstrated that the density of nanocapsules and the band 
thickness increased when the quantity of polymer increased. No contamination of 
nanocapsules with nanospheres was observed. However, Mosqueira et al.3i performed 
similar experiments and observed that nanocapsule preparations obtained 
by nanoprecipitation contained small amounts of nanospheres, as it has previously 
been described by Gallardo et al.5 for nanocapsules prepared by interfacial 
polymerization. When lecithin was present in excess as lipophilic surfactant, liposomes 
were also detected in the nanocapsule preparations. Liposomes could not 
be distinguished from nanocapsules on the basis of density differences, but have 
been detected by electron microscopy52 and by the encapsulation of an acqueous 
4. Drug Release 
Release of encapsulated drugs from nanocapsules made of preformed polymers, 
appears only to be controled by the partition coefficient of the drug between the 
oily core and the acqueous external medium, and the relative volumes of these 
two phases. Except for macromolecules, the rate of diffusion of the drug through 
the thin polymeric coating does not seem to be a limiting factor, nor does the 
nature of the polymeric wall. This clearly suggests that the polymer membrane 
may be porous rather than a continuous film barrier to diffusional release. The 
nature of the external acqueous phase is of prime importance in the release. For 
example, indomethacin release was faster and more complete in the presence of 
albumin, which acts as an acceptor in the acqueous phase.11,52 Similarly, release 
of halofantrine, a highly lipophilic drug, was only observed in the presence of 
serum, because the drug has a high affinity for lipoproteins.36 The presence of 
a hydrophilic poly(ethylene glycol) "brush" at the nanocapsule surface was also 
shown to play a role in drug release. Release of halofantrine and primaquine from 
such surface-modified nanocapsules was reduced, compared with conventional 
In conclusion, it may be considered a challenge to develop nanocapsule systems 
with release profiles, which may be controled not only by the partitioning 
coefficient, but also by the nature or morphology (i.e. thickness or porosity) of the 
surrounding membrane. 
266 Gref & Couvreur 
5. Applications 
Nanocapsules have been proposed as drug delivery systems for several drugs by 
different routes of administration such as oral, ocular or parenteral. Drug-loaded 
nanocapsules were used to improve the stability of the drug either in biological 
fluids, or simply in the formulation. Another goal was to reduce the toxicity of 
some drugs known for their undesirable side effects. 
5.1. Oral route 
Challenging aspects related to oral administration deal with the entrapment of 
unstable molecules, such as peptides or that of anti-inflammatory compounds that 
cause local side effects on the mucosae. Pioneering studies in the mid 1980s dealt 
with indomethacin and insulin entrapment. 
Indomethacin, an anti inflammatory drug, has been successfully encapsulated 
in the polyalkylcyanoacrylate nanocapsules with the aim of reducing its side effects 
on the gastric and intestinal mucosa.11 The drug retained its biological activity after 
nanoencapsulation. Moreover, nanoencapsulated formulations allowed a dramatic 
reduction of the ulcerative side effects usually induced by indomethacin on the 
mucosae.54 This protection was attributed to the combined effect of the sustained 
release of indomethacin from the nanocapsules, with a significant reduction of the 
direct contact between drug and the mucosae. In the case of nanocapsules obtained 
by nanoprecipitation using polyesters, the release kinetics in media mimicking 
pH of the gut were more sensitive to changes in drug partitioning related to the 
change of pH, than to the type of polymer used.55,56 Drug release from nanocapsules 
was accelerated in the presence of digestive enzymes such as proteases and 
esterases. This was correlated with a decrease in polymer molecular weight.55'56 
Diclofenac and indomethacin, two major nonsteroidal anti inflammatory agents, 
have been encapsulated in polyQactic acid) nanocapsules obtained by nanoprecipitation, 
with the aim of reducing their side effects on the gastric mucosa.54,57,58 The 
side effects of both drugs were completely modified and reduced by the encapsulation 
in nanocapsules.54 As in the case of nanocapsules produced by interfacial 
polymerization, a marked protective effect on the gastrointestinal mucosa, as compared 
with the ulcerative effect observed with the drug solutions, was observed. 
Insulin-loaded nanocapsules yielded promising pharmacological results.16,21 
When given orally to diabetic rats and dogs, single administration produced 
a reduction in glycemia after an unusually long lag of several days, and this 
hypoglycemia was sustained for up to 20 days.16,20,21,59 It was suggested that 
nanocapsules could release insulin slowly from a depot within the body. The 
nanocapsules seemed to be involved in carrying the insulin near the intestinal 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 267 
epithelium where they were absorbed and translocated as intact nanocapsules to the 
blood vessels.48,59-61 However, Lowe and Temple16 reported that insulin adsorption 
from orally administered nanocapsules reached a maximum of absorption, 15 min 
after administration and any trace of insulin in blood was detected after a few 
hours. Sai et al.62 have proposed the use of insulin-loaded nanocapsules as a new 
prophylactic tool to prevent diabetes. They showed in a model of non-obese diabetic 
mice that prophylactic injection of such nanocapsules reduced the incidence of 
Anti infectious agents such as atovaquone and rifabutin, two compounds active 
against the opportunistic parasite Toxoplasma gondii, were successfully entrapped in 
poly(lactide) nanocapsules formed by nanoprecipitation. These drugs have a poor 
bioavailability because of their insolubility in water. Nanoencapsulation is allowed 
to decrease in the brain parasitic burden in a higher extent than the same amount 
of free drug.63 
Chitosan-coated nanocapsules were particularly interesting for oral administration, 
probably because their positive charge allow them to stick efficiently along 
the gastro-intestinal mucosa, with a further possible diffusion through the epithelium 
, thus providing a continuous drug delivery into the blood stream.64,65 When 
the peptide salmon calcitonin was entrapped into these nanocapsules, long-lasting 
hypocalcemia effects were observed, following oral administration to rats.66 In contrast, 
calcitonin control emulsions led to negligible responses. 
5.2. Parenteral route 
As far as the parenteral route is concerned, nanocapsules could be useful for the 
formulation of poorly soluble drugs, and for controling the drug biodistribution 
according to the properties of the carrier. In this view, indomethacin and diclofenac 
were entrapped in nanocapsules, but diclofenac in solution or in nanocapsules 
showed similar plasma concentration profiles. After intravenous administration, 
encapsulated indomethacin showed even lower plasma concentrations than the 
free drug because of enhanced hepatic uptake of loaded nanocapsules.57 One possible 
explanation for the absence of the modification of the pharmacokinetics and 
biodistribution profiles of the encapsulated drugs probably results from the rapid 
rate of release of these drugs into the circulation, due to the high blood dilution 
and/or the presence of plasma proteins. Subcutaneous injection did not lead to a 
slow release of the drug either. Nevertheless, after intramuscular administration, 
the nanocapsules containing diclofenac showed a significantly reduced inflammation 
at the site of injection, compared with the free drug in solution.67 Similarly, 
darodipine nanocapsules provided a prolonged antihypertensive effect compared 
with free drug which lasted for at least 24 hrs.68 
268 Gref & Couvreur 
Nanocapsules prepared by interf acial polymerization of the isobutylcyanoacrylate 
monomers were retained longer at the injection site after intramuscular administration 
than the other types of carriers such as emulsions or liposomes.69 Moreover, 
they were taken up to a significant extent by the regional lymph nodes, likely owing 
to the phagocytosis by macrophages. These observations open up the possibility of 
delivering cytostatic drugs and immunomodulators to the lymph node metastases. 
When administered intravenously, nanocapsules made by interfacial polymerization 
or by nanoprecipitation were taken up rapidly by organs of the mononuclear 
phagocyte system, mainly the liver.70 To take advantage of this particular tissue 
distribution, nanocapsules containing muramyltripeptide cholesterol (MTPChol) 
were designed.71,72 This immunostimulating agent, able to activate the 
macrophages and to stimulate their innate defense functions against tumor cells, is a 
useful agent to treat metastatic cancer. In vitro studies with rat alveolar macrophages 
have shown that nanocapsules prepared from poly(D,Llactic acid) containing MTPChol 
were more efficient activators than the free drug. This was attributed to the 
intracellular delivery of the nanoencapsulated immunomodulator after cell phagocytosis; 
an intermediate transfer of the drug to serum proteins was another suggested 
mechanism.73 In vivo, this type of nanocapsules is allowed to obtain significant 
antimetastatic effects in a model of liver metastases.74 
For other types of applications, to avoid the rapid clearance by the mononuclear 
phagocyte system, nanocapsules coated with poly(ethylene glycol) with a 
molar mass of 20,000 g/mole were developed. An antimalarial drug, halofantrine, 
was entrapped with the aim of obtaining a well-tolerated injectable form for the 
treatment of this severe intravascular disease.36 In mice, at an advanced stage of 
infection with Plasmodium berghei, the area under the curve for plasma halofantrine 
was increased six-fold, compared with the free drug when the molecule was presented 
as nanocapsules. Moreover, the toxicity of halofantrine was reduced by 
incorporation into the nanocapsules. Up to 100 mg/kg could be administered intravenously 
without toxicity, yet all mice injected with this dose of free halofantrine 
died instantaneously. However, in vivo, only small differences were observed in 
terms of the therapeutic activity between poly(ethylene glycol) coated nanocapsules 
and the uncoated ones. This was explained by the possible saturation of the 
phagocytic capacity of the liver in severely infected mice, as a result of the uptake of 
parasitized erythrocytes.75 Moreover, it was emphasized that the amount of serum 
lipoproteins, which acted as acceptors for halofantrine released from nanocapsules, 
is reduced during the disease. 
Poly(ethylene glycol) coated nanocapsules were also used to deliver lipophilic 
drugs to the solid tumors. In this case, the vascular endothelium is known to be 
more permeable, thus allowing the extravasation of small-sized colloidal particles. 
This specific distribution of colloids into tumoral sites is known as the enhanced 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 269 
permeability and retention effect (EPR effect). The efficacy of this strategy has been 
demonstrated using a photosensitizer, meta-tetra (hydroxyphenyl) chlorine, encapsulated 
in nanocapsules designed from diblock poly(D,L lactide)-poly(ethylene glycol) 
5.3. Ocular delivery 
The major problems encountered when delivering drugs to the eyes are the poor 
permeability of the corneal epithelium and the rapid clearance because of tear 
turnover and lacrimal drainage. Nanocapsule formulations were developed with 
the aim of improving drug efficacy by retaining it at the level of the ocular tissue, 
thus reducing the number of administrations.7677 
Betaxolol-loaded poly(isobutylcyanoacrylate) nanocapsules made by interracial 
polymerization were prepared for the treatment of glaucoma. Only a marginal 
decrease in the intraocular pressure was observed with this type of formulation, 
compared with the activity obtained with the commercial form (single solution) or 
by other carriers.13 More promising results have been obtained with pilocarpine.14 
In this case, sustained drug release was obtained when incorporating the pilocarpine 
loaded nanocapsules into a Pluronic gel. Thus, a significant increase in the 
bioavailability of the drug was achieved. 
Ganciclovir is an antiviral drug used for the treatment of cytomegalovirus infections. 
In the clinical practice, two to three intravitreal injections per week are needed 
to overcome the rapid clearance of the drug from the eyes. Ganciclovir encapsulation 
in poly(ethylcyanoacrylate) nanocapsules made by interfacial polymerization 
provided a sustained release of the drug over four days.10 Moreover, after intravitreal 
injection of the nanocapsules, the drug could still be detected in the eyes at a 
therapeutic level after ten days. Significant amounts of ganciclovir were found in 
the retina and in the vitreous humor which is considered as beneficial in the treatment 
of cytomegalovirus retinitis. On the contrary, after administration of single 
solutions of the drug free, the maximum concentration of ganciclovir was reached 
in less than one day and no drug could be detected later. However, despite these 
beneficial results, some toxicity (opacification of the lens and vitreous humor turbidity) 
was found as a result of the nanocapsules. 
Antiglaucomatous agents such as carteolol and betaxolol were also encapsulated 
in nanocapsules prepared from preformed polymers, but they only showed 
a reduction of the noncorneal absorption (systemic circulation), leading to lesser 
side effects as compared with the free drug.13'78'79 Encapsulation in nanocapsules 
produced an improved pharmacological effect characterized by a more important 
reduction of the intraocular pressure, compared with the free drug treatment, 
as well as with the same treatment but delivered by nanospheres; reduced 
270 Gref & Couvreur 
cardiovascular systemic side effects were also observed with the nanocapsules. ' 
In the case of betaxolol, the nature of the polymer making up the nanocapsule 
wall was found to play a major role in the pharmacological responses.78'80 Thus, 
poly(e-caprolactone) walls were more efficient than poly(isobutylcyanocrylate) 
or poly(lactide-co-glycolide) ones. Indeed, as shown by the confocal microscopy, 
poly(e-caprolactone) nanocapsules could specifically penetrate the corneal epithelium 
by an endocytic process, without causing any damage to the cells. In contrast, 
poly(isobutylcyanoacrylate) nanoparticles produced a cellular lysis.81 As no differences 
in penetration were observed between nanospheres and nanocapsules, 
the presence of an oily core did not seem to influence activity of the formulation. 
Coating the negatively charged surface of poly(e-caprolactone) nanocapsules with 
chitosan, a cationic polymer, provided the best corneal drug penetration, together 
with preventing the degradation caused by the adsorption of lysozyme, a positively 
charged enzyme found in tear fluid.82 This was explained by the higher penetration 
of the nanocapsules into the corneal epithelial cells and by the mucoadhesion of 
these positively charged particles onto the negatively charged membranes. Additionally, 
a specific effect of chitosan on the tight junctions has been mentioned.83 
Encouraging results were also obtained with nanocapsules containing the 
immunosuppressive peptide cyclosporin A.84 This drug was efficiently entrapped 
in poly(e-caprolactone) nanocapsules, leading to a five-fold increase of the 
cyclosporin A corneal concentrations, compared with an oily solution of the drug. 
Again, chitosan-overcoated nanocapsules were able to provide a selective and prolonged 
delivery of cyclosporine A to the ocular mucosae, without compromising 
the inner ocular tissues and avoiding systemic absorption.84 The mechanism that 
explains the increased ocular penetration was understood as the combination of an 
improved interaction with the corneal epithelium, followed by the penetration of 
the particles into the corneal epithelium.85 In the case of indomethacin associated 
with chitosane-coated nanocapsules, the use of confocal microscopy established the 
fact that the nanocapsules penetrated through the corneal epithelium following a 
transcellular pathway.85,86 
6. Conclusion 
As discussed in this chapter, there are now various technologies for the preparation 
of nanocapsules. These methods which obey a wide variety of principles may either 
start from a monomer or from a preformed polymer. They employ macromolecular 
materials of synthetic or natural origin and they allow the design of nanocapsules 
with either an acqueous or an oily core. Thus, they can efficiently entrap almost 
every molecule. The most significant advantage of nanocapsules over nanospheres 
is that the drug to polymer ratio is generally much higher, which allows the use of 
Nanocapsules: Preparation, Characterization and Therapeutic Applications 271 
lesser polymer to deliver the same amount of drug to the cells and tissues. This is, 
from a toxicological point of view, a substantial advantage of this type of technology. 
On the contrary, drug release from nanocapsules is mainly dependent on the partitioning 
coefficient of the biologically active compound between the nanocapsule 
core and the biological receptor medium. If the nanocapsule thin polymer membrane 
may be a barrier for the diffusion of macromolecules, it is not the case for 
small organic molecules. Thus, to control the drug release kinetic from nanocapsules, 
it is likely to remain the primary challenge to be resolved with this kind of 
technology in the next few years. 
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Dendrimers as Nanoparticulate 
Drug Carriers 
SSnke Svenson and Donald A. Tomalia 
1. Introduction 
The development of molecular nanostructures with well-defined particle size and 
shape is of eminent interest in biomedical applications such as the delivery of active 
pharmaceuticals, imaging agents, or gene transfection. For example, constructs utilized 
as carriers in drug delivery generally should be in the nanometer range and 
uniform in size to enhance their ability to cross cell membranes and reduce the risk 
of undesired clearance from the body through the liver or spleen. Two traditional 
routes to produce particles that will meet some of these requirements have been 
widely investigated. The first route takes advantage of the ability of amphiphilic 
molecules (i.e. molecules consisting of a hydrophilic and hydrophobic moiety) to 
self-assemble in water above a system-specific critical micelle concentration (CMC) 
to form micelles. Size and shape of these micelles depend on the geometry of the 
constituent monomers, intermolecular interactions, and conditions of the bulk solution 
(i.e. concentration, ionic strength, pH, and temperature). Spherical micelles are 
monodisperse in size; however, they are highly dynamic in nature with monomer 
exchange rates in millisecond to microsecond time ranges. Micelles have the ability 
to encapsulate and carry lipophilic actives within their hydrocarbon cores. Depending 
on the specific system, some micelles either spontaneously rearrange to form 
liposomes after a minor change of solution conditions, or when they are exposed 
to external energy input such as agitation, sonication, or extrusion through a filter 
278 Svenson &Tomalia 
membrane. Liposomes consist of bilayer lipid membranes (BLM) enclosing an aqueous 
core, which can be utilized to carry hydrophilic actives. Furthermore, liposomes 
with multilamellar membranes provide cargo space for lipophilic actives as well. 
However, most liposomes are considered energetically metastable, and will eventually 
rearrange to form planar bilayers.1'2 The second route relies on engineering the 
well-defined particles through processing protocols. Examples for this approach 
include (i) shearing or homogenization of oil-in-water (o/w) emulsions or w / o / w 
double emulsions to produce stable and monodisperse droplets, (ii) extrusion of 
polymer strands or viscous gels through nozzles of defined size to manufacture stable 
and monodisperse micro and nanospheres, (iii) layer-by-layer (LbL) deposition 
of polyelectrolytes and other polymeric molecules around colloidal cores, resulting 
in the formation of monodisperse nanocapsules after the removal of the templating 
core, and (iv) controlled precipitation from a solution into an anti-solvent, including 
supercritical fluids. Size, degree of monodispersity, and stability of these structures 
depend on the systems that are being used in these applications.3 These systems 
and their utilization in drug delivery are being discussed in detail in other chapters 
of this book. 
Currently, a new third route to create very well-defined, monodisperse, stable 
molecular level nanostructures is being studied based on the "dendritic state" 
architecture.4 Dendritic architecture is undoubtedly one of the most pervasive 
topologies observed throughout biological systems at virtually all dimensional 
length scales. This architecture is found at the meter scale in tree branching and 
roots, on the centimeter and millimeter scales in circulatory topologies in the human 
anatomy such as lungs, kidney, liver, and spleen, and on the micrometer scale in 
cerebral neurons. On the nanometer level, key examples of dendritic structures 
include glycogen, amylopectin, and proteoglycans. Amylopectins and glycogen 
are critical molecular level constructs involved in energy storage in plants and 
animals, while proteoglycans are an important constituent of connective tissue, 
determining its viscoelastic properties. Upon the analysis of these ubiquitous dendritic 
patterns, it is evident that these highly branched architectures offer unique 
interfacial and functional performance advantages. The objective of this review is to 
study the use of dendrimers in drug delivery applications. Four main properties of 
dendrimers will be discussed: (i) nanoscale container properties (i.e. encapsulation 
and transport of a drug), (ii) nano-scaffolding properties (i.e. surface adsorption 
or attachment of a drug and/or targeting ligand), (iii) dendrimers as drugs, and 
(iv) biocompatibility of dendrimers. In addition, routes of application currently 
investigated will be presented. Particular emphasis will be placed on poly 
(amidoamine) (PAMAM) dendrimers, the first and most extensively studied family 
of dendrimers.4c'5 
Dendrimers as Nanoparticulate Drug Carriers 279 
2. Nanoscale Containers  Micelles, Dendritic Boxes, 
Dendrophanes, and Dendroclefts 
Dendrimers may be visualized as consisting of three critical architectural domains: 
(i) the multivalent surface, containing a larger number of potentially reactive/ 
passive sites (nano-scaffolding), (ii) the interior shells (i.e. branch cell layers defined 
by dendrons) surrounding the core, and (iii) the core to which the dendrons are 
attached. The two latter domains represent well-defined nano-environments, which 
are protected from the outside by the dendrimer surface (nanoscale containers) 
in the case of higher generation dendrimers. These domains can be tailored for 
a specific purpose. The interior is well-suited for host-guest interaction and the 
encapsulation of guest molecules. 
2.1. Dendritic micelles 
Tomalia and coworkers demonstrated by electron microscopy observation that 
sodium carboxylated PAMAM dendrimers possess topologies reminiscent of regular 
classical micelles.4 It was also noted from electron micrographs that a large 
population of individual dendrimers possessed a hollow core. Supporting these 
observations, Turro and colleagues designed a hydrophobic 12-carbon atom alkylene 
chain into the core of a homologous series of PAMAM dendrimers (G = 2, 
3, and 4) to mimic the hydrophobic and hydrophilic core-shell topology of a regular 
micelle. The hosting properties of this series towards a hydrophobic dye as a 
guest molecule were then compared with a PAMAM dendrimer series possessing 
non-hydrophobic cores (e.g. NH3 and ethylenediamine). Dramatically enhanced 
emission of the hydrophobic dye was noted in aqueous solution in the presence 
of hydrophobic versus hydrophilic cored dendrimers.6a Less polar dendrimers 
(i.e. dendrimers containing aryl groups or other hydrophobic moieties as building 
elements), behave as inverse micelles.6b A critical property difference relative 
to micelles is the increased density of surface groups with higher generations. At 
some generational level, the surface groups will reach the so-called "de Gennes 
dense packing" limit and seal the interior from the bulk solution (Fig. I ) . 7 - 9 The 
limit depends on the strength of intramolecular interactions between adjacent surface 
groups, and therefore, on the condition of the bulk solution (i.e. pH, polarity 
and temperature). 
This nanoscale container feature, originally noted for PAMAM dendrimers 
by Tomalia et al. and referred to as "unimolecular encapsulation", can be utilized 
to tailor the encapsulation and release properties of dendrimers in drug delivery 
applications.910 For example, adding up to a limiting amount of Xmmol of 
either 2,4-dichlorophenoxyacetic acid or aspirin (acetylsalicyhc acid) to 1 mmol of 
280 Svenson & Tomalia 
,--oV..> jfc-iiw-'i.. J&:5Sgtfhv ;#88c?pi ;*: 7.:;^- *#.# p f e $$8? 
  J  ^  ' ' ^ ,s&$p?* '%$$? 
4 5 6 7 8 9 10 
Fig. 1. Periodic properties of PAMAM dendrimers generations G = 4-10, depicting the 
decreasing distances between surface charges (Z-Z). The "de Gennes dense packing" appears 
atG = 8. Dendrimers G = 4-6 display "nanoscale container" properties, the larger analogues 
G = 7-10 display "nano-scaffolding" properties. 
STARBURST carbomethoxy-terminated PAMAM dendrimers generations 0.5-5.5 
produced spin-lattice relaxation times (Tj) much lower than the values of these 
guest molecules in solvent without dendrimer. The new relaxation times decreased 
for generations 0.5-3.5, but remained constant for generations 3.5 to 5.5. The maximum 
concentration X varied uniformly from 12 (generation 0.5) to 68 (generation 
5.5). On the basis of these maximum concentrations, the guest-to-host ratios were 
shown to be ~ 4:1 by weight and ~ 3:1 based on a molar comparison of dendrimer 
guest carboxylic acid-to-interior tertiary nitrogen moieties for generations 2.5-5.5. 
Exceeding the maximum concentration X resulted in the appearance of a second 
relaxation time, Tv, characteristic of the guest molecules in bulk solvent phase.10 
2.2. Dendritic box (Nano container) 
Surface-modification of G = 5 poly(propyleneimine) (PPI) dendrimers with 
Boc-protected amino acids induced dendrimer encapsulation properties by the 
formation of dense, hydrogen-bonded surface shells with solid-state character 
("dendritic box").8 Small guest-molecules were captured in such dendrimer interiors 
and were unable to escape even after extensive dialysis. The maximum amount 
of entrapped guest molecules was directly proportional to the shape and size of the 
guest molecules, as well as to the amount, shape and size of the available internal 
dendrimer cavities. Four large guest-molecules (i.e. Rose Bengal) and 8-10 small 
guest-molecules (i.e. p-nitrobenzoic acid) could be simultaneously encapsulated 
within PPI dendrimers containing four large and twelve smaller cavities. Remarkably, 
this dendritic box could be opened under controlled conditions to release either 
some or all of the entrapped guest molecules. For example, partial hydrolysis of 
the hydrogen-bonded Boc-shell liberated only small guest-molecules, whereas total 
hydrolysis released all sizes of entrapped molecules.8'11-12 
Although the "dendritic box" concept demonstrates the unique shapedependent 
cargo space that can be found in certain dendrimers, other parameters 
have to be considered as well for delivering and releasing therapeutic drugs 
Dendrimers as Nanoparticulate Drug Carriers 281 
under physiological conditions. From a thermodynamic perspective, free guestmolecules 
(i.e. drugs) can be distinguished from those encapsulated or bound in a 
complex by finite energy barriers related to the ease of entry and departure to the 
dendrimer cavities. If the drug molecule is incompatible with either the dimension 
or hydrophilic/lipophilic character of the dendrimer cavity, a complex might not 
form, or the guest might only be partially encapsulated within the dendrimer host. A 
hydrophobic drug would be expected to associate with a dendrimer core to achieve 
maximum contact with its hydrophobic domain. In addition, the hydrophobic character 
of this guest molecule would be expected to isolate itself from the dendrimer 
surface and the interface to the bulk solution to afford minimum contact with polar 
and aqueous domains (i.e. physiological media). Notably, the hydrophobic and 
hydrophilic properties, as well as other non-covalent binding properties of these 
spatial binding-sites are expected to strongly influence these guest-host relationships. 
Analysis of a typical symmetrically branched dendrimer makes it apparent 
that there are other subtle and yet important parameters that could control the interior 
space of a dendrimer and influence the guest-host interactions. These include 
components such as branching angles, branching symmetry rotational angles, and 
the length of a repeat-unit segment.13 Of equal importance are the properties of the 
core. Within a homologous PAMAM dendrimer series, the effect of changing the 
length scale of the core on dendrimer guest-host properties was studied. Specifically, 
a series of polyhydroxy-surfaced PAMAM dendrimers with core molecules 
differing in length by one carbon atom (NH2-Cn-NH2 with n = 2-6) were synthesized. 
Three aromatic carboxylic acids, differing systematically by one aromatic 
ring (benzoic acid, 1-naphthoic acid, 9-anthracene carboxylic acid), were examined 
as guest-molecule probes. Two sets of dendrimers, possessing 24 and 48 surface 
hydroxy groups, were investigated.14 The observed trends can be summarized as 
follows: (i) in general, all dendritic hosts accommodated larger amounts of the 
smaller guest-molecule (i.e. molar uptake benzoic > 1-naphthoic > 9-anthracene 
carboxylic acid). This observation was particularly significant for the more congested 
dendrimer surface having 48 surface OH-groups. (ii) Uptake maxima values 
specific to both the core size and the specific guest-probe were noted. This observation 
might be related to the combination of shape and lipophilicity manifested by the 
guest probe, (iii) A decrease in the molar uptake was measured for all probes as the 
core was enhanced beyond an ideal dimension (i.e. 5-6 carbons). It is therefore obvious 
that both core size and surface congestion dramatically affect the cargo-space 
of the dendrimer host. Furthermore, it is apparent that size and shape of the guest 
probe can significantly affect the maximum loading as a function of core size. Finally, 
it should also be noted that for the dendrimers G = 2 (24-OH) and G = 3 (48-OH), 
the guest probes had desirable release properties from the host as a function of time, 
when re-dissolved in water. Performing these same experiments using a dendrimer 
282 Svenson & Tomalia 
with more densely packed surface groups (i.e. G = 4 with 96 surface OH-groups) 
appeared to produce dendritic box behavior. Although guest molecules could be 
encapsulated within the core, the release from the host was delayed as determined 
by analysis after extensive dialysis.14 Structure-property relationships in dendritic 
encapsulation have been studied extensively, mainly using photoactive and redoxactive 
model dendrimers to gain a better understanding of the structural effects 
that cores and branches have on encapsulation.15-17 
2.3. Dendrophanes and dendroclefts 
Specific binding of guest molecules to the dendrimer core can affect the loading 
capacity by enhancing specific interactions between the core and guest (i.e. 
hydrophobic and polar interactions). Dendrimers specifically tailored to bind 
hydrophobic guests to the core have been created by Diederich and coworkers 
and coined "dendrophanes". These water-soluble dendrophanes are built around 
a cyclophane core, and can bind aromatic compounds, presumably via p -p interactions. 
Dendrophanes were shown to be excellent carriers of steroids.18'19 The same 
group synthesized dendrimers tailored to bind more polar bioactive compounds 
to the core, coined "dendroclefts".20'21 In another approach, the surface amines 
of PAMAM dendrimers were modified with tris(hydroxymethyl)aminomethane 
(TRIS) to create water-soluble dendrimers capable of binding carboxylic aromatic, 
antibacterial compounds, which could be released by lowering the pH.14 An alternative 
approach to creating dendritic hosts with highly selective guest recognition 
utilized the principle of "molecular imprinting".22 A dendrimer consisting of a porphyrin 
core and a surface containing terminal double bonds was polymerized into 
a polydendritic network. Subsequently, the base-labile ester bonds between cores 
and dendritic wedges were cleaved, releasing the porphyrin core from the dendritic 
polymer. This polymer was capable of selectively binding porphyrins with 
association constants of 1.4 x 105 M_1. Very recently, an impressive approach has 
been presented, using tandem mass spectrometry, i.e. the combination of electrospray 
ionization (ESI) and collision-induced dissociation (CID) mass spectrometers 
connected in series, to investigate the dynamic behavior of host-guest dendrimer 
complexes.23 This approach offers the potential to provide better insights into these 
3. Dendrimers in Drug Delivery 
Dendrimers have been utilized to carry a variety of small molecule pharmaceuticals 
with the purpose to enhance their solubility and therefore bioavailability, and to 
utilize the passive and active targeting properties of dendrimers, either through the 
Dendrimers as Nanoparticulate Drug Carriers 283 
"Enhanced Permeability and Retention" (EPR)24 effect or specific targeting ligands. 
Some aspects of dendrimers in drug delivery have been reviewed recently.13,25-27 
In the following, selected examples of important drug delivery aspects will be 
3.1. Cisplatin 
Encapsulation of the well-known anticancer drug cisplatin within PAMAM dendrimers 
gives complexes that exhibit slower release, higher accumulation in solid 
tumors, and lower toxicity compared with free cisplatin.28'29 Cisplatin is an antitumor 
drug that exerts its effects by forming stable DNA-cisplatin complexes 
through intrastrand cross-links, resulting in an alteration of the DNA structure that 
prevents replication and activates cell repair mechanisms. The cell detects defective 
DNA and initiates apoptosis. Cisplatin is effective in treating several cancers such 
as ovarian, head and neck, and lung cancers, as well as melanomas, lymphomas, 
osteosarcomas, bladder, cervical, bronchogenic, and oropharyngeal carcinomas. 
Unfortunately, cisplatin has many adverse side effects to the body, the most important 
being nephrotoxicity and cytotoxicity to non-cancerous tissue, because of the 
non-selective interaction between cisplatin and DNA. In addition, the therapeutic 
effect of cisplatin is limited by its poor water solubility (1 mg/mL), low lipophilicity, 
and the development of resistance to cisplatin drugs. Although numerous cisplatin 
derivatives have undergone preclinical and clinical testing, only cisplatin and its 
derivatives carboplatin and oxaliplatin have been approved for routine clinical use 
(Fig. 2).30 
Preliminary studies gave cisplatin loadings of 15-25 wt% for PAMAM dendrimers 
generation 3.5 (size ~ 3.5 nm; MW ~ 13 kDa). In comparison, the cisplatin 
loading of linear poly(amidoamines) and linear N-(2-hydroxypropyl) methacrylamide 
(HPMA; MW 25-31 kDa) was found to be 5-10 wt% and 3-8 wt%, 
respectively. HPMA-cisplatin complexes are currently in clinical trials.31 The 
cisplatin-dendrimer complex could be visualized by Atomic Force Microscopy 
(AFM; carbon nanotip) as shown in Fig. 3. 
H3N, CI 
H3NT \ 
H3N- \ l ^ V 
Fig. 2. Chemical structures of the platinum drugs cisplatin (PLATINOL), carboplatin 
(PARAPLATIN), and oxaliplatin (ELOXATIN). 
284 Svenson & Tomalia 
Fig. 3. AFM images of cisplatin-dendrimer complexes at 120 (left) and 4nm (right) 
Table 1 AUC value (/xg Pt/mLblood or /xg Pt/organ) 
over 48 hours; 5 mice/data point. 
Organ Cisplatin Cisplatin-dendrimer Complex 
Tumor 5.3 25.4 
Blood 9.4 10.7 
Liver 51.6 17.0 
Kidney 57.6 138.1 
The tumor activity of the cisplatin-dendrimer formulation was studied using 
B16F10 cells. These cells were injected into C57 mice subcutaneously (s.c.) to provide 
a solid tumor model. After approximately 12 days, when the tumors had developed 
to a mean area of 50-100 mm2, the animals were injected i.v. with a single dose of 
either cisplatin or cisplatin-dendrimer complex (1 mg/kg cisplatin for both formulations). 
At certain time points within 48 hours, animals were culled and blood and 
tissue samples were taken. Compared with cisplatin alone, the cisplatin-dendrimer 
complex was found to accumulate preferentially in the tumor site relatively quickly 
after the injection. The tumor area under the curve (AUC) for the complex was 
5 times higher than that of free cisplatin, while that in the kidney only increased 
2.4 times, and accumulation in the liver was reduced (Table I).29 
Another recent study revealed a sufficient stability of cisplatin-dendrimer complexes, 
with a 20% release of cisplatin over the first 8 hours, and an additional 60% 
release within 150 hours. In vivo animal efficacy of the platinate was demonstrated 
using B16F10 tumor cells that are subcutaneous implanted into mice. The tumor 
was allowed to grow for 7 days prior to treatment with two doses of drug on day 
7 and day 14, providing equal cisplatin (5 mg/kg) doses in both the dendrimercisplatin 
complex and free cisplatin. A tumor weight reduction of ~ 40% above that 
observed for the free drug was found in this study. 
Dendrimers as Nanoparticulate Drug Carriers 285 
3.2. Silver salts 
The encapsulation of silver salts within PAMAM dendrimers produced conjugates, 
exhibiting slow silver release rates and antimicrobial activity against various Gram 
positive bacteria.32 PAMAM dendrimers, generation four with ethylenediamine 
(EDA) core and tris(2-hydroxymethyl)amidomethane (TRIS) OH-surface and generation 
five, EDA core with carboxylate COO~ surface, were used. Silver containing 
PAMAM complexes were prepared by adding aqueous solutions of the dendrimers 
to the calculated amount of silver acetate powder. Although CHaCOOAg is hardly 
soluble in water, it quickly dissolved in the PAMAM solutions. This enhancement 
is due to the combined action of the silver carboxylate salt formation and/or to the 
complex formation with the internal dendrimer nitrogens. This procedure resulted 
in slightly yellow dendrimer-complex/salt solutions that very slowly photolyzed 
when exposed to light, into dark brown, metallic silver, containing dendrimersilver 
nanocomposite solutions. Final sample concentrations were confirmed by 
atomic absorption spectroscopy. For antimicrobial testing, the standard agar overlay 
method was used. In this test, dendrimer-silver compounds were examined 
for diffusible antimicrobial activity by placing a lO-^L sample of each solution 
onto a 6-mm filter paper disk and applying the disk to a dilute population of the 
test organisms, Staphylococcus aureus, Pseudomonas aeruginosa, and Escherichia coli. 
The silver-dendrimer complexes displayed antimicrobial activity, comparable to or 
better than those of silver nitrate solutions. Interestingly, increased antimicrobial 
activity was observed with dendrimer carboxylate salts, which was attributed to 
the very high local concentration (256 carboxylate groups around a 5.4 nm diameter 
sphere) of nanoscopic size silver composite particles that are accessible for 
microorganisms. The antimicrobial activity was smaller when internal silver complexes 
were applied instead of silver adducts to the surface, indicating that the 
accessibility of the silver is an important factor. 
3.3. Adriamycin, methotrexate, and 5-fluorouracil 
The anticancer drugs, adriamycin and methotrexate, were encapsulated into generations 
3 and 4 PAMAM dendrimers which had poly(ethylene glycol) monomethyl 
ether chains with molecular weights of 550 and 2000 Da attached to their surfaces 
via urethane bonds (Fig. 4). The encapsulation efficiency was dependent on the 
PEG chain length and the size of the dendrimer, with the highest encapsulation 
efficiencies (on average, 6.5 adriamycin molecules and 26 methotrexate molecules 
per dendrimer) found for the G = 4 PAMAM terminated with PEG2000 chains. 
The drug release from this dendrimer was sustained at low ionic strength, again 
reflecting PEG chain length and dendrimer size, but fast in isotonic solution.33 In a 
related study, it was reported that the surface coverage of PAMAM dendrimers with 
286 Svenson & Tomalia 
Fig. 4. Above: Structures of anticancer drugs adriamycin (left) and methotrexate (right). 
Below: Schematic presentations of the encapsulation of methotrexate (left) and 5-fluorouracil 
(right) into PAMAM dendrimers. 
PEG2000 chains had little influence on the encapsulation efficiency of methotrexate, 
but affected the release rate.34 
A similar construct between PEG chains and PAMAM was utilized to deliver the 
anticancer drug 5-fluorouracil. Encapsulation of 5-fluorouracil into G  4 PAMAM 
dendrimers with carboxymethyl PEG5000 surface chains revealed reasonable drug 
loading, a reduced release rate, and reduced hemolytic toxicity compared to the 
non-PEGylated dendrimer (Fig. 4).35 
3.4. Etoposide, mefenamic acid, diclofenac, and venlafaxine 
The combination between dendrimers and hydrophilic and/or hydrophobic polymer 
chains has recently been extended to solubilize the hydrophobic anticancer 
drug etoposide. A star polymer composed of amphiphilic block copolymer arms 
has been synthesized and characterized. The core of the star polymer was a 
generation two PAMAM-OH dendrimer, the inner block of the arm a lipophilic 
poly(e-caprolactone) (PCL) and the outer block of the arm a hydrophilic PEG500o- 
The star-PCL polymer was synthesized first by ring-opening polymerization of 
e-caprolactone with the PAMAM-OH dendrimer as initiator. The PEG polymer 
was then attached to the PCL terminus by an ester-forming reaction. Characterization 
with SEC, 1-H NMR, FTIR, TGA, and DSC confirmed the star structure of the 
polymers. A loading capacity of up to 22% (w/w) was achieved with etoposide. 
Dendrimers as Nanoparticulate Drug Carriers 287 
A cytotoxicity assay demonstrated that the star-PCL-PEG copolymer was nontoxic 
in cell culture.36 
Citric acid-poly(ethylene glycol)-citric acid (CPEGC) triblock dendrimers generations 
1-3 were applied to encapsulate small molecule drugs such as mefenamic 
acid and diclofenac. The formulations were stored at room temperature for up to 
ten months and remained stable with no reported release of the drugs.37 
The attachment of the novel third-generation antidepressant venlafaxine onto 
anionic PAMAM dendrimers (G = 2.5) via a hydrolyzable ester bond and the incorporation 
of this drug-dendrimer complex into a semi-interpenetrating network of 
an acrylamide hydrogel has been studied as a novel drug delivery formulation to 
avoid the currently necessary multiple daily administration of the antidepressant. 
The effect of PEG concentration and molecular weight was studied to find optimal 
release conditions.38 
3.5. Ibuprofen, indomethacin, nifedipine, naproxen, paclitaxel, 
and methylprednisolone 
The anti-inflammatory drug ibuprofen was used as a model compound to study 
its complexation and encapsulation into generations 3 and 4 PAMAM dendrimers 
and a hyperbranched polyester, having approximately 128 surface OH-groups. It 
was found that up to 78 ibuprofen molecules were complexed by the PAMAM dendrimers 
through electrostatic interactions between the dendrimer amines and the 
carboxyl group of the drug. In contrast, up to 24 drug molecules were encapsulated 
into the hyperbranched polyol.39 The drug was successfully transported into A549 
human lung epithelial carcinoma cells by the dendrimers. The PAMAM dendrimers 
with either amino or hydroxy surfaces entered the cells faster (in approximately 1 hr) 
than the hyperbranched polyol (approximately 2 hrs). However, both entries were 
faster than the pure drug. The anti-inflammatory effect of ibuprofen-dendrimer 
complexes was demonstrated by more rapid suppression of COX-2 mRNA levels 
than that achieved by the pure drug.40 
The non-steroidal anti-inflammatory drug (NSAID) indomethacin is practically 
insoluble in water and only sparingly soluble in alcohol. Encapsulation of 
indomethacin into generation 4 PAMAM dendrimers with amino, hydroxy, and 
carboxylate surfaces remarkably enhanced the drug solubility in water, and therefore, 
its bioavailability (Fig. 5).41 The encapsulation efficiency of indomethacin into 
PAMAM dendrimers is dependent on the dendrimer size (G6 > G5 > G4 > G3) 
and the surface functionalization, (NH2 > PEG = PYR > AE) (Fig. 6).42 
The effect of PAMAM dendrimer generation size and surface functional group 
on the aqueous solubility, and therefore, bioavailability of the calcium channel 
blocking agent nifedipine has been studied using PAMAM dendrimers with EDA 
288 Svenson & Tomalia 
E 800^ 
g 400- 
0.1 0.2 
Dendrimerconc. (%v/w) 
Fig. 5. Molecular structure of indomethacin and its solubility profiles in the presence of differing 
concentrations of G4-NH2/ () G4-OH (), and G4.5-COOH (A) PAM AM dcndrimers 
at pH 7 ( = 3, R.S.D. < 5%). 
Encapsulation efficiency of EDA core PAMAM dendrimer 
80 - 
g 40 
3 20- 
,rf sJl 
n .r-dl ^ 
Surface Functionality 
sue TRIS 
Fig. 6. Encapsulation efficiency into PAMAM dendrimers generations 3-6 with amino 
(NH2), poly(ethylene glycol) (PEG), carbomethoxypyrrolidinone (PYR), amidoethanol 
(AE), sodium carboxylate (COONa), succinamic acid (SUC), and tris(hydroxymethyl)- 
aminomethane (TRIS) surface groups. 
core and amino surface (G = 0,1,2,3) or ester surface (G  0.5,1.5,2.5) at pH 4, 
7 and 10. The solubility enhancement of nifedipine was higher in the presence of 
ester-terminated dendrimers than their amino-terminated analogues, possessing 
the same number of surface groups. The nifedipine solubility expectedly increased 
with the size of the dendrimers. For pH 7, the sequence G2.5 > G3 > G1.5 > G2 > 
G0.5 > Gl > GO was reported.43 
In another approach, the non-steroidal anti-inflammatory drug naproxen was 
covalently attached to unsymmetrical poly(arylester) dendrimers to prepare a complex 
with enhanced water solubility of the drug and access for hydrolytic cleavage 
Dendrimers as Nanoparticulate Drug Carriers 289 
1E-3 0.01 0.1 1 0 2k 4k 6k 8k 
Concentration (M) Molecular weight 
Fig. 7. Aqueous paclitaxel solubility as a function of the polyglycerol dendrimer concentration 
(mean  SD, n = 3); G5 (circle), G4 (triangle), G3 (square), and PEG400 (diamond) 
(left). Molecular weight dependency of dendrimers (closed circle) and PEG (open circle) on 
the aqueous paclitaxel solubility. The concentration of dendrimers and PEG was 10wt%. 
(Reproduced with permission from Ref. 45. Copyright 2004 American Chemical Society.) 
of the bond between drug and carrier. Detailed results on the biological evaluation 
of these complexes have not been reported.44 
The anticancer drug paclitaxel, which is being used to treat metastatic breast 
and ovarian cancers and Kaposi's sarcoma, has poor water solubility. To enhance 
its bioavailability, paclitaxel has been encapsulated into polyglycerol dendrimers, 
resulting in a 10,000-fold improved water solubility compared with the pure drug, 
which is much higher than that found for PEG400, a commonly used linear chain 
cosolvent or hydrotropic agent (Fig. 7). The drug release rate was a function of the 
dendrimer generation.45 
Generation 4 PAMAM dendrimers with hydroxy surface have been utilized 
to improve the bioavailability of the corticosteroid methylprednisolone, which 
decreases inflammation by stabilizing leukocyte lysosomal membrane. By connecting 
the drug to the dendrimer using glutaric acid as the spacer, a payload of 32 
wt% was achieved. The drug-dendrimer complex was taken up by A549 human 
lung epithelial carcinoma cells and mostly localized in the cytosol. The complex 
showed a pharmacological activity comparable to the free drug as measured by the 
inhibition of the prostaglandin secretion.46 
3.6. Doxorubicin and camptothecin  self-immolative dendritic 
An exciting new approach to dendritic drug delivery involves the utilization of a 
drug as a part of the dendritic molecule. Self-immolative dendrimers have recently 
been developed and introduced as a potential platform for a multi-prodrug. These 
unique structural dendrimers can release all of their outer branch units through 
jj 0.01- 
I 1E-3, 
290 Svenson & Tomalia 
Fig. 8. Mechanism of dimeric prodrug activation by a single enzymatic cleavage. (Reproduced 
with permission from Ref. 47. Copyright 2004 American Chemical Society.) 
a self-immolative chain fragmentation, initiated by a single cleavage at the dendrimer's 
core. Incorporation of drug molecules as these outer branch units and an 
enzyme substrate as the trigger can generate a multi-prodrug unit that will be activated 
with a single enzymatic cleavage (Fig. 8). The first generation of dendritic 
prodrugs with doxorubicin and camptothecin as branch units and retro-Michael 
focal trigger, which can be cleaved by the catalytic antibody 38C2, has been reported. 
Bioactivation of the dendritic prodrugs was evaluated in cell-growth inhibition 
assay with the Molt-3 leukemia cell line in the presence and absence of antibody 
38C2. A remarkable increase in toxicity was observed. Dependent on the linker 
molecule, different numbers of drug molecules can be released in one single activation 
In a more "classical" approach to deliver doxorubicin, two polyester-based 
dendrimers (generation 4 with trisphenolic core) were synthesized, one carrying a 
hydroxy surface, the other a tri(ethylene glycol) monomethyl ether surface. These 
dendrimers were compared with a 3-arm poly(ethylene oxide) star polymer, carrying 
G = 2 dendritic polyester units at the surface. The star polymer gave the most 
promising results regarding cytotoxicity and systemic circulatory half-life (72hrs). 
Therefore, the anticancer drug doxorubicin was covalently bound to this carrier via 
an acid-labile hydrazone linkage. The cytotoxicity of doxorubicin was significantly 
reduced (80-98%) and the drug was successfully taken up by several cancer cell 
Dendrimers as Nanoparticulate Drug Carriers 291 
3.7. Photodynamic therapy (PDT) and boron neutron capture 
therapy (BNCT) 
Dendrimers have been used to optimize the antitumor effect in photodynamic 
therapy (PDT) and boron neutron capture therapy (BNCT). One of the newest 
developments in the dendrimer field is their application to photodynamic therapy 
(PDT). This cancer treatment involves the administration of a light-activated 
photosensitizing moiety that selectively concentrates in diseased tissue. Subsequent 
activation of the photosensitizer leads to the generation of reactive oxygen, primarily 
singlet oxygen, that damages intracellular components such as lipids and amino 
acid residues through oxidation, ultimately leading to cell death by apoptosis. 
Disadvantages of currently used photosensitizers include skin phototoxicity, poor 
selectivity for tumor tissue, poor water solubility, and difficulties in the treatment 
of solid tumors because of the impermeability of the skin and tissues to the visible 
light required to excite the chromophores. 
In one set of studies, dendrimers have been constructed around a light harvesting 
core (i.e. a porphyrin).50 To reduce the toxicity under non-irradiative conditions 
(dark toxicity) and to prevent aggregation, and consequently, self-quenching of the 
porphyrin cores, these dendrimers have been further encapsulated into micelles. 
For example, poly(ethylene glycol)-b-poly(aspartic acid) and PEG-b-poly(L-lysine) 
micelles have been studied in this regard. These micelles are stable under physiological 
conditions pH 6.2 to 7.4. However, they disintegrate in the acidic intracellular 
endosomal compartment (pH ~ 5.0).51/52 Alternatively, the photosensitizer 
5-aminolevulinic acid has been attached to the surface of dendrimers and studied 
as an agent for PDT of tumorigenic keratinocytes.53 Photosensitive dyes have been 
incorporated into dendrimers and utilized in PDT devices. For example, uptake, 
toxicity, and the mechanism of photosensitization of the dye pheophorbide a (pheo) 
was compared with its complex with diaminobutane poly(propylene imine) (DAB) 
dendrimers in human leukemia cells in vitro.5i 
The second therapy, boron neutron capture therapy, is a cancer treatment based 
on a nuclear capture reaction. When 10B is irradiated with low energy or thermal 
neutrons, highly energetic a-particles and 7Li ions are produced, that are toxic to 
tumor cells. To achieve the desired effects, it is necessary to deliver 10B to tumor cells 
at a concentration of at least 109 atoms per cell. High levels of boron accumulation 
in tumor tissue can be achieved by using boronated antibodies that are targeted 
towards tumor antigens. However, this approach can impair the solubility and 
targeting efficiency of the antibodies. 
One study, involving intratumoral injection of a conjugation between a generation 
5 PAMAM dendrimer carrying 1100 boron atoms and cetuximab, a monoclonal 
antibody specific for the EGF receptor, showed that the conjugate was present 
292 Svenson & Tomalia 
Fig. 9. Schematic presentation of an EDA core G = 3 PAMAM dendrimer (1), the boron 
carrier Na(CH3)3NB10H8NCO (2), and the targeting ligand folic acid (3). (Reproduced with 
permission from Ref. 56. Copyright 2003 American Chemical Society.) 
at an almost 10-fold higher concentration in brain tumors than in normal brain 
tissue.55 To reduce the liver uptake observed for boronated PAMAM dendrimer 
conjugates, PEG chains were attached onto the dendrimer surface, in addition to 
the borane clusters, to provide steric shielding. As compared with a dendrimer without 
PEG chains, the amount of liver uptake was found to be less for PEG-conjugated 
dendrimers with an average of 1.0-1.5 chains of PEG2000/ but higher for dendrimers 
with 11 chains of PEG550. Folic acid moieties were also conjugated to the ends of 
the PEG chains to enhance the uptake of the dendrimers by tumors overexpressing 
folate receptors. Although this strategy was successful in enhancing localization 
of the molecules to tumors in mice bearing 24JK-FBP tumors expressing the folate 
receptor, it also led to an increase in the uptake of the dendrimers by the liver and 
4. Nano-Scaffolds for Targeting Ligands 
The surface of dendrimers provides an excellent platform for the attachment of cellspecific 
ligands, solubility modifiers, stealth molecules, reducing the interaction 
with macromolecules from the body defense system, and imaging tags. The ability 
to attach any or all of these molecules in a well-defined and controllable manner 
onto a robust dendritic surface, clearly differentiates dendrimers from other carriers 
such as micelles, liposomes, emulsion droplets, and engineered particles. 
4.1. Folic acid 
One example of cell-specific dendritic carriers is a dendrimer modified with folic 
acid. The membrane-associated high affinity folate receptor (hFR) is a folate binding 
protein that is overexpressed on the surface of a variety of cancer cells, and 
Dendrimers as Nanoparticulate Drug Carriers 293 
therefore, folate-modified dendrimers would be expected to internalize into these 
cells preferentially over normal cells via receptor-mediated endocytosis. Folatedendrimer 
conjugates have been shown to be well-suited for targeted, cancerspecific 
drug delivery of cytotoxic substances.56-59 
In a very recent study, branched poly(L-glutamic acid) chains were centered 
around PAMAM dendrimers generations 2 and 3 and poly(ethylene imine) (PEI) 
cores to create new biodegradable polymers with improved biodistribution and targeting 
ability. These constructs were surface-terminated with poly(ethylene glycol) 
chains to enhance their biocompatibility, and folic acid ligands to introduce cellspecific 
targeting. Cell binding studies have been performed using the epidermal 
carcinoma cell line, KB.60 
4.2. Carbohydrates 
In addition to folates, carbohydrates constitute another important class of biological 
recognition molecules, displaying a wide variety of spatial structures due to 
their branching possibility and anomericity. To achieve sufficiently high binding 
affinities between simple mono- and oligosaccharide ligands and cell membrane 
receptors, these ligands have to be presented to the receptors in a multivalent or 
cluster fashion.61,62 The highly functionalized surface of dendrimers provides an 
excellent platform for such presentations. The design, synthesis, and biomedical 
use of glycodendrimers, as well as their application in diagnostic and for vaccinations, 
have been thoroughly reviewed recently.63-69 For example, the Thomsen- 
Friedenreich carbohydrate antigen (T-antigen), /J-Gal-(l-3)-a-GalNAc, which has 
been well documented as an important antigen for the detection and immunotherapy 
of carcinomas, especially relevant to breast cancer, has been attached to the 
surface of PAMAM and other dendrimers.70-72 An enhanced binding affinity was 
observed for all glycodendrimers. These constructs could have potential in blocking 
the metastatic sites of invasive tumor cells. A series of dendritic ,6-cyclodextrin 
derivatives, bearing multivalent mannosyl ligands, has been prepared and their 
binding efficiency towards the plant lectin concanavalin A (Con A) and a mammalian 
mannose-specific cell surface receptor from macrophages has been studied. 
The effects of glycodendritic architecture on binding efficiency, molecular inclusion, 
lectin-binding properties, and the consequence of complex formation using 
the anticancer drug docetaxel on biological recognition were investigated.73 Di- to 
tetravalent dendritic galabiosides, carrying (Galal-4Gal) moieties on their surfaces, 
were studied as inhibitors of pathogens based on bacterial species such as E. colt and 
Streptococcus suis. Attachment of dendritic galabiosides onto cell surfaces would be 
expected to inhibit the attachment of bacteria using the same sugar ligand-receptor 
interactions. The study revealed a clear enhancement of the binding affinity between 
294 Svenson & Tomalia 
glycodendrons and cell surfaces, with an increasing number of sugar moieties.74 In 
a similar approach, glycodendrons carrying two to four /i-D-galactose moieties on 
their surface, while the dendron core was connected to a protein-degrading enzyme, 
were synthesized. These glycodendriproteins are expected to attach to the surface of 
bacteria, allowing the enzyme to degrade the bacterial adhesin, hence rendering the 
bacteria incapable of attaching to the cell surfaces.75 Anionic PAMAM dendrimers 
(G = 3.5) were conjugated to D(+)-glucosamine and D(+)-glucoseamine 6-sulfate. 
These water-soluble conjugates not only revealed immuno-modulatory and antiangiogenic 
properties, but synergistically prevented scar tissue formation after glaucoma 
filtration surgery. In a validated and clinically relevant rabbit study, the longterm 
success rate was increased from 30 to 80% using these dendrimer-conjugates.76 
4.3. Antibodies and biotin-avidin binding 
Generation 5 PAMAM dendrimerendrimers with amino surface were conjugated 
to fluorescein isothiocyanate as a means to analyze cell binding and internalization. 
Two different antibodies, 60bca and J591, which bind to CD14 and prostate-specific 
membrane antigen (PSMA) respectively, were used as model targeting molecules. 
The binding of the antibody-conjugated dendrimers to antigen-expressing cells 
was evaluated by flow cytometry and confocal microscopy. The conjugates specifically 
bound to the antigen-expressing cells in a time- and dose-dependent fashion, 
with affinity similar to that of the free antibody (Fig. 10). Confocal microscopic 
analysis suggested at least some cellular internalization of the dendrimer conjugate. 
Dendrimer-antibody conjugates are, therefore, a suitable platform for targeted 
molecule delivery into antigen-expressing cells.77 
Monolayers formed by generation 4 PAMAM dendrimers on a gold surface 
were functionalized with biotin and produced a biomolecular interface that was 
Fig. 10. Confocal microscopic analysis of HL60 cells, which were incubated (1 h at 4C) 
with 12.5nM G5 PAMAM carrying fluorescence dye and 60bca antibody on the surface. The 
cells were rinsed and confocal images were taken. The left and right panels represent the FITC 
fluorescence and light images taken in the same cell. The arrow indicates the binding of the 
conjugate on the cell surface at 4C. (Reproduced with permission from Ref. 77. Copyright 
2004 American Chemical Society.) 
Dendrimers as Nanoparticulate Drug Carriers 295 
capable of binding high levels of avidin. Avidin binding as high as 88% coverage 
of the surface was observed despite conditions that should cause serious steric 
hindrance. These dendritic monolayers were utilized as a model to study proteinligand 
4.4. Penicillins 
The surfaces of PAMAM dendrimers, generations 0 to 3, were decorated with benzylpenicillin 
in an attempt to develop a new in vitro test to quantify IgE antibodies 
to specific ^-lactam conjugates, with the goal of improving the existing methods for 
diagnosing allergy to this type of antibiotic. The monodispersity of dendrimers is 
advantageous over conventional peptide carrier conjugates such as human serum 
albumin (non-precise density of haptens in their structure) and poly-L-lysine (mixture 
of heterogeneous molecular weight peptides). Preliminary radioallergosorbent 
tests (RAST), using sera from patients allergic to penicillin, have confirmed the usefulness 
of penicilloylated dendrimers.79 
Penicillin V was used as a model drug containing a carboxylic group and 
attached to the surface of PAMAM dendrimers generations 2.5 and 3, both containing 
32 surface functionalities. The drug was complexed to the dendrimers via 
amide or ester bonds. It was found in tests using a single-strain bacterium, Staphylococcus 
aureus, that the bioavailability of the penicillin was unaltered after the drug 
was released from the complex through ester bond hydrolysis.80 
5. Dendrimers as Nano-Drugs 
Dendrimers have been studied as antitumor, antiviral and antibacterial drugs.25 
The most prominent and advanced example is the use of poly(lysine) dendrimers, 
modified with sulfonated naphthyl groups, as antiviral drugs against the herpes 
simplex virus.81 Such a conjugate based on dendritic poly(lysine) scaffolding is 
VivaGel, a topical agent currently under development by Starpharma Ltd., Melbourne, 
Australia, that can potentially prevent/reduce transmission of HIV and 
other sexually transmitted diseases (STDs). VivaGel (SPL 7013) is being offered 
as a water-based gel, with the purpose to prevent HIV from binding to cells in the 
body. The gel differs from physical barriers to STDs such as condoms, by exhibiting 
inhibitory activity against HIV and other STDs. In July 2003, following submission 
of an Investigational New Drug (IND) application, Starpharma gained clearance 
under U.S. FDA regulations to proceed with a Phase I clinical study to assess 
the safety of VivaGel in healthy human subjects. This phase 1 study, representing 
for the first time a dendrimer pharmaceutical tested in humans, compared 36 
women who received either various intra-vaginal doses of VivaGel or a placebo 
gel daily for one week. The trial was double blinded so that the volunteers, principal 
296 Svenson & Tomalia 
investigator and Starpharma did not know who was receiving placebo or VivaGel. 
Study participants were assessed for possible irritant effects of the gel. Additionally, 
the women were assessed for any possible effect upon vaginal microflora (natural 
micro-organisms in the vagina) or absorption into the blood of the active ingredient 
of VivaGel. A thorough review of the complete data revealed no evidence of irritation 
or inflammation. Preclinical development studies had demonstrated that 
VivaGel was 100% effective at preventing infection of primates exposed to a 
humanized strain of simian immunodeficiency virus (SHIV).82 In earlier studies, 
it was found that PAMAM dendrimers covalently modified with naphthyl sulfonate 
residues on the surface, also exhibited antiviral activity against HIV. This 
dendrimer-based nano-drug inhibited early stage virus/cell adsorption and later 
stage viral replication, by interfering with reverse transcriptase and/or integrase 
enzyme activities.83,84 
The general mode of action of antibacterial dendrimers is to adhere to and 
damage the anionic bacterial membrane, causing bacterial lysis.25,85 PPI dendrimers 
with tertiary alkyl ammonium groups attached to the surface have been shown to 
be potent antibacterial biocides against Gram positive and Gram negative bacteria. 
The nature of the counterion is important, as tetraalkylammonium bromides 
were found to be more potent antibacterials over the corresponding chlorides.86 
Poly(lysine) dendrimers with mannosyl surface groups are effective inhibitors of 
the adhesion of E. coli to horse blood cells in a haemagglutination assay, making 
these structures promising antibacterial agents.87 Chitosan-dendrimer hybrids 
have been found to be useful as antibacterial agents, carriers in drug delivery systems, 
and in other biomedical applications. Their behavior have been reviewed 
very recently88 Triazine-based antibiotics were loaded into dendrimer beads at 
high yields. The release of the antibiotic compounds from a single bead was sufficient 
to give a clear inhibition effect.89 In many cases, dendritic constructs were 
more potent than analogous systems based on hyperbranched polymers. 
The anti-prion activity of cationic phosphorus-containing dendrimers with tertiary 
amine surface groups has been evaluated. These molecules had a strong anti 
prion activity at non-toxic doses. They have been found to decrease the amount of 
pre-existing PrPSc from several prion starins, including the BSE strain. In addition, 
these dendrimers were able to reduce PrPSc accumulation in the spleen by more 
than 80%.90 
6. Routes of Application 
Most commonly, dendrimers are applied as parenteral injections, either directly 
into the tumor tissue or intravenous for systemic delivery. However, recent oral 
drug delivery studies using the human colon adenocarcinoma cell line, Caco-2, 
Dendrimers as Nanoparticulate Drug Carriers 297 
have indicated that low generation PAMAM dendrimers cross cell membranes, 
presumably through a combination of two processes, i.e. paracellular transport 
and adsorptive endocytosis.91 The P-glycoprotein (P-gp) efflux transporter does 
not effect dendrimers, and therefore, drug-dendrimer complexes are able to bypass 
the efflux transporter.92 
Furthermore, recent work has shown that PAMAM dendrimers enhanced the 
bioavailability of indomethacin in transdermal delivery applications.93 Similarly, 
the drug tamsulosin was used as a model to study transdermal delivery utilizing 
PAMAM dendrimers. The dendrimers were found to be weak penetration 
enhancers.94 However, no dendrimer-driven effect was observed for the drugs ketoprofen 
and clonidine. As an explanation, dendrimer-triggered drug crystallization 
within the transdermal delivery matrix was discussed, allowing the formation of 
drug polymorphs that can or cannot facilitate transdermal delivery95 
Several PAMAM dendrimers (generations 1.5, 2-3.5 and 4) with amine, carboxylate 
and hydroxyl surface groups were studied for controlled ocular drug 
delivery. The duration of residence time was evaluated after solubilization of these 
dendrimers in buffered phosphate solutions containing 2 parts per thousand (w/v) 
of fluorescein. The New Zealand albino rabbit was used as an in vivo model for qualitative 
and quantitative assessment of ocular tolerance and retention time, after a single 
application of 25 /xL of dendrimer solution to the eye. The same model was also 
used to determine the prolonged miotic or mydriatic activities of dendrimer solutions, 
some containing pilocarpine nitrate and some tropicamide, respectively. Residence 
time was longer for the solutions containing dendrimers with carboxylic and 
hydroxyl surface groups. No prolongation of remanence time was observed when 
dendrimer concentration (0.25-2%) increased. The remanence time of PAMAM dendrimer 
solutions on the cornea showed size and molecular weight dependency. This 
study allowed novel macromolecular carriers to be designed with prolonged drug 
residence time for the ophthalmic route.96 
7. Biocompatibility of Dendrimers 
Dendrimers have to exhibit low toxicity and be non-immunogenic in order to be 
widely used in biomedical applications. To date, the cytotoxicity of dendrimers 
has been primarily studied in vitro, however, a few in vivo studies have been 
published.25 As observed for other cationic macromolecules, including liposomes 
and micelles, dendrimers with positively charged surface groups are prone to destabilize 
cell membranes and cause cell lysis. For example, in vitro cytotoxicity IC50 
measurements (i.e. the concentration where 50% of cell lysis is observed) for aminoterminated 
PAMAM dendrimers revealed significant cytotoxicity on human intestinal 
adenocarcinoma Caco-2 cells.97,98 Furthermore, the cytotoxicity was found to be 
298 Svenson & Tomalia 
generation-dependent, with higher generation dendrimers being the most toxic. ' 
A similar generation dependence of amino-terminated PAMAM dendrimers was 
observed for the haemolytic effect, studied on a solution of rat blood cells.100 However, 
some recent studies have shown that amino-terminated PAMAM dendrimers 
exhibit lower toxicity than more flexible amino-functionalized linear polymers perhaps 
due to lower adherence of the rigid globular dendrimers to cellular surfaces. 
The degree of substitution, as well as the type of amine functionality, is important, 
with primary amines being more toxic than secondary or tertiary amines." 
Amino-terminated PPI and PAMAM dendrimers behave similarly with regard to 
cytotoxicity and haemolytic effects, including the generation-dependent increase 
of both.100'101 
Comparative toxicity studies on anionic (carboxylate-terminated) and cationic 
(amino-terminated) PAMAM dendrimers using Caco-2 cells have shown a significantly 
lower cytotoxicity of the anionic compounds.97 In fact, lower generation 
PAMAM dendrimers possessing carboxylate surface groups show neither haematotoxicity 
nor cytotoxicity at concentrations up to 2 mg/ml.100 The biocompatability 
of dendrimers is not solely determined by the surface groups. Dendrimers containing 
an aromatic polyether core and anionic carboxylate surface groups have shown 
to be haemolytic on a solution of rat blood cells after 24hrs. It is suggested that 
the aromatic interior of the dendrimer may cause haemolysis through hydrophobic 
membrane contact.100 
One way to reduce the cytotoxicity of cationic dendrimers may reside in partial 
surface derivatization with chemically inert functionalities such as PEG or fatty 
acids. The cytotoxicity towards Caco-2 cells can be reduced significantly (from 
IC50 ~ 0.13 mM to >lmM) after such a modification. This observation can be 
explained by the reduced overall positive charge of these surface-modified dendrimers. 
Apartial derivatization with as few as six lipid chains or four PEG chains on 
a G4-PAMAM, respectively, was sufficient to lower the cytotoxicity substantially.98 
In studies conducted at Dendritic Nano Technologies, Inc. using Caco-2 and two 
other cell lines, it was found that besides (partial) PEGylation of the surface, surface 
modification with pyrrolidone, another biocompatible compound, can significantly 
reduce cytotoxicity to levels far better than those of currently available products.102 
In some cases, the cytotoxicity of PAMAM dendrimers could be reduced by additives 
such as fetal calf serum.103 
Only a few systematic studies on the in vivo toxicity of dendrimers have been 
reported so far. Upon injection into mice, doses of 10 mg/kg of PAMAM dendrimers 
(up to G = 5), displaying either unmodified or modified amino-terminated surfaces, 
did not appear to be toxic.81-104 Hydroxy- or methoxy-terminated dendrimers 
based on a polyester dendrimer scaffold have been shown to be of low toxicity 
both in vitro and in vivo. At very high concentrations (40 mg/ml), these polyester 
Dendrimers as Nanoparticulate Drug Carriers 299 
dendrimers induced some inhibition of cell growth in vitro, but no increase in cell 
death was observed. Upon injection into mice, no acute or long-term toxicity problems 
were observed. The non-toxic properties make these new dendritic motifs very 
promising candidates for drug delivery devices.49 
Initial immunogenicity studies performed on unmodified amino-terminated 
PAMAM dendrimers showed no or weak immunogenicity of the G3-G7 dendrimers. 
However, later studies indicated some immunogenicity of these dendrimers, 
which could be reduced by surface-modification utilizing PEG chains.105 
8. Conclusions 
The high level of control over the architecture of dendrimers, their size, shape, 
branching length and density, and their surface functionality, makes these compounds 
ideal carriers in drug delivery applications. The bioactive agents may either 
be encapsulated into the interior of the dendrimers or they may be chemically 
attached or physically adsorbed onto the dendrimer surface, with the option to 
tailor the properties of the carrier to the specific needs of the active material and its 
therapeutic applications. Furthermore, the high density of surface groups allows 
attachment of targeting groups as well as groups that modify the solution behavior 
or toxicity of dendrimers. Surface-modified dendrimers themselves may act 
as nano-drugs against tumors, bacteria and viruses. This review of drug delivery 
applications of dendrimers clearly illustrates the potential of this new "fourth architectural 
class of polymers"106 and substantiates the high optimism for the future of 
dendrimers in this important field. 
The authors wish to thank all contributors to this fascinating field of research, as 
well as the funding agents that have supported this work over the years. In particular, 
DNT would like to acknowledge current funding by the US Army Research 
Laboratory (ARL) (Contract # W911NF-04-2-0030). 
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306 Svenson & Tomalia 
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Drug Nanocrystals/Nanosuspensions for 
the Delivery of Poorly Soluble Drugs 
Rainer H. Muller and Jens-Uwe A. H. Junghanns 
1. Introduction 
Since the last ten years, the number of poorly soluble drugs is steadily increasing. 
According to estimates, about 40% of the drugs in the pipelines have solubility 
problems.1 The increased use of high throughput screening methods leads to the 
discovery of more drugs being poorly water soluble. In the literature, figures are 
quoted that about 60 percent of the drugs coming directly from synthesis are nowadays 
poorly soluble.2 Poor solubility is not only a problem for the formulation 
development and clinical testing, it is also an obstacle at the very beginning when 
screening new compounds for pharmacological activity. From this, there is a definite 
need for smart technological formulation approaches to make such poorly 
soluble drugs bioavailable. Making such drugs bioavailable means that they show 
sufficiently high absorption after oral administration, or they can alternatively be 
injected intravenously. 
There is quite a number of formulation approaches for poorly soluble drugs 
which can be specified as "specific approaches". These approaches are suitable 
for molecules having special properties with regard to their chemistry (e.g. solubility 
in certain organic media) or to the molecular size or conformation (e.g. 
molecules to be incorporated into the cyclodextrin ring structure). Of course it 
would be much smarter to have a "universal formulation approach" applicable 
to any molecule. Such a universal formulation approach to increase the oral 
308 Muller & Junghanns 
bioavailability is micronization, meaning the transfer of drug powders into the size 
range between typically 1-10 /j-va. However, nowadays many drugs are so poorly 
soluble that micronization is not sufficient. The increase in surface area, and thus 
consequently in dissolution velocity, is not sufficient to overcome the bioavailability 
problems of very poorly soluble drugs of the biopharmaceutical specification 
class II. A consequent next step was to move from micronization to nanonization. 
Since the beginning of the 90s, the company Nanosystems propagated the use of 
nanocrystals (instead of microcrystals) for oral bioavailability enhancement, and 
also to use nanocrystals suspended in water (nanosuspensions) for intravenous or 
pulmonary drug delivery. 
The solution was simple; in general, simple solutions possess the smartness that 
they can be realized easier than complex systems and introduction to the market is 
faster. Nevertheless, it took about ten years before the first nanocrystals in a tablet 
appeared on the market, the product Rapamune by the company Wyeth in 2000. 
Compared with liposomes developed in 19683 with the first products on the market 
around 1990 (e.g. Alveofact, a lung surfactant), this was still relatively fast. What 
were the reasons that it took about one decade for nanocrystals to enter the market? 
From our point of view, pharmaceutical companies prefer to use formulation 
technology already established with know how available in the company. In addition, 
if formulation technologies are established, a company also has the possibility 
for production of the final product. Therefore, all the traditional formulation 
approaches were exploited to solve a formulation problem. In addition, formulation 
approaches were preferred, being even simpler than nanocrystals. For example, 
production of drug-containing microemulsions administered in a capsule is, 
in many cases, even simpler. Another reason for the reluctance of pharmaceutical 
companies at the beginning was the lack of large scale production methods. These 
were not available at the very beginning of the development of the nanocrystal technology. 
Meanwhile, this has changed and the major pharmaceutical companies try 
to secure or have already secured their access to nanocrystal technology. Access 
to nanocrystal technology is possible either by licencing in or alternatively by the 
attempt to develop one's own production technologies for the nanocrystals, which 
do not depend on already existing intelectual property (IP). This chapter discusses 
the physicochemical properties of nanocrystals which make them interesting for 
drug delivery, reviews and discusses briefly the various production methods available 
and highlights the opportunities for improved drug delivery using different 
application routes. 
2. Definitions 
Drug nanocrystals are crystals with a size in the nanometer range, meaning that 
they are nanoparticles with a crystalline character. There are discussions about the 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 309 
definition of a nanoparticle, referring to the size of a particle to be classified as a 
nanoparticle. Depending on the discipline, e.g. in colloid chemistry, particles are 
only considered as nanoparticles when they are in sizes below 100 nm or even below 
20 nm. Based on the size unit, in the pharmaceutical area, nanoparticles should be 
defined as having a size between a few nanometers and 1000 nm (1 /im); thus, 
microparticles possess consequently a size 1-1000 micrometer. 
A further characteristic is that drug nanocrystals are composed of 100% drug; 
there is no carrier material as in polymeric nanoparticles. Dispersion of drug 
nanocrystals in liquid media leads to "nanosuspensions", in contrast to "microsuspensions" 
or "macrosuspensions". In general, the dispersed particles need to 
be stabilized, e.g. by surfactants or polymeric stabilizers. Dispersion media can be 
water, aqueous solutions or non-aqueous media [e.g. liquid polyethylene glycol 
(PEG), oils]. Depending on the production technology, processing of drug microcrystals 
to drug nanoparticles can lead to either a crystalline or to an amorphous 
product, especially when applying precipitation. In the strict sense, such an amorphous 
drug nanoparticle should not be called nanocrystal. However, one often 
refers to "nanocrystals in the amorphous state". 
3. Physicochemical Properties of Drug Nanocrystals 
3.1. Change of dissolution velocity 
The reason for micronization is to increase the surface area, thus consequently 
according to the Noyes-Whitney equation, increasing the dissolution velocity. 
Therefore, micronization can be succesfuUy employed if the dissolution velocity 
is the rate-limiting step for oral absorption (drugs of BSC II). Of course, by moving 
one dimension further to smaller particles, the surface area is further enlarged 
and consequently, the dissolution velocity is further enhanced. In most cases, a low 
dissolution velocity is correlated with a low saturation solubility. 
3.2. Saturation solubility 
The general textbook statement is that the saturation solubility cs is a constant 
depending on the compound, the dissolution medium and the temperature. This 
is valid for powders of daily life with a size in the micrometer range or above. 
However, below a critical size of 1-2 /zm, the saturation solubility is also a function 
of the particle size. It increases with decreasing particle size below 1000 nm. 
Therefore, drug nanocrystals possess an increased saturation solubility. This has 
two advantages: 
1. According to Noyes-Whitney, the dissolution velocity is further enhanced 
because dc/dt is proportional to the concentration gradient (cs  cx)/h (cx  
bulk concentration, h  diffusional distance). 
310 Muller & Junghanns 
2. Due to the increased saturation solubility, the concentration gradient between 
gut lumen and blood is increased, consequently, the absorption by passive 
The interesting question very often asked is "How manyfold is the increased saturation 
solubility?". Data published in the literature or available to us from discussions 
range from 2-14 fold. What are the factors affecting the increase in saturation solubility? 
The factors can be identified when looking at the theoretical background. 
The Kelvin equation describes the increase in the vapor pressure of droplets in a 
gas medium as a function of their particle size, i.e. as a function of their curvature: 
, P -y*VL*cos6 
[pj~ rK*RT 
Fig. 1. The Kelvin equation. 
P = vapor pressure 
Po = equilibrium pressure of a flat liquid surface 
y = surface tension 
VL = molar volume 
cos(#) = contact angle 
rK = radius of droplet 
R = universal gas constant 
T = absolute temperature (K) 
The vapor pressure increases with increasing curvature of the surface, that means 
decreasing particle size. Each liquid has its compound specific vapor pressure, 
thus the increase in vapor pressure will be influenced by the available compoundspecific 
vapor pressure. The situation of a transfer of molecules from a liquid phase 
(droplet) to a qas phase is in principal identical to the transfer of molecules from a 
solid phase (nanocrystal) to a liquid phase (dispersion medium). The vapor pressure 
is equivalent to the dissolution pressure. In the state of saturation solubility, 
there is an equilibrium of molecules dissolving and molecules recrystallizing. This 
equilibrium can be shifted in case the dissolution pressure increases, thus increasing 
the saturation solubility. Identical to liquids with different vapor pressures under 
normal conditions (micrometer droplet size), each drug crystal has a specific dissolution 
pressure in micrometer size. 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 311 
Relative vapor pressure at 293 K 
Droplet size [u,m] 
Fig. 2. Comparison of the relative increase in vapor pressure between water, ether and oleic 
acid (calculated using the Kelvin equation) as a function of the droplet size (with permission 
The important question is how the dissolution pressure changes, depending on 
the specific dissolution pressure of each compound and on the particle size. Model 
calculations were performed applying the Kelvin equation to compounds with 
different vapor pressures (droplets) as a function of droplet size (Fig. 2). Liquids 
with low medium and high vapor pressure were selected, such as oleic acid as an 
oil, water and ether. The important result for a drug formulation was: 
1. The increase in vapor pressure is more pronounced for compounds having 
a priori a low vapor pressure. Applied to solid compounds, increase in dissolution 
pressure will be more pronounced for compounds having a priori a low 
dissolution pressure, i.e. the relative increase is highest for poorly soluble drugs. 
2. The increase in vapor pressure is exponential, with a very pronounced increase 
occurring at droplet sizes below 100 nm. 
Figure 3 shows a calculated increase for barium sulfate as solid model compound. 
3.3. Does size really matter? 
Transferring this to drug nanocrystals means that really smart crystals with highest 
increase in saturation solubility should have a size of e.g. 50 nm or 20-30 nm. 
From this, it can be concluded that the slogan "size matters" is correct regarding 
the increase in saturation solubility, and consequently, the increase in dissolution 
312 Muller & Junghanns 
Saturation solubility of BaS04 in water at 293 K 
C^ 1,0- 
1 0,8- 
c 0,6- 
jjj 0,4- 
K 0,0- 
- 0 , 2 - 
. BaS04 properties: 
\ M = 233.40 g/mol 
\ p = 4.50 g/cm3 
\ cs = 2.22 mg/L 
\ o = 26.7mN/m 
-> i- i i i -! n - ' i i i . . . i i i . |  i 1 '' I ' ' ' '' ' ' ' 111 
0,1 1 10 100 
Drug size [jam] 
Fig. 3. Increase in saturation solubility of BaS04 in water as a function of the particle size 
calculated using the Kelvin equation (with permission after4). 
velocity caused by a higher cs. It needs to be kept in mind which blood profile is 
anticipated with a certain drug. In many cases, too fast a dissolution is not desired 
(creation of high plasma peaks, reduction of tmax). There is the request to combine 
drug nanocrystals with traditional controlled release technology (e.g. coated pellets) 
to avoid too fast a dissolution, too high plasma peaks, too early a tmax and to 
reach prolonged blood levels. To summarize, the optimal drug nanocrystal size will 
depend on: 
1. Required blood profile 
2. Administration route 
In the case of i.v. injected nanocrystals, the size should be as small as possible in case 
the pharmacokinetics of a solution should be mimicked. In the event that a targeting 
is the aim (e.g. to the brain by PathFinder technology,5 the drug nanocrystals should 
possess a certain size to delay dissolution and to give them the chance to reach the 
blood-brain barrier (BBB) for internalization by the endothelial cells of the BBB.6 
3.4. Effect of amorphous particle state 
It is well known that amorphous drugs possess a higher saturation solubility, 
compared with crystalline drug material. A classical example from the literature 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 313 
is chloramphenicol palmitate. The polymorphic modification I has a solubility of 
0.13, the high energy modification II a solubility of 0.43 and the amorphous material 
of 1.6mg/mL.7'8 The same is valid for drug nanoparticles, amorphous drug 
nanoparticles possess a higher saturation solubility, compared with equally sized 
drug nanocrystals in the crystalline state. Therefore, to reach highest saturation 
solubility increase, a combination of nanometer size and amorphous state is ideal. 
However, prerequisite for exploitation in pharmaceutical products is that the amorphous 
state can be maintained for the shelf life of the product. 
4. Production Methods 
4.1. Precipitation methods 
4.1.1. Hydrosols 
The Hydrosol technology was developed by Sucker and the intellectual property 
owned by the company Sandoz, now known as Novartis.9,10 It is basically a classical 
precipitation process known to pharmacists under the term "via humida paratum" 
(v.h.p.). This v.h.p. process was employed to prepare ointments containing finely 
dispersed, precipitated drugs. The drug is dissolved in a solvent, the solvent added 
to a non-solvent leading to the precipitation of finely dispersed drug nanocrystals. 
A problem associated with this technology is that the formed nanoparticles need to 
be stabilized to avoid growth in micrometer crystals. In addition, the drug needs 
to be soluble at least in one solvent. This creates problems for the newly synthesized 
or discovered drugs, being poorly soluble in water and simultaneously in 
organic media. Lyophilization is recommended to preserve the particle size.1 To 
our knowledge, this technology has not been applied to a product to date. 
4.1.2. Amorphous drug nanoparticles (NanoMorph) 
Depending on the precipitation methodology, drug nanoparticles can be generated 
which are in the amorphous state. A nice example are carotine nanoparticles in food 
A solution of the carotinoid, together with a surfactant and a digestible oil, are 
admixed into an appropriate solvent at a specific temperature. The solution is mixed 
with a protective colloid. This tranforms the hydrophilic solvent components into 
the water phase and the hydrophobic phase of the carotinoid forms a monodisperse 
phase. X-ray analysis after subsequent lyophilization shows that approximately 
90% of the carotinoid is in the amorphous state.11 
Amorphous precipitation technology is used by the company Soliqs and the 
technology is advertised under the tradename NanoMorph. The preservation of 
314 Miiller & Junghanns 
the amorphous state could be achieved successfully for food products. To exploit 
the amorphous technology for pharmaceutical products, the stricter requirements 
for pharmaceuticals need to be met. 
4.2. Homogenization methods 
4.2A. Microfluidizertechnology 
The previous Canadian company RTP (Montreal, now Skyepharma Canada Inc.) 
employed the microfluidizer to homogenize drug suspensions. The microfluidizer 
is a jet stream homogenizer of two fluid streams collied frontally with high velocity 
(up to 1000m/sec)12 under pressures up to 4000 bar. There is a turbulant flow, 
high shear forces, particles collied leading to particle diminution to the nanometer 
range.13-15 The high pressure applied and the high streaming velocity of the lipid 
can also lead to cavitation additionally, contributing to size diminution. The patent 
describes examples requiring up to 50 passes through the microfluidizer to obtain 
a nanosuspension.16 Sometimes, up to 100 cycles are required when applying the 
microfluidizer technology. This does not pose any problem on the small lab scale, 
but it is not production friendly for larger lab scale. The dispersion medium is water. 
4.2.2. Piston-gap homogenization in water (Dissocubes) 
In 1994, Mueller et al.17-18 developed a high pressure homogenization method 
based on piston-gap homogenizers for drug nanosuspension production. Dispersion 
medium of the suspensions was water. A piston in a large bore cylinder creates 
pressure up to 2000 bar. The suspension is pressed through a very narrow 
ring gap. The gap width is typically in the range of 3-15 micrometer at pressures 
between 1500-150 bar. There is a high streaming velocity in the gap according to the 
Bernouli equation.19 Due to the reduction in diameter from the large bore cylinder 
(e.g. 3 cm) to the homogenization gap, the dynamic pressure (streaming velocity) 
increases and simultaneously decreases the static pressure on the liquid. The liquid 
starts boiling, and gas bubbles occur which subsequently implode, when the suspension 
leaves the gap and is again under normal pressure (cavitation). Gas bubble 
formation and implosion lead to shock waves which cause particle diminution. 
The patent describes cavitation as the reason for the achieved size diminution.17,20 
Piston-gap homogenizers which can be used for the production of nanosuspensions 
are e.g. from the companies APV Gaulin, Avestin or Niro Soavi. The technology was 
aquired by Skyepharma PLC at the end of the 90s and employed in its formulation 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 315 
4.2.3. Nanopure technology 
For oral administration, the drug nanosuspensions themselves are, in most cases, 
not the final products. For patient's convenience, the drug nanocrystals should be 
incorporated in traditional dry dosage form, e.g. tablets, pellets and capsules. An 
elegant method to obtain a final formulation directly is the production of nanocrystals 
in non-aqueous homogenization media. Drug nanocrystals dispersed in liquid 
polyethylene glycol (PEG) or oils can be directly filled as drug suspensions into gelatine 
or HPMC capsules. The non-aqueous homogenization technology was established 
against the teaching that cavitation is the major diminution force in high 
pressure homogenization. Efficient particle diminution could also be obtained in 
non-aqueous media.24-30 
To prepare tablets or pellets, the dispersion medium of the nanosuspension 
needs to be removed, i.e. in general, evaporated. Evaporation is faster and possible 
under milder conditions when mixtures of water with water miscible liquids are 
used, e.g. water-ethanol. To obtain isotonic nanosuspensions for intravenous injection, 
it is beneficial to homogenize in water-glycerol mixtures. The IP owned by 
Pharmasol covers, therefore, water-free dispersion media (e.g. PEG, oils) and also 
water mixtures. 
4.3. Combination Technologies 
4.3.1. Microprecipitation and High Shear Forces (NANOEDGE) 
The Nanoedge technology by the company Baxter covers a combination of precipitation 
and subsequent application of high energy shear forces, preferentially high 
pressure homogenization with piston-gap homogenizers.31 As outlined in Sec. 4.1.1, 
the precipitated particles have a tendancy to grow. According to the patent by Kipp 
et ah, treatment of a precipitated suspension with energy (e.g. high shear forces) 
avoids particle growth in precipitated suspensions (= annealing process). The 
relative complex patent description can be summarized in a simplified way that 
the subsequent annealing stabilizes the obtained particle size by precipitation. As 
described in Sec. 4.1.2, precipitated particles can be amorphous or partially amorphous. 
This implies the risk that during the shelf life of a product, the amorphous 
particles can recrystalize, leading subsequently to a reduction in oral bioavailability 
or a change in pharmacokinetics after intravenous injection. The annealing process 
by Baxter converts amorphous or partially amorphous particles to completely crystalline 
31 6 Muller & Junghanns 
4.3.2. Nanopure XP technology 
An important criteria for a technology is its scaling up ability and the possibility 
to produce on large scale, applying "normal" production conditions. The number 
of 50-100 passes through a homogenizer as partially required for the microfluidizer 
technology16 is not production friendly. Piston-gap homogenizers (Sec. 4.2.2) 
proved to be more efficient, typically between 10-20 homogenization cycles are 
sufficient to obtain a nanosuspension. However, it would of course be desirable 
to apply even less homogenization cycles, reducing production time, potential 
product contamination by wearing of the machine and production costs. Pharmasol 
developed a new combination process, Nanopure XP (Xtended Performance)32 
leading to: 
1. Identical particle sizes compared with high pressure homogenization in water 
(Sec. 4.2.2), but at half the cycle numbers or less. 
2. Lower particle sizes at identical cycle numbers. 
The process is again a combination technology, a pre-treatment step is followd 
by a high pressure homogenization step, typically performed with a piston-gap 
homogenizer.34'35 The code for this homogenization technology is H42. Figure 4 
LD - volume size distribution 
cycle 15 
new (H42) 
cycle 40 
old technology 
Fig. 4. Comparison of the old homogenization technology (homogenization in water, 
piston-gap homogenizer) on the right side to the new technology on the left side, presented 
are the laser diffractometry (LD) diameters 50%, 90% and 99% (volume distribution, Coulter 
LS230, Beckman-Coulter/Germany) (with permission after33). 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 31 7 
demonstrates the efficiency of method processing a very hard drug material. Applying 
the novel H42 technology leads to distinctly smaller crystals after just 15 cycles, 
compared with the "old" technology of applying 40 cycles. 
5. Application Routes and Final Formulations 
5.1. Oral administra tion 
Most attractive regarding regulatory and commercial aspects is the oral administration 
route. Compared with parenteral administration, the regulatory hurdles are 
much lower. In addition, the patient prefers an oral dosage form, that is why oral 
products possess the largest percentage of the pharmaceutical market. However, 
for the oral administration route, it is generally necessary to transfer the liquid 
nanosuspension into a solid dosage form. 
Aqueous nanosuspensions can be used as a granulation fluid for producing 
tablets or as a wetting agent for pellet production. In addition, spray drying can be 
performed in order to obtain a product which can subsequently be processed to oral 
products. The first nanosuspension product in the market was Rapamune, introduced 
in 2000 by the company Wyeth. Rapamune is available on the market as 
oral solution, and alternatively as tablet. The tablet is more user-friendly. Comparing 
the oral bioavailabilities of solution and nanocrystal tablet, the bioavailability 
of the nanocrystals is 21 % higher compared with the solution. The oral single dose 
of Rapamune is 1 or 2 mg, the total tablet weight being 364 mg for 1 mg formulation 
and 372 mg for the 2 mg formulation, meaning that it contains a very low 
percentage of its total weight as nanocrystals. An important point is that the drug 
nanocrystals are released from the tablet as ultrafine suspension. In the event that 
crystal aggregation takes place to a pronounced extent, the dissolution velocity, and 
subsequently, the oral bioavailability of the BSCII drugs will be reduced. Therefore, 
there is an upper limit to load tablets with nanocrystals. In case the limit is exceeded 
and nanocrystals get in contact with each other within the excipient mixture of the 
tablet, the nanocrystals might fuse to larger crystals under the compression pressure 
during tablet production. For drugs with a low oral single dose such as Sirolimus 
in Rapamune, incorporation into tablets causes little issues. A total nanoparticle 
load of less than 1% is well below the percentage being critical.36 
The second product on the market was Emend, introduced in 2001 by the 
Company Merck. The drug Aprepiptant is for the treatment of emesis (single dose is 
either 80 or 125 mg). Aprepiptant will only be absorbed in the upper gastrointestinal 
tract. Bearing this in mind, nanoparticles proved to be ideal in overcoming this narrow 
absorption window. The large increase in surface area due to nanonization leads 
to rapid in vivo dissolution, fast absorption and increased bioavailability.37,38 The 
formulation of a tablet from micronized bulk powder made higher doses necessary, 
31 8 Muller & Junghanns 
leading to increased side effects.39 The drug nanocrystals are contained within the 
hard gelatin capsules as pellets. Aprepiptant was formulated as capsules for it to be 
user friendly by healthcare providers and patients, and on the other hand, to make 
it applicable as pellets via a stomach tube. Currently, studies are being undertaken 
to evaluate the change in pharmacokinectics (if any) between the pellets and the 
All nanocrystals in these first two products were produced using the pearl 
mill technology by Nanosystems/Elan. The prerequiste was the bioavailability of 
sufficient large scale production facilities for the respective product. In general, 
the candidates of first choice for nanosuspension technology are drugs with a relatively 
low dose. It is interesting that drugs such as Naproxen are formulated as 
nanosuspension (e.g. for fast action onset and reduced gastric irritancy),40 however, 
it requires more sophisticated formulation technology to ensure the release of the 
drug nanocrystals as fine suspension when incorporated in a tablet in a relatively 
high concentration of a single dose of 250 mg. The tablet size (weight) has to be 
acceptable for the patient and that a dosing with two tablets should be avoided, for 
reasons of patient's compliance and marketing purposes. 
Alternatively, to aqueous nanosuspensions, nanosuspensions in nonaqueous 
media can be produced by the Nanopure technology (Pharmasol). Nanocrystals 
dispersed in liquid PEG or oil can be directly filled into gelatine or HPMC capsules.25 
It saves the step of water removal and subsequent dispersion of the powder in a 
liquid capsule filling medium. 
The Nanopure technology also allows production of nanocrystals in melted 
PEG (at 60 C). After solidification of the PEG nanosuspension, the drug nanocrystals 
are fixed (and kept seperated) in the solid PEG matrix. The solidified drug 
nanocrystal containing PEG can either be milled into powders and filled into the 
capsules, or alternatively, the hot liquid PEG nanosuspension can be directly filled 
into the capsules (Fig. 5, upper). 
Instead of using aqueous nanosuspensions as fluids for the wet granulation 
process or extrusion of pellet mass, the nanosuspensions can be converted into a 
dry powder which is subsequently further processed into a tablet or a capsule. It 
also appears attractive to package such powders in sachets for redispersion in water 
or soft drinks prior to oral administration. Spray-drying is the only feasable costeffective 
way to produce such powders. An attractive approach is the production 
of so-called "compounds" as described in the direct compress technology.41 The 
term "compound" does not mean a chemical compound; in powder technology, 
"compounds" are defined as freely flowable granulate powders. In the direct compress 
technology, water-insoluble polymeric particles (e.g. Eudragit RSPO, ethylcellulose) 
are dispersed in the aqueous drug suspension, and lactose is dissolved. 
The mixture is a freely flowable compound yielded by spray-drying. The lactose 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 319 
Fig. 5. Gelatin capsules filled directly with hot liquid PEG nanosuspension, solidification 
takes place in the capsules (upper) or filling of the capsules with milled solidified PEG 
nanosuspension (lower). (From Ref. 36 with permissions.) 
is responsible for the good flowing properties. The water-insoluble polymeric 
particles also contribute to the formation of flowable granules, while at the same 
time allowing the compound to be compressed in a direct compaction process into 
tablets. The polymers form the matrix structure of the tablet. Depending on the percentage 
of the insoluble polymeric particles added, the resulting tablets may disintegrate 
fast or present a prolonged release system. A prolonged release of dissolving 
nanocrystal is desired in the case of high plasma that peaks at very early times (short 
tmax) and a targeted sustained blood level. Alternatively, the drug nanocrystal compound 
can be filled into hardgelatine capsules. Due to the presence of lactose and 
surfactant from the original nanosuspension, the compounds disperse relatively 
fast in liquids. Figure 6 shows the dispersion process of a compound after layering 
it on the surface of water in a beaker. As outlined above, efficient release and redispersion 
of the drug nanocrystals in a fine, nonaggregated state is a prerequisite for 
benefiting fully from the drug nanocrystal features.42 
5.2. Parenteral administration 
Intravenous administration is the second frequently investigated route. The 
company Baxter, with its technology NANOEDGE, is presently focusing on 
intravenous nanosuspensions. They investigated Itraconazole nanosuspensions 
intensive.44 It could be nicely shown that the side effects of the commercial product 
Sporanox could be distinctly reduced by the administration of a nanosuspension. 
The nephrotoxicity of Sporanox is not caused by the drug, but by the excipient 
320 Miiller &/unghanns 
0 sec 15 sec 30 sec 60 sec 120 sec 
Fig. 6. Dispersion of a drug nanocrystal compound as a function of time after layering it 
on the surface of water in a beaker (with permission after43) (Compound: Aquacoat 40%, 
Lactose, 60%.) 
used for solubilizing the drug, the hydroxypropyl-^-cyclodextrin.45'46 The itraconzole 
nanosuspension was stabilized with Tween 80 surfactant being well tolerated 
Administration of nanosuspsensions into body cavities is also of great interest, 
e.g. to increase the tolerability of the drug, to achieve a local treatment or to have a 
depot with slow release (e.g. into the blood). It could be shown that intraperitonal 
administration of a nanosuspension was well tolerated, whereas administration of a 
macrosuspension leads to irritancy [azodicarbonamide (ADA), unpublished data]. 
Intraperitonal administration can be used for local treatment or to obtain a depot 
with prolonged release into the blood. Interesting therapeutic targets include local 
inflammations, e.g. in joints. For instance, arthritic joint inflammations are caused 
by secretion products of activated macrophages. An interesting approach is therefore 
the administration of a corticoid nanosuspension directly into the joint capsule. 
The drug particles will be phagocytosed, the drug dissolves and reduces the hyperactivity 
of the macrophages. This concept is not new, being adopted by the company 
Boots in the 80s in an attempt to incorporate the corticoid prednisolone into polymeric 
nanoparticles made from PLA-GA-copolymer.47 The particle load (polymer 
load) required to achieve a therapeutic drug level was being calculated. However, 
incubating macrophage cell cultures with the required particle concentration lead 
to cytotoxicity. The concept could not be realized, as it cannot occur with drug 
nanocrystals since no carrier polymer to required and present. 
Producing parenteral products with drug nanocrystals has to meet higher 
regulatory hurdles and product quality standards distinctly. The produced drug 
nanosuspensions need to be terminally sterilized or alternatively produced in an 
aseptic process. In principal, sterilization is possible by autoclaving. However, the 
increase in temperature can reduce hydration of steric stabilizers, thus leading to 
some aggregation during the sterilization process. Gamma irradiation is a priori a 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 321 
non-preferred process by the industry due to the necessary analytics (i.e. proof of 
absence of toxic irradiation products). In addition, it was also observed that irradiation 
can cause aggregation not by directly interacting with the drug nanocrystals, 
but with the stabilizing surfactant. Irradiation of Tarazepide nanosuspension leads 
to aggregation; simultaneously, a decrease in zeta potential also occurred during 
the irradiation process. A decrease in zeta potential, i.e. electrostatic repulsion, was 
considered as the cause for the aggregation process. It can be concluded that the 
production of drug nanosuspensions in an aseptic, controlled process has to be 
preferred, compared with the terminal sterilization by irradiation. The aseptic production 
process can be validated and documented relatively easy, therefore, being 
simpler to handle as an irradiation sterilization with accompanied analytics. 
5.3. Miscellaneous administration routes 
Oral and parenteral/intravenous routes are the ones in which developments are 
focusing, clearly due to the commercial background and the relation between the 
development costs for a market product versus its potential annual sales. However, 
drug delivery could also be improved when using drug nanocrystals for pulmonary 
and ophthalmic adminstration or dermal application. 
Poorly soluble drugs could be inhaled as drug nanosuspension. The drug 
nanosuspension can be nebulized using commercially available nebulizers.48,49 Disposition 
in the lungs can be controlled via the size distribution of the generated 
aerosol droplets. Compared with microcrystals, the drug is more evenly distributed 
in the droplets when using a nanosuspension. The number of crystals are higher, 
consequently, the possibility that one or more drug crystals are present in each 
droplet is higher. 
It could be shown that nanoparticles possess a prolonged retention time in the 
eye, most likely due to their adhesive properties.50-52 From this, poorly soluble 
drugs could be administered as a nanosuspension. However, the major obstacles 
are the commercial considerations. In many cases, the sales volume do not justify 
the costs for the development of a new market product. This is especially the case 
when a company has already a drug formulation which might be less efficient, but 
is already a product on the market. The price achievable with an improved product 
is not sufficiently high to cover the development costs of this new product. An 
additional major obstacle for the development of such improved products is the cost 
reduction policy of the healthcare systems worldwide. A longer treatment time with 
a less efficient product might still be less expensive for the healthcare system than a 
shorter treatment time with a more efficient, but distinctly more expensive product. 
The same is valid for dermal products. Sales per product are lower compared 
with e.g. oral products, as the dermal market is smaller. Dermal nanosuspensions 
322 MiJller&Junghanns 
are mainly of interest if conventional formulation technology fails or if it is distinctly 
less efficient. Dermal drug nanosuspensions lead to a supersaturated system 
because of their increased saturation solubility. The higher concentration gradient 
between topical formulation and skin can improve drug penetration into the skin. 
In addition, because of their small size, drug nanocrystals could target the hair follicle 
by protruding into the gap around the hairs. This was illustrated in solid lipid 
nanoparticles of a similar size.53 Adhesive properties of drug nanocrystals are also 
an area of interest. Adherence to the skin reduces the "loss" of drug to the environment/
third persons. This is especially so in the event that highly active compounds 
are applied, e.g. hormones. For this reason, the drug estradiole was incorporated 
into solid lipid nanoparticles to better localize it on the skin.54 
6. Nanosuspensions as Intermediate Products 
As described above, nanosuspensions can be produced such that nanocrystals 
appear in final products. Alternatively, drug nanosuspensions can be used as intermediate 
product, i.e. the drug nanocrystals do not appear in the final product. 
Recently, the SolEmuls technology was developed to produce drug-loaded emulsions 
for intravenous injection, i.e. localizing poorly soluble drugs in the interfacial 
layer of lecithin emulsions.55-57 The applicability of the technology has been proven 
for several drugs including amphotericin B,58 itraconazole,59,60 ketoconazole,61 
and carbamazepine,62'63 among others. The drug Amphotericin B is on the market 
as a solution (Fungizone), but also in liposomes (Ambisome); the latter 
having the benefit of reduced nephrotoxicity.64 Liposomes are relatively expensive 
(daily treatment costs approximately EUR 1000-200064,65), therefore Amphotericin 
B was incorporated into parenteral emulsions. These emulsions can also 
reduce nephrotoxicity,66 but for their production, it was necessary to use organic 
solvents. Egg lecithin and amphotericin B were dissolved in an organic solvent, 
the solvent evaporated and the obtained drug-lecithin mixture was used to produce 
an o /w emulsion. In these emulsions, amphotericin B was located in the 
interfacial lecithin layer as Amphotericin B is simultaneously poorly soluble in 
water and in oils.67 There were also attempts to incorporate amphotericin B in the 
emulsion by simply adding Amphotericin B powder to the emulsion and subsequently 
shaking it. However, even shaking for 18 hours with 1800rph was unable 
to completely dissolve the Amphotericin B. The reason was simply due to its low 
solubility in the water, and the dissolution velocity was also extremely low, i.e. 
the process of dissolution and redistribution into the lecithin layer takes too long 
for it to be used in pharmaceutical production. The problem was solved by the 
SolEmuls technology, i.e. simple co-homogenization of oil droplets and microcrystals. 
For a de novo production, a coarse pre-emulsion of lecithin stabilized 
Drug Nanocrystals/Nanosuspensions for the Delivery of Poorly Soluble Drugs 323 
oil droplets in water is prepared, the drug powder is admixed under stirring 
and the obtained hybrid suspension subsequently homogenized at 600 bar (pressure 
being in the range to be used in pharmaceutical production lines). The high 
streaming velocities in the homogenization process lead to fast dissolution of 
the drug microcrystals and the re-distribution into the interfacial lecithin layer 
(Fig. 7). 
Depending on the size of the drug crystals, 5-10 homogenization cycles are 
required. The number of homogenization cycles can be reduced when adding the 
. drug not as microcrystals, but as nanocrystals in the form of a nanosuspension. A 
concentrated nanosuspension is prepared (e.g. 20-30% solid content) and added to 
the pre-emulsion. Ideally the nanosuspension is also stabilized by lecithin, i.e. the 
same emulsifier for the suspension and the emulsion. Alternatively, intravenously 
accepted stabilizers such as Tween 80 or Poloxamer 188 can be used. They are 
accepted intravenously without posing any regulatory issues. In addition, mixing 
the emulsion and nanosuspension at a ratio of 10:1 or higher will dilute the stabilizer 
concentration used in the nanosuspension by at least a factor of 10, meaning that 
in the final product, the nanosuspension surfactant concentration is typically 0.1 
or 0.01%. The question might arise as to why an emulsion should be prepared 
using a nanosuspension as an intermediate product, when it can administer the 
nanosuspension itself intravenously? 
r ~\ 
iecithin / drug mixture 
O rf-Q 
O ' o T 
direct production 
through highpressure 
organic solution t lecithin + drug 
drug crystal or 
Fig. 7. Drug incorporation through various methods in comparison. Left: traditional 
attempt of shaking or alternatively use of organic solvent; Right: SolEmuls process. 
324 Muller & Junghanns 
The reason is that drug-loaded parenteral emulsions are already products on the 
market (e.g. Diazepam-Lipuro, Etomidate-Lipuro, etc.), i.e. in a dosage form with 
which the regulatory authorities are already familiar with. Applying the SolEmuls 
technology and using lecithin-stabilized nanosuspension, the final product will 
only contain the excipients of an o /w emulsion for parenteral nutrition, without 
additional excipient plus the drug. It is an accepted known system with regard 
to the excipient status and its perfomance after intravenous injection. In contrast, 
drug nanosuspensions represent a new dosage form not yet present as intravenous 
formulations on the market. Registration of a completely new dosage form for a 
certain administration route is just more complicated and timely than registration 
of a product based on an established, known technology. 
7. Perspectives 
There was a "delayed" acceptance of the nanocrystal technology in the 90s. Pharmaceutical 
companies tried to solve their formulation problems with the traditional 
formulation approaches. However, the increasing number of drugs having a very 
low solubility, and not able to be formulated with these traditional formulation 
approaches, lead to a broad acceptance of the drug nanocrystal technology. This is 
clearly reflected in the increasing number of licensing agreements between companies 
holding nanocrystal IP and a number of medium and large pharmaceutical 
companies. The smartness of the technology is that it can be universally applied 
to practically any drug. Identical to micronization, it is a universal formulation 
principle, but limited to BSC drugs class II. The time between the beginning of 
intensive research in the drug nanocrystal technology and the first products on the 
market was relatively short, about one decade. The value of a formulation principle 
or technology can be clearly judged by looking at the number of products on the 
market, in the clinical phases, and/or the time of entry into the market. Based on 
these criteria, the drug nanocrystal technology is a successful emerging technology. 
Meanwhile, "Big Pharma" also realized the drug nanocrystal value. In combination 
with the further increasing number of poorly soluble drugs, a distinct increase in 
drug nanocrystal-based products on the market can be expected. In many cases, oral 
products will dominate because of the market share, higher sales volumes and less 
regulatory hurdles and quality requirements, compared with parenteral products. 
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Cells and Cell Ghosts as Drug Carriers 
Jose M. Lanao and M. Luisa Sayalero 
1. Introduction 
Microparticle and nanoparticle polymeric systems currently occupy an important 
place in the field of drug delivery and targeting.1 Nevertheless, there are biological 
drug carriers that offer an efficient alternative to such systems. Within the different 
systems of biological carriers, of great importance are cells and cell ghosts, which are 
both efficient and highly compatible systems from the biological point of view, capable 
of providing the sustained release and specific delivery to tissues, organs and 
cells of drugs, enzymatic systems and genetic material. Cell systems such as bacterial 
ghosts, erythrocyte ghosts, polymorphonuclear leukocytes, apoptotic cells, 
tumor cells, dendritic cells, and more recently, genetically engineered stem cells, 
are all examples of how cell systems of very diverse nature can be suitably manipulated 
and loaded with drugs and other substances, to permit specific drug delivery 
in vivo with important therapeutic applications.2-8 Cell carriers for drug delivery are 
used in very different applications such as cancer therapy, cardiovascular disease, 
Parkinson's, AIDS, gene therapy, etc. Table 1 shows the classification of biological 
carriers for drug delivery based on the use of cells and cell ghosts. 
2. Bacterial Ghosts 
Bacterial ghosts are intact, non-living, non-denatured bacterial cell envelopes 
devoid of cytoplasmic contents. They are created by lysis of bacteria, but maintain 
330 Lanao & Sayalero 
Table 1 Kinds of cells and cell ghosts used for drug and gene delivery. 
Cell carrier 
Bacterial ghost 
Erythrocyte ghost 
Engineered stem cells 
Apoptopic cells 
Tumor cells 
Denditric cells 
Tissues, macrophages, cells 
RES, macrophages 
Tumor cells, T cells, 
Tumor cells 
Tumor cells 
T cells 
Encapsulated substance 
Drugs, vaccines, genetic material 
Drugs, enzymes, peptides 
Genetic material 
their cellular morphology and native surface antigenic structures, including their 
bioadhesive properties.3,9 
Bacterial ghosts allow the encapsulation of drugs and other substances, and 
their specific attachment to mammalian tissues and cells. This kind of cell carrier acts 
as a true drug delivery system, allowing the permanency of drugs in the systemic 
circulation to be increased together with tissue-specific targeting. They are thus a 
promising alternative to conventional drug delivery systems such as liposomes or 
The main advantages of bacterial ghosts as delivery systems are the fact that 
they are non-living, i.e. they can act as delivery systems of drugs, antigens or DNA; 
allow specific delivery to different tissues and cell types; and are well captured by 
phagocytic cells and antigen-presenting cells as dendritic cells. Among the drawback 
of bacterial ghosts is the possibility that they might revert to being virulent, the 
possibility of horizontal gene transfer, the stability of the recombinant phenotype, 
and pre-existing immunity against the carrier used.10 
Usually, bacterial ghosts are produced by protein E-mediated lysis of Gramnegative 
bacteria.11 The production of bacterial ghosts is based on the controlled 
expression of the bacteriophage PhiX174-derived lysis gene E. Expression of this 
gene from a plasmid in Gram-negative bacteria leads to the formation of a transmembrane 
lysis tunnel structure that penetrates the inner and outer membranes, 
and is formed by protein E with border values fluctuating between 40-200 nm 
in diameter. Protein E is a hydrophobic protein localized exclusively in the cell 
envelope.12 E-mediated lysis has been achieved in many Gram-negative bacteria.13 
Scanning electron micrographs of E-lysed cells reveal that bacterial ghosts contain 
only one E hole in a bacterial ghost, although in a few cases, there are two holes. 
The cytoplasm is expelled as a consequence of the high osmotic pressure inside the 
Cells and Cell Ghosts as Drug Carriers 331 
cell. The collapse of membrane potential and the release of cytoplasmic components 
such as proteins, nucleic acids, etc occur simultaneously.14 In the case of strains of 
E. coli, this effect occurs within a period of 10 min after the induction of expression.15 
The resulting empty bacterial cell envelope is considered a bacterial ghost. Bacterial 
ghosts show all the morphological, structural and immunogenic properties of 
a living cell.9-15-17 Since bacterial ghosts are derived from Gram-negative bacteria 
that are able to adhere to structures such as fimbriae and lipopolysaccharide, they 
are used for specific binding to human tissue.18 
Bacterial ghost drug-loading is accomplished by the suspension of lyophilised 
bacterial ghosts in a buffered medium containing the drug. The ghosts are then 
subjected to an incubation process varying from 5 to 30 min at 24C. They are then 
washed to remove excess drug.18'19 
In order to prevent rapid leakage of loaded water-soluble drugs or other substances, 
the bacterial ghosts are sealed by fusion of the cell membrane with membrane 
vesicles at the edges of the lysis pore. For the sealing step, the bacterial ghosts 
suspension is incubated in the fusion buffer at 28C for 10 min.18 Figure 1 shows a 
scheme of the production of bacterial ghosts by protein E-mediated bacterial lysis. 
The in vitro release of drugs from loaded bacterial ghosts is performed from a 
suspension of drug-loaded bacterial ghosts that is dialysed through a membrane 
suitable for excluding the ghosts. Dialysis is performed at 28C in PBS buffer.19 The 
concentrations of drug released into the buffer at preset times are quantified using 
an appropriate analytical technique. 
In studies addressing the adherence and capture of loaded bacterial ghosts by 
target cells such as macrophages, human colorectal adenocarcinoma cells (Caco-2) 
or dendritic cells, fluorescent markers such as fluorescein isothiocyanate (FITC) 
are used. These allow adherence to be assessed using fluorescence microscopy 
and flow cytometry techniques.18,19 Macrophages internalize bacterial ghosts to 
a greater extent than Caco-2 cells.18,19 Studies carried out using confocal laser scanning 
microscopy with M. haemolytica ghosts loaded with Doxorubicin have shown 
that the drug was associated with the ghosts membranes and the inner lumen.19 
Denditric cells that are professional phagocytic cells displaying the phagocytic 
capacity of antigens also have a good capacity for capturing bacterial ghosts, allowing 
the latter to be used as a vehicle for immunization and immunotherapy.20 
2.1. Application of bacterial ghosts as a delivery system 
Bacterial ghosts have important therapeutic applications. They can be loaded 
with drugs, proteins and other substances, and can be targeted selectively to 
macrophages, tumors or endothelial cells.10,19 
332 Lanao & Sayalero 
Cytoplasmic content 
Inner Membrane 
Outer Membrane 
Cytoplasmic Membrane 
Protein E-mediated lysis 
E hole (40-200 nm) 
Fig. 1. Production and drug loading of bacterial ghosts. 
Bacterial ghosts have been used as efficient drug delivery systems10 in the 
field of anti-cancer drugs.18 Bacterial ghosts obtained have been used as a delivery 
system of doxorubicin to human colorectal carcinoma cells. Cytotoxicity assays 
revealed that doxorubicin-loaded ghosts show better antiproliferative capacity in 
Caco-2 cells than when free doxorubicin is used at the same concentration.18 Experiments 
have also been carried out using E.coli ghosts containing streptavidin, in order 
to increase the affinity of streptavidin for biotinylated compounds. Streptavidinloaded 
ghosts permit specific targeting to mucosal surfaces of the gastrointestinal 
and respiratory tracts, and also to phagocytic cells.3 Bacterial ghosts have been used 
as veterinary vaccines for the immunization of different animal species.9 
Pasteurella multocida is a pathogen that causes morbidity and mortality in rabbits 
and its importance as a human pathogen has also been recognized. P. multocida 
Cells and Cell Ghosts as Drug Carriers 333 
ghosts have been used to immunize rabbits and mice.17 Similar results have been 
obtained in the immunization of cattle against pasteurellosis using Pasteurella 
haemolytica ghosts.11 
Actinobacillus pleuropneumoniae is a highly contagious microorganism and is 
the cause of porcine pleuropneumonia, infecting 30-50% of pig populations. However, 
Actinobacillus pleuropneumoniae vaccines provide limited protection, since they 
decrease mortality but not morbidity in swine. Comparative studies have been carried 
out on immunization using a aerosol infection model for pigs vaccinated with 
loaded-ghosts or formalin inactivated Actinobacillus pleuropneumoniae bacterins. The 
results obtained showed that immunization with bacterial ghosts is more efficient 
in protecting pigs than bacteria.21,22 
Bacterial ghosts can also be used as carriers of therapeutic DNA or RNA.3/13 
The use of nucleic acid vaccines currently offers a technique for the development of 
prophylactic or therapeutic vaccines, based on the use of DNA plasmids to induce 
immune responses by direct administration of DNA-encoding antigenic proteins 
into animals, and this is also suitable for the induction of cytotoxic T cells.23,24 
Bacterial ghosts loaded with DNA produce a high level of gene expression. They 
can be used to enhance the mucosal immune response to target antigens expressed 
in the bacterial ghost system. They can also be used for the specific targeting of 
DNA-encoded antibodies to primary antigens located in cells.13 Ghosts of Vibrium 
cholerae have been tested as antigen carriers of Chlamidia trachomatis as potential 
vaccines for the control of genital infections produced by this bacteria. Recombinant 
Vibrium cholerae ghosts, previously cloned with a major outer membrane protein of 
C. trachomatis, afforded a high level of protective immunity against Chlamydia in a 
murine model.25,26 Mannheimia haemolytica is a pathogen that causes ovine mastitis. 
M. haemolytica ghosts loaded with plasmid DNA stimulate the elicitation of efficient 
immune responses in mice, with no symptoms of acute or subacute toxicity during 
the experiment.27 
3. Erythrocyte Ghosts 
Erythrocytes constitute the largest population of blood cells and are produced in 
the bone marrow. They are mature blood cells that produce haemoglobin and carry 
out the exchange of oxygen and carbon dioxide between the lungs and the body 
The term "erythrocyte ghost" attempts to define the resulting cell-like structure 
when erythrocytes are subjected to a reversible process of osmotic lysis.28 For more 
than 30 years, many studies, both in vivo and in vitro, have been carried out to explore 
the use of erythrocyte ghosts as delivery systems of drugs and other substances.2 
334 Lanao & Sayalero 
Erythrocyte ghosts are obtained from fresh erythrocytes coming from human 
blood or the blood of different animal species such as the rat, mouse, rabbit, etc, and 
are loaded with different types of substance, mainly drugs, peptides and enzymes, 
using different encapsulation methods. The most frequent methods for collecting 
erythrocyte ghosts are osmosis-based methods such as hypotonic dialysis.2,29 
Autologous erythrocyte ghosts offer a drug delivery system that can act as a 
reservoir of the drug or substance encapsulated, providing the sustained release 
of the drug into the organism together with selective targeting of the drugs to the 
reticuloendothelial system (RES) of the liver, spleen and bone marrow.2 
The main advantages of carrier erythrocytes as drug delivery systems are their 
high degree of biocompatibility, the possibility of encapsulating the drug in a small 
amount of cells, the sustained release of the encapsulated drug or substance into 
the body, the selective targeting to the RES, and the possibility of encapsulating 
substances of high molecular weight such as peptides. Among the drawbacks of 
these systems are the rapid leakage of some drugs out of the loaded erythrocytes 
and other problems related to their standardized preparation, storage and potential 
Erythrocyte ghosts can be obtained by diverse procedures such as hypotonic 
dilution, hypotonic pre-swelling, osmotic pulse, hypotonic hemolysis, hypotonic 
dialysis, electroporation, drug-induced endocytosis and chemical methods.2,30 Of 
the different ways of obtaining carrier erythrocytes, hypotonic dialysis is undoubtedly 
the most frequently used encapsulation method. The reasons why it is so popular 
are its simplicity, its ease of application for a large number of drugs, enzymes 
and other substances, and because it is the method that best conserves the morphological 
and haematological properties of the erythrocyte ghosts obtained. 
Hypotonic dialysis is based on the exposure of red cells to the action of a hypotonic 
buffer, inducing cell swelling and the formation of pores that permit the drug 
to enter erythrocytes by means of a passive mechanism. Figure 2 shows a scheme 
of the production of erythrocyte ghosts using a hypotonic dialysis method. 
Morphological inspection of erythrocyte ghosts is usually performed using 
transmission (TEM) or scanning (SEM) electron microscopy.2 Some physical parameters 
of red cell membranes can also be studied from the diffusion of haemoglobin.28 
The haemolytic methods employed in the production of erythrocyte ghosts normally 
affect the haemolytic volume, surface area and tension.28 Figure 3 shows the 
morphological changes observed by SEM that occur in amikacin-loaded erythrocytes 
due to hypotonic dialysis.31 
Haematological parameters determine the effects of the procedure used to collect 
erythrocyte ghosts on their haematological properties. Among others, parameters 
such as reduced glutathione (GSH), mean corpuscular volume (MCV) or red 
cell distribution width (RDW), may be evaluated using a haematology analyzer. 
Cells and Cell Ghosts as Drug Carriers 335 
Dialysis bag 
suspension T \ 
(Isotonic buffer) 
(Hypertonic buffer) 
(10 min, 37C, pH 7.4) (30 min, 37C, pH 7.4) 
Hypotonic buffer 
(45 min, 4"C, pH 7.4) 
Fig. 2. Production and drug loading of erythrocyte ghosts using a hypotonic dialysis 
o V 
Fig. 3. SEM micrographs of amikacin carrier erythrocytes obtained by hypotonic dialysis31 
(Copyright 2005 from Encapsulation and in vitro Evaluation of Amikacin-Loaded Erythrocytes 
by C. Gutierrez Millan. Reproduced by permission of Taylor & Francis Group, LLC, 
http: / / www. taylorandfrancis .com). 
Erythrocyte ghosts obtained by hypotonic dialysis show a decrease in the mean corpuscular 
volume and an increase in size dispersion.28'29 Erythrocyte ghosts show a 
greater degree of haemolysis than normal erythrocytes.29 
3.1. Applications of erythrocyte ghosts as a delivery system 
Erythrocyte ghosts can be used as potential drug delivery systems for enzymes, 
proteins and peptides, allowing sustained release into the systemic circulation and 
the delivery of these substances into the RES.2 
In vitro release of drugs from loaded erythrocyte ghosts is usually tested using 
autologous plasma or an isoosmotic buffer at 37C; alternatively, a dialysis bag 
may be used.32 The in vitro release of drugs and substances from loaded erythrocytes 
is usually a first-order process, suggesting that the drug crosses the plasma 
membrane through a passive diffusion mechanism.33 However, zero-order release 
336 Lanao & Sayalero 
kinetics from loaded erythrocytes has also been described.34 In vitro studies about 
the release kinetics of different drugs, enzymes and peptides from loaded erythrocytes 
have shown a slow release of the encapsulated substance.2 
When loaded erythrocyte ghosts are administered in vivo, changes in the pharmacokinetics 
of the encapsulated drugs occur, involving a systemic drug clearance 
related to the biological half-life of the erythrocytes.35 Increased serum half-lives 
and the areas under the curve of drugs encapsulated in loaded erythrocyte ghosts, 
in comparison with the free drug, have been observed in animals and humans.36,37 
At the same time, erythrocyte ghosts show a greater accumulation in tissues such 
as liver and spleen.38,39 
Surface treatment of erythrocyte ghosts with substances such as glutaraldehyde, 
ascorbate, Fe(+2), diamide, band 3-cross-linking reagents, trypsin, phenylhydrazine 
and the N-hydroxysuccinimide ester of biotin (NHS-biotin), enhances 
the recognition of erythrocyte ghosts by macrophages in vitro and liver targeting 
in vivo.40~i3 
Red cells may be used as carriers for some drugs such as antineoplastics, antiinfective 
agents, antihypertensives, corticosteroids, etc.2 Thus, carrier erythrocytes 
have been widely studied as delivery systems of antineoplastic drugs for targeting 
the RES located in organs such as liver and spleen. 
Different antineoplastic drugs have been encapsulated in erythrocyte ghosts 
in both in vitro and in vivo experiments.2 Increases have been obtained in average 
survival times in the treatment of mice bearing hepatomas, using methotrexateloaded 
carrier erythrocytes.44 Better recognition and capture of erythrocyte ghosts 
by macrophages have been obtained by using biotinylated erythrocytes containing 
methotrexate,45 by alterations to the membrane using band-3 cross-linkers of erythrocyte 
ghosts containing etoposide,46 or by treatment of erythrocytes containing 
doxorubicin with glutaraldehyde.47 
Anti-infective agents such as gentamicin, metronidazole, primaquine or imizol 
have also been encapsulated in erythrocytes.2 Human erythrocytes containing 
gentamicin have proven to act as an efficient slow-release system in ofco.48,49 
Erythrocyte ghosts containing dexamethasone have been used in vivo in rabbits 
and humans. A sustained release of dexamethasone in vivo in animals and humans 
was observed using carrier erythrocytes. An increased anti-inflammatory effect of 
the drug using carrier erythrocytes was observed in rabbits.50,51 
Moreover, new prodrugs of anti-opioid drugs such as naltrexone and naloxone 
have been encapsulated in erythrocytes to solve stability problems of the primary 
drug within the erythrocyte. The encapsulated prodrugs are transformed into the 
active compound, following their release from erythrocyte ghosts.52 
In the fields of biochemistry and enzymatic therapeutics, the encapsulation 
of enzymes in erythrocytes has been studied in some depth. Enzymatic 
Cells and Cell Ghosts as Drug Carriers 337 
deficiencies or the treatment of specific illnesses may be approached using carrier 
erythrocytes loaded with enzymes. The encapsulation of enzymes in erythrocytes 
solves some of the problems associated with enzyme therapy, such as 
the short half-life deriving from the action of plasma proteases, intolerant reactions, 
and the immunological disorders or allergic problems associated with 
the use of enzymes in therapeutics. In vitro or in vivo studies with enzyme 
carrier erythrocytes have been performed using L-asparaginase,53 hexokinase,54 
alcohol dehydrogenase,55 aldehyde dehydrogenase,56 alcohol oxidase,57 glutamate 
dehydrogenase,58 uricase,59 urokinase,60 lactate 2-mono oxigenase,61 
arginase,62 rhodanase,63 recombinant phosphotriestearase,64 delta-aminolevulinate 
dehydratase,65 urease,66 pegademase,67 brinase68 and alglucerase.69 One of the best 
examples of the use in therapeutics of carrier erythrocytes containing enzymes, is 
that of L-asparaginase encapsulated in human erythrocytes. This has been successfully 
used in the treatment of acute lymphoblastic leukaemia in paediatrics.70 
Erythrocyte ghosts may act as carrier systems for the delivery of peptides 
and proteins. One of the main therapeutic applications of carrier erythrocytes in 
this field is that of anti-HIV peptides. Nucleoside analogues successfully inhibit 
the replication of immunodeficiency virases. In view of the importance of the 
monocyte-macrophage system in infection by HIV-1, it would be of maximum 
therapeutic interest to have available, the specific delivery of these therapeutic 
peptides into macrophages, which act as an important reservoir for the virus. Carrier 
erythrocytes containing anti-HIV peptides such as azidothimidine (AZT) and 
didanosine (DDI), significantly reduced the pro-viral DNA content in comparison 
with the administration of free peptides in a murine AIDS model.71 Similar 
results have been obtained with 2',3'-dideoxycytidine 5'-triphosphate'(ddCTP),72 
2',3'-dideoxycytidine (ddCyd)73 and AZT prodrugs74 encapsulated in erythrocytes. 
Anti-neoplastic peptides such as 2-fluoro-ara-AMP (fludarabine) and 5'- 
fluoro-2'-deoxyuridine 5'-monophosphate (FdUMP), a pro-drug of 5-fluro-2'- 
deoxyuridine (FdUrd), have been encapsulated in human carrier erythrocytes, 
behaving as a slow-release delivery system.75,76 
Macrophage uptake in vitro of antisense oligonucleotides may be increased by 
using carrier erythrocytes.77,78 Other peptides, such as erythropoietin,79 heparin,80 
dermaseptin S3,81 interleukin-382 or vaccines,83 have also been encapsulated in erythrocytes 
to increase their stability,84 acting as a slow release system with a prolonged 
half-life,80,84 or for their specific targeting to bacterial membranes.85 
Erythrocyte ghost derivatives can also be used as drug delivery systems. 
Nanoerythrosomes are erythrocyte membrane derivatives formed by spheroid vesicles, 
obtained by consecutive extrusion under nitrogen pressure through a polycarbonate 
filter membrane of a erythrocyte ghost suspension to produce small vesicles 
having an average diameter of 100 nm. In vitro and in vivo studies, carried out with 
338 Lanao & Sayalero 
nanoerythrosomes loaded with daunorubicin, have shown that when linked covalently 
to nanoerythrosomes, the drug produces slow release of daunorubicin to the 
organism over a prolonged period of time and also that, in comparison with the free 
drug, cytotoxicity is greater.86 The advantage of nanoerythrosomes, as compared 
with erythrocyte ghosts as drug delivery system, is that the former are able to escape 
from the reticuloendothelial system faster.86,87 In vitro studies have shown that the 
nanoerythrosome-daunorubicin complex is rapidly adsorbed and phagocytosed by 
macrophages.88 Liver, spleen and lungs are the organs in which nanoerythrosomes 
show the greatest capacity of accumulation.89 
Another derivative of erythrocyte ghosts are reverse biomembrane vesicles 
loaded with drugs.90 Reverse biomembrane vesicles are produced by spontaneous 
vesiculation of the ghost erythrocyte membrane by endocytosis, using an appropriate 
vesiculating medium, producing small vesicles containing the drug within the 
parent ghost. In vivo studies carried out using reverse biomembrane vesicles from 
erythrocyte ghosts loaded with doxorubicin in rats have revealed increases in the 
half-life and bioavailability of the drug, the liver and spleen, being the main organs 
for the clearance of this drug delivery system.90 
4. Stem Cells 
In gene therapy, a therapeutic transgene is introduced into the patient with a view of 
supplementing the functions of an abnormal gene. To achieve the delivery of genetic 
material into the target cell, it is necessary to have a suitable carrier. One important 
aim in the field of gene therapy is the design and development of gene carriers that 
encapsulate and protect the nucleic acid, and selectively release the vector/nucleic 
acid complex to the target tissue, so that the genetic material will be released at the 
cellular level later. In practice, there are several ways to achieve this. The first is 
through the use of modified viruses containing the genetic material of interest. The 
use of viruses for gene delivery has some drawbacks since it is limited to specific 
cells susceptible to being infected by the virus, and also the administration itself 
of the virus, has some immunological problems among others.91-93 The second 
alternative is to use living cells modified genetically, such as stem cells, to deliver 
transgenic material into the body.8,94 
Stem cell therapy is a new form of treatment, in which cells that have died or 
lost their function are replaced by healthy adult stem cells. One advantage of this 
kind of cell is that it is possible to use samples from adult tissues or cells from the 
actual patient, for culture and subsequent implantation. 
Within the framework of stem cell research, the use of stem cells as delivery 
systems is a novel and attractive technique in the field of gene therapy, in which the 
cells of the patients themselves are genetically engineered, in order to introduce a 
therapeutic transgene used to deliver the genetic material. A promising therapeutic 
Cells and Cell Ghosts as Drug Carriers 339 
strategy is the use of stem cells such as lymphocytes or fibroblasts as drug delivery 
systems. Experimental studies using stem cells as such systems have been tested 
in different therapeutic applications, especially in the field of cancer therapy. Considering 
the affinity of stem cells for tumor tissue, engineered stem cells have been 
successfully used for direct drug delivery to cancer cells.8'94 In vitro cultures have 
been made of human mesenchymal stem cells from bone marrow that are transduced 
with an adenovirus vector carrying the human interferon beta-gene, which 
exerts therapeutic action against cancer. Engineered stem cells administered in vivo 
allow the delivery of the genetic material to cancer cells. This new drug delivery 
system has proven to be efficient in the treatment of experimental neoplasms, such 
as lung cancer, in mice.94 Figure 4 shows a scheme of the application of stem cells 
as carriers for gene delivery in experimental cancer therapy. 
In vivo studies have also been carried out with neural stem cells engineered 
using adenoviral vectors to express interleukm-12, an oncolytic gene, whose efficiency 
has been demonstrated in the treatment of intracranial malignant gliomas 
in mice.8,95 
The used of haematopoietic stem cells has allowed antiviral genes to be introduced 
in both T cells and macrophages for the treatment of AIDS.96 The use of 
stem cells as vehicles for gene therapy has also been suggested for the treatment of 
ischaemic heart disease,97 
Stem cells have also been employed in the field of antiepileptic therapy. Glial 
precursor cells, which release adenosine, have been derived from adenosine kinase 
embryonic stem cells. In these experiments, the fibroblasts were engineered to 
release adenosine by inactivating adenosine metabolising enzymes. After encapsulation 
within polyethersulfone hollow-fibre capsules, and the introduction into 
Stem cells Interferon 
Engineered i vitro 
Stem cells expansion 
Ficoll ; ', 
Fig. 4. Application of stem cells as carriers for gene delivery in experimental cancer. 
340 Lanao & Sayalero 
the brain ventricles in a rat epilepsy model, the local release of adenosine allows 
drug-resistant focal epilepsy to be treated. These engineered cells were shown to 
suppress seizure activity.98-99 
5. Polymorphonuclear Leucocytes 
Polymorphonuclear leucocytes (PMN) can be used as carriers of antibiotics in view 
of their selective targeting to sites of infection. Simply incubating PMN in the presence 
of high concentrations of antibiotic for 1 hr at 37 C guarantees cell loading with 
the antibiotic. PMN loaded with the macrolide azithromycin have been found to be 
efficient in an in vitro model that permits the delivery of the antibiotic in a bioactive 
form to Chlamydia inclusions in polarized human endometrial epithelial (HEC-1B) 
cells infected with Chlamydia trachomatis. PMN carriers allow the accumulation of 
large amounts of antibiotic in endometrial epithelial cells and its retention over 
long periods of time.4 
6. Apoptopic Cells 
Programmed cellular death or apoptosis is a process that is controlled genetically 
in which the cells induce their own death in response to different types of stimulus 
such as the binding of death-inducing ligands to cell surface receptors. 
A new strategy for drug delivery, called apoptopic induced drug delivery 
(AIDD), allows drug delivery to tumor cells upon the initiation of apoptosis by 
using a biological mechanism to achieve drug delivery.5 This new system is based 
on the fact that apoptosis produces many changes in cell morphology that can be 
taken advantage of to achieve drug delivery. 
Apoptosis is reflected in enhanced membrane permeability, which favors the 
release of the encapsulated drug from the apoptotic cells to the tissue. Phagocytosis 
of drug- loaded apoptotic carrier cells by tumor cells facilitates the localization 
of the drug within the tumor cell. One advantage of the apoptotic induced drug 
delivery system (AIDD) is that the drug carrier cells may be genetically engineered 
to modify their properties. 
In vitro studies have been performed using S49 mouse lymphoma cells in which 
apoptosis is produced by exposure to dexamethasone. The cytotoxicity of RG-2 cells 
caused by temazolamide-loaded-S49 apoptotic cells was from 4 to 7 times higher 
than that of control temazolamide-loaded S49 cells.5 
7. Tumor Cells 
A novel strategy for drug delivery based on the use of cell systems is the drugloaded 
tumor cell system (DLTC), developed for drug delivery and targeting in 
Cells and Cell Ghosts as Drug Carriers 341 
lung metastasis.6'100 The tumor cells as drug carriers permit drug targeting to the 
blood-borne cancerous cells and the lungs as potential metastatic organs. In practice, 
there is affinity between the plasma membrane of malignant tumor cells and the 
metastatic addressins expressed by the endothelial cells of the targeted organ.6101 
In vivo studies have been carried out with DLTC based on Doxorubicin-loaded 
B16-F10 murine melanoma cells. Doxorubicin accumulation in the mouse lung was 
several times higher than that seen after administering free Doxorubicin.6 
8. Dendritic Cells 
Dendritic cells (DC) are antigen-presenting cells. They ingest antigen by phagocytosis, 
degrade it, and present fragments of the antigen at their surface. Dendritic cells 
have huge potential for immunization against a broad variety of diseases, because 
they travel throughout the body in search of pathogens indicative of infection or 
disease. They are very important for the induction of T cell responses, which result 
in cell-mediated immunity. 
Selective targeting of drugs incorporated in dendritic cells to T cells allows 
the response of these cells to be manipulated in vivo. It has been shown that when 
incorporated into dendritic cells, the drug O-galactosylceramide improves their 
anti-tumor activity.7 
9. Conclusions 
This chapter has focused on the use of cells and cell ghosts as delivery systems of 
drugs, enzymes or therapeutic genes. The use of carrier cells such as bacterial ghosts, 
erythrocyte ghosts and engineered stem cells, for drug delivery and targeting are 
reviewed among others. Their high biocompatibility, together with their capacity 
for selective delivery and targeting in cells and specific tissues mean that these types 
of carrier are promising alternatives to the use of nano- and microparticle systems, 
with applications in the fields of interest such as cancer therapy, cardiovascular 
therapy, AIDS, gene therapy, etc. As an alternative to the use of cell carriers, modified 
viruses can also be used as drug delivery systems, especially in the field of gene 
therapy. Despite their potential interest, clinical studies with these types of carrier 
are still very limited, although in the near future, increase in the use and therapeutic 
applications of cell delivery systems is expected. 
This chapter was supported in part by a project of the National Research and Development 
Plan (Project: SAF 2001-0740). 
342 Lanao & Sayalero 
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Cochleates as Nanoparticular Drug 
Leila Zarif 
1. Introduction 
In spite of the availability of many non-traditional novel dosage forms, oral route 
remains the most attractive way for administration of therapeutical materials. 
However, many therapeutic agents, especially the increasing number of biological 
molecules cannot be taken up by intestine due to their intrinsic impermeability 
to tissue membranes and the enzymatic degradation through the wall of the GI tract. 
Carrier systems that facilitate intestine uptake of these molecules are of major interests 
in the drug delivery arena. Moreover, drug delivery systems that provide a route 
of administration that does not involve injection can improve patient compliance 
and expand the market for existing, injectable, drugs. The factors which are important 
for the oral efficiency of a vehicle system have been repeatedly summarized in 
the literature.1,2 Small particle size, appropriate surface properties, mucoadhesive 
and targeting moieties, stability, as well as dose are the major factors imparting the 
efficiency of oral uptake. 
Producing formulations of poorly soluble drugs with high bioavailability is 
an even higher challenge. Known technologies are nanocrystals and nanoparticles 
which use the approach of enhancing the bioavailability by a decrease in particle 
size, resulting in an increase of surface area and subsequently a faster dissolution. 
Other technologies such as solid dispersions, polymeric micelles and selfemulsifying 
systems were developed to increase the drug solubility. 
350 Zarif 
Many lipid-based systems were developed to enhance oral bioavailability3,4 
Examples are lipid-based emulsions & microemulsions5-7; Solid lipid nanoparticles 
(SLN), a high melting point lipids enclosed in a surfactant layer8,9 adequate 
to enhance the oral bioavailability of poorly absorbed drugs; Lipid nanocapsules 
(LNC) for oral, injectable use10 and improved bioavailability11; Lipid nanospheres 
prepared from egg lecithin and soybean, described for their low toxicity12 and 
higher efficacy, compared with other delivery systems when incorporating amphotericin 
B,13 due to their smaller particle size and lower uptake by reticuloendothelial 
system.14,15 Recently, solid lipid microparticles, prepared by the solvent-in-wateremulsion-
diffusion technique, were described for the encapsulation and oral delivery 
of insulin.16 
In particular, lipid-based cochleate delivery system appears to provide answers 
to oral delivery challenges by (1) formulating different kind of molecules, especially 
hydrophobic ones17,18 and (2) protecting the sensitive and biologically active 
molecules from harsh environmental conditions. 
In this review, we will focus on cochleates nanoparticular drug carrier and will 
present the main features and the state of the art of this delivery technology. 
2. Cochleates Nanoparticles in Oral Delivery 
2.1. Cochleate structure 
Cochleates were first described by Dimitrious Papahadjopoulos and his co-workers 
in 1975 as precipitates formed by the interaction of negatively charged phosphatidylserine 
and calcium.19-21 He named these cylindrical structures "cochleate", 
meaning shell in the Greek language because of their rolled-up form, and explained 
the mechanism of cochleates formation by the fusion of negatively charged vesicles 
induced by the calcium cation22 (Fig. 1). 
These cigar-like structures have gained interest as antigen delivery system for 
vaccine applications.23 More recently cochleates were studied as tools to deliver 
small molecule drugs.17,18,24 A cochleate lipid formulation of amphotericin B has 
been developed as an oral composition to treat systemic fungal infections.24-26 
Other medical and non-medical applications are also under investigation.27 
2.2. Cochleate preparation 
2.2 A. Which phospholipid and which cation to use? 
Cochleates are a phospholipid-ion precipitates. Does that mean that cochleate is a 
structure obtained from precipitation of any phospholipid with any ion as presented 
in some litterature?,28 i.e. a complex of negatively charged phospholipid with any 
cation or a complex made from a positively charged lipid with any anion? 
Cochleates as Nanoparticular Drug Carriers 3 51 
f) Ca 0 fusion ( \ I 1 ///EPTA. 
A B C D E F 
Fig. 1. Cochleate cylindrical structure and mechanism of formation (adapted from Refs. 19 
and 69 with permission). 
Papahadjopoulos has given in 1975 this appellation to a rolled phospholipid 
structure. So far, to our knowledge no physico-chemical evidence on the obtention 
of such cigar-like structure from positively charged phospholipid with an anion 
had been described; on the contrary, extensive litterature is available on obtaining 
these cigar-like structure when negatively charged phospholipid such as phosphosphatidylserine 
(PS) had been precipitated with a cation such as calcium.17'18'20-22'29-30 
Other negatively charged phospholipids, such as phosphatidic acid (PA) or phosphatidyl 
glycerol derivatives, have been studied as well. Mixture of negatively 
charged phospholipids with other lipids can lead to cochleate formation. In this 
case, the cochleate formation depends on the negatively charged lipid/other lipid 
ratio and depends on the nature of the negatively charged lipid in the mixed lipid 
system. For example, PA derivatives form cochleate domains after the addition of 
calcium cation. However, when mixed with the corresponding diacylphosphatidylcholine 
(PC) and diacylphosphatidylethanolamine (PE), it was found that up to 
20 mole% of PC or PE can be introduced into the cochleate phase of PA(Ca2+), 
above which a distinct PC rich or PE-rich phase appears.31 
Other phospholipid derivatives such as galactosphingolipid hydroxy fatty acid 
cerebroside were reported to form cochleate cylinders by thermal mechanical treatment 
of glycol suspensions.32 However, the addition of conjugated lipid, such as 
352 Zarif 
poly(ethylene glycol)-lipid conjugates to PS vesicles, inhibited the calcium-induced 
In general, an additional desired feature of an oral drug delivery system is 
that the excipient permitting this transport to be classified is generally regarded 
as safe (GRAS). Soy phosphatidylserine fits this criteria. Furthermore, Soy PS 
has been used as a nutrient supplement since early 1980s. Clinical trials showed 
that PS may play a role in supporting mental functions in aging brains such as 
enhancing the memory, improving learning ability,34-41 reducing the stress42'43 and 
Cochleates can be made from purified soy phosphatidylserine, which represents 
an affordable source of raw material.45 A study comparing the purified soy 
phosphatidylserine (PSPS) to non-purified soy PS (NPSPS) has been disclosed in 
this patent, showing that PS should be present in an amount of at least 75% of the 
total lipid in order to allow the formation of cochleates. The other 25% phospholipids 
present can be selected either from the anionic group such as phosphatidic 
acid, phosphatidylglycerol, phosphatidyl inositol or phosphatidylcholine. PSPS 
cochleates can be loaded with different bioactive materials such as nutritional supplement, 
vitamins, antiviral, antifungal, small peptides. Proof of principle of the 
use of purified soy PS has been achieved using a polyene antifungal agent, amphotericin 
B. The preparation method for amphotericin B cochleates can be either via 
High pH-trapping or film method18 or by hydrogel method;29 the latter leading to 
nanocochleates formation. 
The nature of the cation is an important factor in cochleate formation. In the 
precipitation process, divalent cations are preferred to monovalent cations. Monovalent 
cations such as Na+ were described to prevent the cochleate formation.46 
Increases concentration of Na+ ions was shown to interfere with the destabilization 
effect of Ca2+. A critical Ca/PS ratio is necessary for the destabilization effect 
of divalent cations and the formation of cochleate phases.46 
The formation of cochleate is easier from small unilamellar vesicles (SUV). 
However, multilamellar vesicles (MLV) can also lead to cochleate formation. In this 
case, the first mechanism is a destabilization of the outer bilayer of PS by Ca2+ 
which causes its collapse, leading to a higher access of Ca2+ to inner PS bilayers 
and so forth. 
2.2.2. Which molecules can be entrapped in cochleates nanoparticles 
Due to the intrinsic nature of the lipid-contained cochleates, these nanoparticles 
can encapsulate a variety of molecules of all shapes and sizes. Preference is given, 
however, to hydrophobic molecules, for which a need to enhance chemical stability 
or bioavailability is desired [Fig. 2(a)]. Amphiphatic molecules which can easily 
Cochleates as Nanoparticular Drug Carriers 353 
Fig. 2. Type of molecules which can be encapsulated into lipid based cochleate (adapted 
from Ref. 18 with permission). 
insert in the membrane bilayers [Fig. 2(b)], negatively charged moiety [Fig. 2(c)] 
or positively charged moiety [Fig. 2(d)] could be encapsulated in the cochleate 
nanoparticle structure. 
The nature of the drug influence the percentage of encapsulation. Hydrophobic 
drag shows a quantitative encapsulation, whereas less was seen for amphiphatic 
molecules. For instance, doxorubicin which presents hydrophobic regions is a 
water-soluble drug, has a partition between the bilayers and the external aqueous 
phase [Fig. 2(b)]. As calcium induces dehydration of the interbilayer domains, 
the amount of water in this region is low,47 therefore, small hydrophilic molecules 
will not be suitable for cochleate system. 
2.2.3. Multiple ways of preparing cochleates 
Several processes were developed to obtain cochleates with a nanosize range, with 
the objective to allow oral delivery.24,29'48-59 Particle size is process dependent. When 
a small nanosized particle is desired, the "hydrogel method" can be used, based 
on the use of an aqueous-aqueous emulsion system.29 Briefly, this method consists 
of 2 steps: The preparation of small size liposomes either by high pH method18'25 
or by film method,18 then the liposomes are mixed with a high viscosity polymer 
354 Zarif 
such as dextran. The dextran/liposome phase is then injected into a second, nonmiscible, 
polymer (i.e. PEG). The calcium was then added and diffused slowly from 
one phase to another, resulting in the formation of nanocochleates. The final step 
is the washing of the gel. These nanosized cochleates showed potential in the oral 
delivery of drugs.18,29,48,59 
Electron microscopy and X-Ray crystallography of the nanoparticles show a 
unique multilayered structure consisting of continuous, solid lipid bilayer sheets, 
rolled up in a spiral with no internal aqueous space and the localization of AmB in 
the lipid bilayer.25 
Other preparation techniques are known, e.g. the trapping method, useful for 
the encapsulation of hydrophilic and hydrophobic molecules,17'18 which consist 
in the preparation of the liposomal suspension containing the drug either in the 
aqueous space of liposome (when hydrophilic) or intercalated in between the bilayers 
(when hydrophobic). A step of addition of calcium follows, and an aggregate 
of cochleates are formed. The cochleates made by the Trapping method present 
higher aggregation compared with other methods. This has been demonstrated 
using Electron microscopy after Freeze-fracture.25 
Another method was developed for hydrophobic drugs,61 known as "the solvent 
drip method" which consists of preparing a liposomal suspension separately 
based on soy PS and a hydrophobic or amphipathic cargo moiety solution. Solvent 
for hydrophobic drug can be selected from DMSO, DMF. The solution is then added 
to liposomal suspension. Since the solvent is miscible in water, a decrease of the 
solubility of the cargo moiety is observed, which associates at least in part with the 
lipid-hydrophobic liposomal bilayers. The cochleates are then obtained by addition 
of calcium and the excess solvent is being washed. 
Usually, the cochleate formation can be characterized by optical microscopy 
when they are present in needle form in the micrometer size range. In this case, 
direct observation using a higher magnification can be used.25 When nanocochleate 
are obtained, optical microscope can be used as an indirect method to assess the 
formation of cochleate, i.e. observation of the liposome formation after chelation 
of the calcium present, by addition of EDTA (ethylene diamine tetraacetate) to 
nanocochleate. A more sophisticated method is the electron microscopy after freezefracture18'
25 which allows the observation of the tighted packed bilayers. Recently, 
other methods were described using Laurdan (6-dodecanoyl-2-dimethylamino 
naphtalene) to monitor the cochleate phase formation.62 In this case, the lipid vesicles 
are labeled with Laurdan and the addition of calcium to the laurdan labeled 
vesicles resulted in a shift in the emission peak maximum of Laurdan. Due to 
dipolar relaxation, excitation and emission, generalized polarization (GPgx and 
GPEm) indicates the transition from a LC to a rigid and dehydrated cochleate 
Cochleates as Nanoparticular Drug Carriers 355 
2.3. Cochleates as oral delivery system for antifungal agent, 
amphotericin B 
Among the drug of choice using nanocochleate delivery system, amphotericin B 
(AmB) presented all aspects of a good candidate. Amphotericin B is a hydrophobic 
drug with poor oral bioavailability. This drug had been used for decades in 
injectable form to treat systemic fungal infections of Candida, cryptococcus and 
aspergillosis species.63-65 
Lipid formulations of Amphotericin B such as liposomes, lipid complexes, lipid 
emulsions and colloidal dispersions, were developed with the aim to achieve a 
higher therapeutic index.26-66 These formulations indeed showed enhanced therapeutic 
index, even though none of these formulations showed ability to deliver 
AmB orally. Cocheate technology seems to offer the advantage over other delivery 
systems in providing the possibility for the oral delivery of AmB. Oral administration 
of amphotericin B cochleates (CAMB) to healthy mice achieved potentially 
therapeutic concentrations in key target tissues.51 
Preclinical studies demonstrate a promising activity of CAMB in murine 
models of clinically relevant invasive fungal infections such as disseminated 
candidiasis,25'48,67 disseminated aspergillosis17,18-58'59 and central nervous system 
2.3.1. In candidiasis animal model 
In Candida albicans infected murine animal model, AmB cochleates showed potential 
either after intraperitoneal (i.p.) or oral (p.o.) administration.17,18,48,49,54,55,57,60,66-68 
After i.p. administration CAMB provided protection against C. albicans at doses 
as low as 0.1 mg/kg/day, kidney tissues burden showed that CAMB was more 
potent than Fungizone at 1 mg/kg/day and was equivalent to AmBisome at 
10 mg/kg/day18,25,60 (Fig. 3). CAMB was also effective after oral administration. 
Complete eradication of C. albicans from the lungs was noticed after p.o. administration 
at 2.5 mg/kg/day. These results were comparable to i.p. Fungizone at 
2.0 mg/kg/day.48,54-56 
2.3.2. In aspergillosis animal model 
Oral administration of CAMB was shown to be protective in a dosedependent 
manner against systemic infection of Aspergillus fumigatus in animals 
immunosusppressed with cyclophosphamide.58,59 In this mouse model, intragastric 
administration of CAMB at 40 mg/kg/day for 15 days resulted in 80% survival, 
while Fungizone at 4 mg/kg/day (i.p.) resulted in 20% survival; higher doses of 
Fungizone were lethal to animals. 
356 Zarif 
~-~ :-> Uo 
106 - 
105 - 
104 - 
1 0 '  
1 U 1 1 . 1 1 1 1 1  > 
Control 0.1 1.0 10.0 0.1 1.0 10.0 0.1 1.0 
AmB Dose Concentration (mg/kg) 
Fig. 3. Kidneys tissue burden of infected mice treated with either CAMB (), Fungizone 
() or AmBisome (), compared with controls (T) (from Ref. 18 with permission) 
 1 1 I J ffl -1 u 
1 i* 
* JL 
' l 5 ! 
U JJ M  M. , , 
control DAMB Smg/kg lOmg/kg 20mg/kg 30mg/kg 40mg/kg 
a a c; 
s? ~   < 
Concentration of Drug 
Fig. 4. Tissue burden for mice infected in a model of invasive aspergillosis after oral administration 
of CAMB (from Ref. 58 with permission). 
The tissue fungal burden for target organs, kidneys, liver and lungs, demonstrated 
the benefic effect of CAMB (Fig. 4 ). CAMB showed a pronounced dosedependent 
reduction in the fungal burden in all organs. The near eradication 
of Aspergillus was observed above a concentration of 20mg/kg/day. CAMB at 
30 mg/kg (PO) was as effective as CAMB at 20 mg/kg (PO) in reducing fungal 
tissue burden.58 
... . t 
Cochleates as Nanoparticular Drug Carriers 357 
2.3.3. In cryptococcal meningitis animal model 
Oral amphotericin B cochleates were effective in a murine cryptococcal meningitis 
model with an 80% survival after 17 days, obtained after oral treatment 
with CAMB (lOmg/kg) to mice having intracerebral infection with cryptococcus 
2.3.4. Toxicity of amphotericin B cochleates 
In vitro, Amphotericin B cochleates (CAMB) showed a low toxicity on red blood cells 
when compared with Fungizone (DAMB). CAMB showed no hemoglobin release 
and therefore no hemolysis of red blood cells when incubated at 500 ^g/ml. In 
contrast, DAMB was hemolytic at 10 /xg/ml due to the presence of the detergent, 
sodium desoxycholate.25 
In vivo, CAMB was non toxic to mice when administered orally at 
50mg/kg/day for 14 days. No nephrotoxicity was observed as demonstrated by 
the normal BUN level, and the histopathology of kidneys, lungs, liver, spleen and 
GI tract showed that animals dosed with CAMB were comparable to controls.18 
2.3.5. Pharmacokinetics of amphotericin B cochleates 
Oral pharmacokinetics?)^ 
Pharmacokinetic studies have shown that after oral administration of CAMB, AmB 
is distributed into the target tissues (e.g. brain, liver, lung, spleen and kidneys)18,50'52 
in healthy mice and AmB tissue level suggests a zero-order uptake process for all 
When CAMB was administered po to C57BL/6 mice at lOmg/kg (n = 5), 
and blood and tissues collected and AmB level measured by HPLC, blood 
shows a plateau-shaped profile with Tmax = 6h and Cmax = 0.05mg/ml. Noncompartmental 
(NCA) analysis showed blood AUC0-oo = 1.20/xg*h/ml, ti/2 = 
12.8 h, MRTo_oo = 21.1 h, Cl/F = 139.2ml/min/kg, Vz /F = 153.91 L/kg. AmB tissue 
exposure (AUCo-oo, .ig*h/g) evaluated using NCA was greater for lungs (23.11), 
followed by liver (16.91), spleen (15.40) kidneys (14.97) and heart (3.34). Tissue elution 
ti/2(h): kidneys 9.3, lungs 5.6, heart 5.3, liver 4.9 and spleen 4.3. For all tissues, 
Tmax = 12 h and Cmax ranged between 0.23/zg/ml for heart and 1.58/xg/ml for 
The delivery of AmB by cochleates after multiple oral doses (10) was assessed 
in the same mouse model and was compared with AmBisome. It was found 
that cochleate provides therapeutic levels in tissue and presents better delivery 
and transfer efficiency of AmB to the target tissue, as well as better tissue 
358 Zarif 
The ability of cochleate vehicles to deliver systemic AmB after single or multiple 
oral dosing suggest the potential of CAMB formulations to treat and prevent 
systemic fungal infections. 
AmB given intraveneously (IV) to mice showed a two-phase pharmacokinetic 
profile.69,70 Pharmacokinetic analysis in target tissues (liver, spleen, kidney and 
lungs) shows a multi-peak profile, large AUC and MRT. 
After IV administration of 0.625 mg/kg, AMB presented a two-phase blood 
concentration time course [Fig. 5(A)]. This profile is characterized by a very fast 
distribution phase and an elimination phase with t1/2 = 11.68 hrs. The AUCo-oo w a s 
1.006 A<,g*h/ml, CI = 10.36 ml/min/kg, MRT0_oo = 15.41 hrs and Vs s = 9.587 L/kg. 
This pharmacokinetic profile indicates that CAMB is removed fast from blood. 
In addition, the large Vss also indicates a large distribution into the tissues. The 
results obtained in target tissues showed this extensive distribution and penetration 
[Fig. 5(B)]. 
Calculation of pharmacokinetic parameters showed that the main target tissues 
have a large AMB exposure reflected in the AUC and CMAX values (Table 1), as well 
as the tissue to blood AUC ratio. 
The large AMB exposure in liver and spleen suggests involvement of the 
mononuclear phagocyte system (MPS) in the removal of CAMB. Cochleates are 
particulates that can be quickly cleared from the circulation by the macrophages of 
the reticular endothelial system (RES) related to the liver and the spleen. In addition, 
"physical retention" seems to play a role in the kinetic profile of the lungs due 
to its capillary nature. 
Time (hrs) "* 
0 10 20 30 40 50 
Time (hours) 
Fig. 5. (A) AMB profile in blood after a single dose (B) IV PK profile of AMB in target 
tissues, (from Ref. 69, with permission). 
Cochleates as Nanoparticular Drug Carriers 359 
Table 1 Pharmacokinetics parameters for CAMB in different 
target organs after IV administration to C57BL/6 
mice (n = 5 per time point) (From Ref. 69, with 
T max 
*~ max 
t l/2>-z
This phenomenon and the mobility of the macrophages seem to cause certain 
redistribution of cochleates that gives a multi-peak and plateau shape profiles in 
liver and spleen. Finally, AMB was also detected in bile and intestine contents, 
suggesting that bile excretion may be an additional elimination route. 
2.4. Other potential applications for cochleates 
2.4.1. Cochieate for the delivery of antibiotics 
As cochieate has shown a high affinity to be engulfed by macrophages [Fig. 6(A)] 
probably due to a dual mechanism, the cochieate essential particulate feature71 and 
possibly a PS receptor mediated internalization of the cochieate into macrophage.72 
Fig. 6. Uptake of amphotericin B cochleates by J774 macrophages as seen by (A) fluorescence 
microscopy, (B) confocal microscopy (from Ref. 17, with permission). 
360 Zarif 
This particulate system would have potential for the delivery of antibacterial 
agents such as aminoglycosides and vancomycin.17 Illustration is given by the 
encapsulation of clofazimine, an anti-TB drug, and tobramycin, an aminoglycoside 
antibiotic used in treating bacterial infections, both given intraveneously thus far. 
The cochleate system may possibly offer a new oral way of delivery. 
2.4.2. Delivery of clofazimine 
Clofazimine cochleates were prepared by the Trapping method.18 Clofazimine 
is a known hydrophobic anti-TB drug, the efficacy of Clofazimine cochleate 
was assessed by measuring the IC50 in Vero Cells and in bone marrow derived 
macrophage (BM-M).73 Clofazimine cochleates exhibit a greater decrease in toxicity 
versus free clofazimine and had a higher efficacy in killing intracellular 
M. Tuberculosis than free clofazimine:2 Log reduction (CE99) was achieved at 
20.9 /xg/ml for cochleates, while free clofazimine was toxic at this concentration. 
This shows that encapsulation of clofazimine in cochleates potentiates the antimicrobial 
efficacy of the drug, i.e. when higher concentration of drug can be used 
because of less toxicity, bactericidal levels of the drug could be attained. 
2.4.3. Delivery of tobramycin 
A recent research work has been published on the possible use of nanocochleates as 
an oral delivery system for Tobramycin.74 Tobramycin is a well known aminoglycoside 
antibiotic used in treating bacterial infections, and is usually administered by 
intravenous (i.v.) infusion, intramuscular (i.m.) injection, or inhalation. This aminogycoside 
drug is known for its side effects such as mineral depletion (i.e. calcium, 
magnesium, potassium) after i.v. administration.75,76 
In this work, the author described that tobramycin which is positively charged 
at low pH, will be encapsulated in the inter-bilayer space of cochleates. The fusion 
of unilamellar liposomes is no longer induced by a metal cation such as Ca2+, 
but by the organic molecule to be encapsulated. The cochleate cylinders formation 
has been described by Papahadjoupolos as resulting partly from the intrinsic 
properties of the calcium cation. Indeed, phosphatidylserine shows considerable 
selectivity for calcium due to the propensity of calcium to lose part of its hydration 
shell, and to displace water upon complex formation.19'77 In the cochleate solid 
crystalline structures formation, calcium plays a crucial role in bringing bilayers 
together closely through partial dehydration of the membrane surface and the crosslinking 
of opposing molecules of phosphatidylserine. In our opinion, in this recent 
work where formation of cochleate is claimed with no calcium present, additional 
Cochleates as Nanoparticular Drug Carriers 361 
relevant physico-chemical evidence on cochleate formation and the localization of 
the drug in the interbilayer space will be needed. 
2.4.4. Cochleate for the delivery of anti-inflammatory drugs 
As a result of the deep embedding of the molecules in the cochleates structures, 
drug molecules are hidden from the outside environment. This should have two 
beneficial effects: one is to hide and protect the molecule from the degradation due 
to environment; the other is to protect, the environment when needed, from the 
active molecule when such molecule presents side effects. 
This is the case of anti-inflammatory drugs, which associates cure to the disturbance 
of GI tract (stomach for instance). Cochleates were described to act beneficially 
in this area, reducing the stomach irritation when anti-inflammatory drugs 
such as aspirin is hidden in the cochleate structure, and administered to a carrageenan 
rat model for acute inflammation.27,61 
2.5. Othet uses of cochlea tes 
Cochleates were also described as vehicles for nutrients27 as an improved drug 
and contrast agent delivery system,28 as well as intermediate in the preparation of 
special liposomes such as Large Unilamellar Vesicles (LUV) and proteoliposomes. 
In fact, the discovery of the cochleate structures was a result of the desire to prepare 
LUV by Pr papahadjoupoulos,19'20 which were developed for the delivery of 
hydrophilic drugs. Proteoliposomes prepared from cochleates intermediates were 
described for vaccine applications in general,78 and more recently, when containing 
lipopolysaccharide as a novel adjuvant.79 
3. Conclusion 
Cochleates lipid-based nanocarrier appears to have potential for the oral delivery 
of bioactive molecules. Future work should be directed towards more fundamental 
science, as many research aspects of the cochleate drug carrier system are still hardly 
known (e.g. localization of the drug in lipid bilayers, impact of multivalent cations 
on the cochleate formation, mechanism of action of cochleate after oral uptake). In 
addition, the development of friendly analytical assays to monitor the drug localization 
and loading percentage in cochleates will be desired. This nano drug carrier 
is currently under development by Biodelivery Sciences International.27 Having 
the first drug-cochleate in the market place represents a big challenge. For instance, 
when oral amphotericin B cochleates are ultimately available for patients, thus will 
provide a new opening in the treatment of systemic fungal infections. 
362 Zarif 
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Aerosols as Drug Carriers 
N. Renee Labiris, Andrew P. Bosco 
and My ma B. Dolovich 
1. Introduction 
As the end organ for the treatment of local diseases or as the route of administration 
for systemic therapies, the lung is a very attractive target for drug delivery (Table 1). 
The lung provides direct access to the site of disease for the treatment of respiratory 
illness, without the inefficiencies and unwanted effects of systemic drug delivery. 
In addition, it provides an enormous surface area and a relatively low enzymatic 
environment for the absorption of drugs to treat systemic diseases (Table 1). 
Inhaled medications have been available for many years for the treatment of 
lung diseases. Inhalational delivery has been widely accepted as being the optimal 
route of administration of first line therapy for asthmatic and chronic obstructive 
pulmonary diseases. Drug formulation plays an important role in producing an 
effective inhalable medication. In addition to being pharmacologically active, it is 
important that a drug be efficiently delivered into the lungs, to the appropriate site 
of action and remain in the lungs until the desired pharmacological effect occurs. 
A drug designed to treat a systemic disease, such as insulin for diabetes, must be 
deposited in the lung periphery to ensure maximum systemic bioavailability. For 
gene therapy, anti cancer or anti infective treatment, cellular uptake and prolonged 
residence in the lungs of the drug may be required to obtain the optimal therapeutic 
effect. Thus, a formulation that is retained in the lungs for the desired length of time 
and avoids the clearance mechanisms of the lung may be necessary. 
The human lung contains airways and approximately 300 million alveoli with 
a surface area of 140 m2, equivalent to that of a tennis court.1 As a major port of 
368 Labiris, Bosco & Dolovich 
Table 1 
Advantages of pulmonary delivery of drugs to treat respiratory and systemic 
Treatment of respiratory diseases Treatment of systemic diseases 
Deliver high drug concentrations directly 
to the disease site 
Minimizes risk of systemic side effects 
Rapid clinical response 
Bypass the barriers to therapeutic 
efficacy, such as poor gastrointestinal 
absorption and first-pass metabolism in 
the liver 
Achieve a similar or superior therapeutic 
effect at a fraction of the systemic dose. 
For example, oral salbutamol 2-4 mg is 
therapeutically equivalent to 100-200 /xg 
A non-invasive Needle-free delivery 
Suitable for a wide range of substances 
from small molecules to very large 
Enormous absorptive surface area 
(140 m2) and a highly permeable 
membrane (0.2 to 0.7 /xm thickness) in 
the alveolar region. 
Large molecules with very low 
absorption rates can be absorbed in 
significant quantities; the slow 
mucociliary clearance in the lung 
periphery results in prolonged residency 
in the lung. 
A less harsh, low enzymatic 
Avoids first-pass metabolism. 
Reproducible absorption kinetics. 
Pulmonary delivery is independent 
of dietary complications, extracellular 
enzymes and inter-patient metabolic 
differences that affect gastrointestinal 
entry, the lung has evolved to prevent the invasion of unwanted airborne particles 
from entering into the body. Airway geometry, humidity, mucociliary clearance 
and alveolar macrophages play a vital role in maintaining the sterility of the lung, 
and consequently, they can be barriers to the therapeutic effectiveness of inhaled 
The size of the drug particle can play an important role in avoiding the physiological 
barriers of the lung and targeting to the appropriate lung region (Fig. 1). 
Nanoparticles are solid colloidal particles ranging in size from 10 to 1000 nm.2 
Studies have demonstrated that they are taken up by macrophages, cancer cells, 
and epithelial cells.3-6 Their small size ensures the particles containing the active 
pharmacological ingredient will reach the alveolar regions. However, the use of an 
aerosol delivery system that generates nano-sized particles for inhalation, places 
these particles at risk of being exhaled, leaving very few drug particles to be 
deposited in the periphery of the lung. Residence time is not long enough for the 
particles to be deposited by sedimentation or diffusion.7 
Aerosols as Drug Carriers 369 
05 ID 2.0 5.0 
Fig. 1. Relationship between particle size and lung deposition. 
2. Pulmonary Drug Delivery Devices 
The origin of inhaled therapies can be traced back 4000 years ago to 
India, where people smoked the leaves of the Atropa belladonna plant to 
suppress cough. In the 19th and early 20th centuries, asthmatics smoked 
asthma cigarettes that contained stramonium powder mixed with tobacco 
to treat the symptoms of their disease. Modern inhalation devices can be 
divided into three different categories (Fig. 2), the refinement categories 
(Fig. 2), the refinement of the nebulizer and the of compact portable 
devices, the pressurized metered dose inhaler (pMDI), and the dry powder 
inhaler (DPI). The advantages and disadvantages of each are summarized in 
Table 2. 
2.1. Nebulizers 
Nebulizers have been used for many years to treat asthma and other respiratory 
diseases. There are 2 basic types of nebulizers, jet and ultrasonic nebulizers. The 
jet nebulizer functions by the Bernoulli principle by which compressed gas (air 
or oxygen) passes through a narrow orifice, creating an area of low pressure at 
the outlet of the adjacent liquid feed tube. This results in the drug solution being 
drawn up from the fluid reservoir and shatter into droplets in the gas stream. The 
ultrasonic nebulizer uses a piezoelectric crystal, vibrating at a high frequency (usually 
1 to 3 MHz), to generate a fountain of liquid in the nebulizer chamber; the 
higher the frequency, the smaller the droplets produced. Nebulizers can aerosolize 
3 70 Labiris, Bosco & Dolovich 
Glass Nebulizer 
(Late 19* century) 
Hand Bulb Nebulizer 
Metered Dose Inhalers (MDI) 
(1956, CFC prcmellant) 
Metered Dose 
Liquid Inhalers 
Dry Powder Inhaler 
Passive Active 
Fig. 2. Evolution of pulmonary delivery devices. 
most drug solutions and provide large doses with very little patient coordination 
or skill. However, treatments using these nebulizers can be time consuming and 
inefficient, with large amounts of drug wastage e.g. 50% loss with continuously 
operated nebulizers.8 Most of the prescribed drug never reaches the lung with nebulization. 
The majority of the drug is either retained within the nebulizer (referred 
to as residual or dead volume) or released into the environment during expiration. 
On average, only 10% of the dose placed in a continuous output jet nebulizer is 
actually deposited in the lungs.8 Advances in technology have led to the development 
of novel nebulizers that reduce drug wastage and improve delivery efficiency. 
Breath-enhanced jet nebulizers such as the Pari LC Star, (PARI, Germany) increase 
aerosol output by directing auxiliary air, entrained during inspiration, through 
the nebulizer, causing more of the generated aerosol to be swept out of the nebulizer 
and available for inhalation. Drug wastage during exhalation is reduced to 
the amount of aerosol produced by the jet airflow rate that exceeds the storage 
volume of the nebulizer. Adaptive aerosol delivery (Halolite, Medic-Aid, Bognor 
Regis, UK) monitors a patient's breathing pattern in the first 3 breaths and then targets 
the aerosol delivery into the first 50% of each inhalation. This ensures that the 
aerosol is delivered to the patient during inspiration only, thereby eliminating drug 
loss during expiration that occurs with continuous output nebulizers.9 A number 
of metered dose liquid inhalers, including AERx (Aradigm, Hayward, CA), Aero- 
Dose (AeroGen, Sunnyvale, CA) and Respimat (Boehringer Ingelheim, Ingelheim 
Rhein, Germany), have been developed to produce a fine aerosol in the respirable 
Aerosols as Drug Carriers 371 
Table 2 Advantages and disadvantages of inhalation devices. 
Inhalation device Advantages Disadvantages 
Nebulizers (jet, ultrasonic) no specific inhalation 
technique or coordination 
aerosolizes most drug 
delivers large doses 
suitable for infants and 
people too sick or 
physically unable to use 
other devices 
time consuming 
contents easily 
relatively expensive 
poor delivery efficiency 
drug wastage 
wide performance 
variation between 
models and operating 
pressurized Metered Dose 
Inhalers (pMDI) 
Dry Powder Inhalers (DPI) 
multi-dose (-200 
sealed environment (no 
degradation of drug) 
reproducible dosing 
breath actuated 
easy to use 
no hand-mouth 
coordination required 
inhalation technique 
and patient coordination 
high oral deposition 
maximum dose of 5 mg 
limited range of drugs 
respirable dose 
dependent on IFR* 
humidity may cause 
powders to aggregate 
and capsules to soften 
dose lost if patient 
inadvertently exhales 
into the DPI 
most DPIs contain 
*IFR = Inspiratory Flow Rate 
range by forcing the drug solution through an array of nozzles, using vibrating 
mesh or electronic micropump platforms with 30 to 75% of the emitted dose being 
deposited in the lungs.10,11 
2.2. Metered-dose inhalers 
The pressurized metered-dose inhaler (pMDI) was a revolutionary invention that 
overcame the problems of the hand-bulb nebulizer, and it is the most widely 
used aerosol delivery device today. The pMDI emits a drug aerosol driven by 
372 Labiris, Bosco & Dolovich 
propellants, such as chlorofluorocarbons (CFC) and more recently, hydrofluoroalkanes 
(HFAs) through a nozzle at high velocity (>30m/sec). pMDIs deliver only a 
small fraction of the drug dose to the lung. Typically, only 10 to 20% of the emitted 
dose is deposited in the lung.12 The high velocity and large particle size of 
the spray causes approximately 50% to 80% of the drug aerosol to impact in the 
oropharygeal region.13 Hand-mouth discoordination is another obstacle in the optimal 
use of the pMDI. Crompton and colleagues14 found 51% of patients experienced 
problems coordinating the actuation of the device with inhalation, 24% of 
patients halted inspiration upon firing the aerosol into the mouth, and 12% inspired 
through the nose instead of the mouth when the aerosol was actuated into the 
The delivery efficiency of a pMDI depends on a patient's breathing pattern, 
inspiratory flow rate and hand-mouth coordination. The studies by Bennett15 and 
Dolovich16 demonstrated that for any particle size between 1 to 5 /tm mass median 
aerodynamic diameter (MMAD), deposition was more dependent on inspiratory 
flow rate than any other variable. Fast inhalations (>60 L/min) result in a reduced 
peripheral deposition because the aerosol is more readily deposited by inertial 
impaction in the conducting airway and oropharyngeal regions. When aerosols 
are inhaled slowly, deposition by gravitational sedimentation in peripheral lung 
regions are enhanced.17 Peripheral deposition has also been shown to increase with 
an increase in tidal volume and a decrease in respiratory frequency. As the inhaled 
volume is increased, aerosols are able to penetrate more distally into the lungs.18 
A period of breath holding on completion of inhalation enhances deposition of 
particles in the periphery, thus preventing the particles from being exhaled during 
the expiratory phase. Thus, the optimal conditions for inhaling pMDI aerosols are 
from a starting volume equivalent to the functional residual capacity, the actuation 
of the device at the start of inhalation, inspiratory flow rate of <60 L/min, followed 
by a 10 second breath-hold at the end of inspiration.17,19 
Spacer tubes, valved holding chambers and mouthpiece extensions have been 
developed to eliminate coordination requirements and reduce the amount of drug 
deposited in the oropharynx, by decreasing the particle size distribution and slowing 
the aerosol's velocity. Spacer geometry and materials of manufacture influence 
the quality and quantity of aerosol available. The aerosols from a pMDI and the 
holding chamber are finer than that with the pMDI alone, with an approximate 
25% decrease in the mass median aerodynamic diameter (MMAD), compared with 
the original aerosol.20,21 This finer aerosol is more uniformly distributed in the normal 
lung, with increased delivery to the peripheral airway. However, in patients 
with airway obstructions, the addition of a holding chamber to the pMDI may not 
change the distribution of the aerosol.22 
Aerosols as Drug Carriers 373 
2.3. Dry powder inhalers 
Dry powder inhalers (DPIs) were designed to eliminate the coordination difficulties 
associated with the pMDI. There are a wide range of DPI devices on the market from 
single-dose devices loaded by the patient (e.g. Aerolizer from Novartis, Rotahaler 
from GSK, Ware UK) to multi unit dose devices provided in a blister pack (e.g. 
Diskhaler, GSK, Ware UK), multiple unit doses sealed in blisters on a strip that 
moves through the inhaler (e.g. Diskus, GSK, Ware UK) or reservoir-type (bulk 
powder) systems (e.g. Turbuhaler, AstraZeneca, Lund Sweden). 
Lung deposition varies among the different DPIs. Approximately 12% to 40% 
of the emitted dose is delivered to the lungs with 20 to 25% of the drug being 
retained within the device.10,23,24 Poor drug deposition with DPIs can be attributed 
to inefficient deaggregation of the fine drug particles from coarser carrier lactose 
particles or drug pellets. Slow inspiratory flow rate, high humidity and rapid, large 
changes in temperature are known to affect drug deaggregation and hence the efficiency 
of pulmonary drug delivery with DPIs.25,26 With most DPIs, drug delivery 
to the lungs is augmented by fast inhalation. Borgstrom and colleagues27 demonstrated 
that increasing inspiratory flow from 35L/min to 60L/min through the 
Turbuhaler7, increased the total lung dose of terbutaline from 14.8% of nominal 
dose to 27.7%. This is in contrast to the MDI which requires slow inhalation and 
breath holding to enhance lung deposition of the drug. Each DPI has a different 
air flow resistance that governs the required inspiratory effort.28,29 The higher the 
resistance of the device, the more difficult it is to generate an inspiratory flow great 
enough to achieve the maximum dose from the inhaler.30-32 However, deposition 
in the lung tends to increase when using high resistance inhalers.32-36 
Active DPIs are being investigated to reduce the importance of a patient's inspiratory 
effort. By adding either a battery driven propeller that aids in the dispersion 
of the powder (Spiros, Elan Pharmaceuticals, San Diego, CA), or using compressed 
air to aerosolize the powder and converting it into a standing cloud in a holding 
chamber, the generation of a respirable aerosol becomes independent of a patient's 
inspiratory effort (Inhance Pulmonary Delivery System, Nektar Therapeutic, San 
Carlos, CA). 
3. Aerosol Particle Size 
Aerosol particle size is one of the most important variables in defining the dose 
deposited and the distribution of drug aerosol in the lung (Fig. 3). Fine aerosols 
are distributed on peripheral airways, but deposit less drug per unit surface area 
than larger particle aerosols which deposit more drug per unit surface area, but on 
3 74 Labiris, Bosco & Dolovich 
(a) JOO*. 
% 50 _ 
0.1 i )0 100 
100 | - / > 
% 50 
/ MMAD = 2.25 ftm 
0.1 i 10 
Fig. 3. Frequency (a) and cumulative (b) distribution curves for Beclovent MDI used with 
an Aerochamber, in terms of number of particles and volume (mass) of particles vs. particle 
aerodynamic diameter. The volume distribution curves are displaced to the right of the 
number distribution curves. The smaller number of large particles within the aerosol carry the 
greater mass of the drug; this is reflected in the larger, second peak of the volume distribution 
curve, which corresponds to the smaller second peak of the number distribution curve. 
MMAD is read from the cumulative distribution curve at the 50% point and if the distribution 
is log-normal, the GSD can be calculated as the ration of the diameter at the 84.1% point to 
the MMAD. Particle distribution was measured using the Anderson Cascade Impactor.105 
the larger, more central airways.37 Most therapeutic aerosols are nearly always heterodisperse, 
consisting of a wide range of particle sizes. These aerosols are described 
by the log-normal distribution, with the log of the particle diameters plotted against 
particle number, surface area or volume (mass) on a linear or probability scale and 
expressed as absolute values or cumulative %. Since delivered dose is very important 
when studying medical aerosols, particle number may be misleading as smaller 
particles contain less drug than larger ones. Particle size is defined from this distribution 
by several parameters. Mass median diameter of an aerosol refers to the 
Aerosols as Drug Carriers 375 
particle diameter that has 50% of the aerosol mass residing above and 50% of its 
mass below it. The aerodynamic diameter relates the particle to the diameter of a 
sphere of unit density that has the same settling velocity as the particle of interest, 
regardless of its shape or density. MMAD is read from the cumulative distribution 
curve at the 50% point (Fig. 3). Geometric standard deviation (GSD) is a measure of 
the variability of the particle diameters within the aerosol, and is calculated from 
the ratio of the particle diameter at the 84.1% point on the cumulative distribution 
curve to the MMAD. For a log-normal distribution, the GSD is the same for the 
number, surface area or mass distributions. A GSD of 1 indicates a monodispersed 
aerosol, while a GSD of > 1.2 indicates a heterodispersed aerosol. 
Particles can be deposited by inertial impaction, gravitational sedimentation 
or diffusion (Brownian motion), depending on their size. While deposition occurs 
throughout the airways, inertial impaction usually occurs in the first 10 generations 
of the lung, where air velocity is high and airflow is turbulent.38 Most particles above 
10 /xm are deposited in the oropharyngeal region with a large amount impacting 
on the larynx, particularly when the drug is inhaled from devices requiring a high 
inspiratory flow rate (DPIs) or when the drug is dispensed from a device at a high 
forward velocity (MDIs).39,40 The large particles are subsequently swallowed and 
contributed minimally, if at all, to the therapeutic response. In the tracheobronchial 
region, inertial impaction also plays a significant role in the deposition of particles, 
particularly at bends and airway bifurcations. Deposition by gravitational sedimentation 
predominates in the last 5 to 6 generation of airways (smaller bronchi 
and bronchioles), where air velocity is low.38 In the alveolar region, air velocity 
is negligible and thus the contribution to deposition by inertial impaction is also 
negligible. Particles in this region have a longer residence time and are deposited 
by both sedimentation and diffusion. Particles not deposited during inhalation are 
exhaled. Deposition due to sedimentation affects particles down to 0.5 ^tm in diameter, 
whereas below 0.5 /xm, the main mechanism for deposition is by diffusion. 
Targeting the aerosol to conducting or peripheral airways can be accomplished 
by altering the particle size of the aerosol. It is difficult to predict the actual site 
of deposition, since airway calibre and anatomy differ among people. However, 
in general, aerosols with a MMAD of 5 to 10 /xm are mainly deposited in the large 
conducting airways and the oropharyngeal region.41 Particles 1 to 5 /xm in diameter 
are deposited in the small airways and alveoli with greater than 50% of the 3 /tm 
diameter particles being deposited in the alveolar region. In the case of pulmonary 
drug delivery for systemic absorption, aerosols with a small particle size would 
be required to ensure peripheral penetration of the drug.42 Particles <3 /xm have 
approximately 80% chance of reaching the lower airways, with 50 to 60% being 
deposited in the alveoli.43'44 Nanoparticles <100nm are deposited mainly in the 
alveolar region. 
376 Labiris, Bosco & Dolovich 
4. Targeting Drug Delivery in the Lung 
The therapeutic effect of aerosolized therapies is dependent on the dose deposited 
and its distribution within the lung. If a drug aerosol is delivered at a suboptimal 
dose or to a part of the lung, devoid of the targeted disease or receptors, the 
effectiveness of therapy may be compromised. For example, the receptors for the fc 
agonist, salbutamol and the muscarine (M3) agonist, ipratropium bromide, are not 
uniformly distributed throughout the lung. Autoradiographic studies have shown 
P2 adrenergic receptors are present in high density in the airway epithelium from 
the large bronchi to the terminal bronchioles. Airway smooth muscle has a lower 
/S-receptor density, greater in the bronchioles than bronchi.45 However, greater than 
90% of all /3 receptors are located in the alveolar wall, a region where no smooth 
muscle exists and whose functional significance is unknown. Another autoradiographic 
study has shown a high density of M3 receptors in submucosal glands 
and airway ganglia, and a moderate density in smooth muscles throughout the 
airways, nerves in intrapulmonary bronchi and in alveolar walls.46 The location of 
these receptors in the lung suggests that ipratropium bromide needs to be delivered 
to the conducting airways, while salbutamol requires a more peripheral delivery 
to the medium and small airways to produce a therapeutic effect. 
Since particle size affects the lung deposition of an aerosol, it can also influence 
the clinical effectiveness of a drug. Rees et al. reported the varying clinical effect 
of 250 /xg of aerosolized terbutaline from a pMDI, given in three different particle 
sizes of <5 /xm, 5 to 10 /xm, and 10 to 15 /xm.47 In asthmatics, the greatest increase in 
forced expiratory volume in one second (FEVi) was found with the smallest particle 
size (<5/xm), suggesting that the smaller particle aerosol was considerably more 
effective than larger particle size aerosols in producing bronchodilation, since it has 
the best penetration and retention in the lungs in the presence of airway narrowing. 
Using three monodisperse salbutamol aerosols (MMAD of 1.5 /xm, 2.8 /xm, 5 Aim), 
Zanen and colleagues demonstrated in patients with mild to moderate asthma 
that the 2.8 /xm particle size aerosol produced a superior bronchodilation, compared 
with the other two aerosols.48 In patients with severe airflow obstruction 
(FEVi < 40%), Zanen et al. demonstrated that the optimal particle size for /J2 agonist 
or anticholinergic aerosols is approximately 3 /xm.49 They examined the effect 
on lung function of equal doses of three different sizes of monodisperse aerosols, 
1.5 /xm, 2.8 /xm and 5 /xm, of salbutamol and ipratropium bromide. Their findings 
suggest that small particles penetrate more deeply into the lung and more effectively 
dilate the small airways than larger particles, which are filtered out in the 
upper airways. The 1.5 /xm aerosol induced significantly less bronchodilation than 
the 2.8 /xm aerosol, suggesting that this fine aerosol may be deposited too peripherally 
to be effective, since smooth muscle is not present in the alveolar region. 
Aerosols as Drug Carriers 377 
The optimal site of deposition in the respiratory tract for aerosolized antibiotics 
depends on the infection being treated. Pneumonias represent a mixture of purulent 
tracheobronchitis and alveolar infection. Successful therapy would theoretically 
require the antibiotic to be evenly distributed throughout the lungs. However, those 
confined to the alveolar region would most likely benefit from a greater peripheral 
deposition. Pneumocystis carinii pneumonia, the most common life-threatening 
infection among patients infected with HIV, is found predominately within the 
alveolar spaces, with relapses occurring in the apical region of the lung after treatment 
with inhaled pentamidine given as a 1 fim MMAD aerosol.50 The mechanism 
suggested for this atypical relapse is the poorer apical deposition of the aerosol. 
Regional changes in intrapleural pressure result in the lower lung regions receiving 
relatively more of the inspired volume than the upper lung, when sitting in an 
upright position or standing. This influence on deposition has been shown to occur 
in an experimental lung model, analyzing sites of aerosol deposition in a normal 
lung. The experiment showed a 2:1 ratio in the overall deposition for a 4 /xm aerodynamic 
diameter aerosol between the lower and upper lobes when in the upright 
Chronic lung infection with Pseudomonas aeruginosa, in patients with cystic 
fibrosis or non-CF bronchiectasis, resides in the airway lumen with limited invasion 
of the lung parenchyma.52'53 Infection starts in the smaller airways, the bronchioles, 
and moves into the larger airways. The optimal site of deposition for inhaled 
antimicrobial therapy would, therefore, be a uniform distribution on the conducting 
airways. Mucus plugs in the bronchi and bronchioles may prevent deposition 
of even small particle aerosols in regions distal to the airway obstruction, possibly 
the regions of highest infection, and thereby limiting the therapeutic effectiveness 
of the aerosolized antibiotic.54-56 
Until recently, aerosol drug delivery has been limited to topical therapy for 
the lung and nose. The major contributing factor to this restriction was the inefficiencies 
of available inhalation devices that deposit only 10% to 15% of the emitted 
dose in the lungs. While appropriate lung doses of steroids and bronchodilators can 
be achieved with these devices, for systemic therapies, large amounts of the drug 
are necessary to achieve therapeutic drug levels systemically. Recent advances in 
aerosol and formulation technologies have led to the development of delivery systems 
that are more efficient and that which produce small particle aerosols, allowing 
higher drug doses to be deposited in the alveolar region of the lungs, where they 
are available for systemic absorption. 
Most macromolecules cannot be administered orally because proteins are 
digested before they are absorbed into the bloodstream. In addition, their large 
size prevents them from naturally passing through the skin or nasal membrane; 
therefore, they cannot be administered intranasally or transdermally without the 
378 Labiris, Bosco & Dolovich 
use of penetration enhancers. Thus, the easiest route of administration for proteins 
has been through intravenous or intramuscular/subcutaneous injection. It has been 
known for many years that proteins can be absorbed from the lung as demonstrated 
with insulin in 1925.57 Macromolecules < 40 kiloDaltons (kDa) (<5-6nm 
in diameter) appear rapidly in the blood following inhalation into the airways. 
Insulin which has a molecular weight (mw) of 5.7 kDa and a diameter of 2.2 nm 
peaks in the blood 15 to 60 min after inhalation.58-62 Macromolecules >40 kDa (>5- 
6 nm in diameter) are slowly absorbed over many hours; inhaled albumin (68 kDa) 
and alphai-antitrypsin (45-51 kDa) have a Tmax of 20hrs and between 12 to 48hrs 
The lung is the only organ through which the entire cardiac output passes. 
Before the inhaled drug can be absorbed into the blood from the lung periphery, 
it has several barriers to overcome such as lung surfactant, surface lining fluid, 
epithelium, interstitium and basement membrane, and the endothelium. Drug 
absorption in the lung periphery is regulated by a thin alveolar-vascular permeable 
barrier. An enormous alveolar surface area with epithelium, consisting of a 
thin single cellular layer (0.2 to 0.7 /xm thickness), promotes efficient gas exchange 
through passive transport, but also provides a mechanism for efficient drug delivery 
into the bloodstream.64 Although the mechanism of absorption is unknown, 
it has been hypothesized that macromolecules either pass through the cells via 
absorptive transcytosis (adsorptive or receptor mediated), paracellular transport 
between bijunctions or trijunctions or through large transitory pores in the epithelium 
caused by cell injury or apoptosis.65 Thus, the high bioavailability of macromolecules 
deposited in the lung (10 to 200 times greater than nasal and gastrointestinal 
values) may be due to its enormous surface area, very thin diffusion layer, 
slow surface clearance and anti-protease defense system. 
5. Clearance of Particles from the Lung 
Like all major points of contact with the external environment, the lung has evolved 
to prevent the invasion of unwanted airborne particles from entering into the body. 
Airway geometry, humidity and clearance mechanisms contribute to this filtration 
process. The challenge in developing therapeutic aerosols is to produce an aerosol 
that eludes the lung's various lines of defense. 
5.1. Airway geometry and humidity 
Progressive branching and narrowing of the airways encourages impaction of particles. 
The larger the particle size, the greater the velocity of incoming air, while 
the greater the bend angle of bifurcations and the smaller the airway radius, the 
Aerosols as Drug Carriers 379 
greater the probability of deposition by impaction.66 Drug particles are known to 
be hygroscopic and grow in size in high humidity environments, such as the lung 
which has a relative humidity of approximately 99.5%. The addition and removal of 
water can significantly affect the particle size and thus deposition of a hygroscopic 
aerosol.67 A hygroscopic aerosol that is delivered at relatively low temperature and 
humidity into one of high humidity and temperature would be expected to increase 
in size when inhaled into the lung. The rate of growth is a function of the initial 
diameter of the particle, with the potential for the diameter of fine particles less 
than 1 /xm to increase 5-fold, compared with 2 to 3-fold for particles greater than 
2 /xm.68 The increase in particle size above the initial size should affect the amount 
of drug deposited, and particularly, the distribution of the aerosolized drug within 
the lung. Ferron and colleagues have predicted that for initial sizes between 0.7 /xm 
and 10 /xm, total deposition of hygroscopic aerosols increases by a factor of 2.69 
For particles with an initial size of 1 /xm, Xu and Yu were able to predict changes 
in the distribution pattern due to particle growth.70 The calculations showed a 
shift from deposition due to sedimentation to primarily impaction on more central 
5.2. Lung clearance mechanisms 
Once deposited in the lungs, inhaled drugs are either cleared from the lungs, 
absorbed into the circulatory or lymphatic systems, or metabolized. Drug particles 
deposited in the conducting airways are primarily removed through mucociliary 
clearance, and to a lesser extent, are absorbed through the airway epithelium into 
the blood or lymphatic system. Ciliated epithelium extends from the trachea to 
the terminal bronchioles. The airway epithelial goblet cells and submucosal glands 
secrete mucus forming a two-layer mucus blanket over the ciliated epithelium: a 
low-viscosity periciliary or sol layer covered by a high-viscosity gel layer. Insoluble 
particles are trapped in the gel layer and moved towards the pharynx (and 
ultimately to the gastrointestinal tract) by the upward movement of mucus generated 
by the metachronous beating of cilia. In the normal lung, the rate of mucus 
movement varies with the airway region and is determined by the number of 
ciliated cells and their beat frequency. Movement is faster in the trachea than in 
the small airways, and is affected by factors influencing ciliary functioning and 
the quantity and quality of the mucus.40'71 For normal mucociliary clearance to 
occur, airway epithelial cells must be intact, ciliary structure and activity normal, 
the depth and chemical composition of the sol layer optimal, and the rheology of 
the mucus within the physiological range. Mucociliary clearance is impaired in 
lung diseases such as immotile cilia syndrome, bronchiectasis, cystic fibrosis and 
asthma.72 In immotile cilia syndrome and bronchiectasis, the ciliary function can be 
380 Labiris, Bosco & Dolovich 
either impaired or nonexistent. In cystic fibrosis, the ciliary structure and function 
are normal, however, the copious amounts of thick, tenacious mucus present in 
the airways impairs their ability to clear the mucus effectively73 In these diseases, 
clearance of aerosolized drugs deposited in the conducting airways is generally 
decreased and secretions are cleared from the lung by cough.74-76 
In addition to mucociliary clearance, soluble particles can also be removed 
by absorptive mechanisms in the conducting airways.77 Lipophilic molecules pass 
easily through the airway epithelium via passive transport. Hydrophilic molecules 
cross via extracellular pathways such as tight junctions or by active transport via 
endocytosis and exocytosis.78 From the submucosal region, particles are absorbed 
either into systemic circulation, bronchial circulation or lymphatic systems. 
Drugs deposited in the alveolar region may be phagocytosed and cleared 
by alveolar macrophages or absorbed into the pulmonary circulation. Alveolar 
macrophages are the predominant phagocytic cell for the lung defense against 
inhaled microorganisms, particles and other toxic agents. There are approximately 
5 to 7 alveolar macrophages per alveolus in the lungs of healthy, non-smokers.79 
Macrophages phagocytose insoluble particles deposited in the alveolar region are 
either cleared by the lymphatic system or moved into the ciliated airways along currents 
in alveolar fluid and then cleared via the mucociliary escalator.65 This process 
can take weeks or months to complete.7 As discussed above, soluble drug particles 
deposited in the alveolar region can be absorbed into the systemic circulation. The 
pulmonary epithelium appears to be more resistant to soluble particle transport 
than to the endothelium or the interstitium.42 
The lung-blood barrier may behave as a molecular sieve, allowing the passage 
of small solutes but restricting the passage of macromolecules. Conhaim and colleagues 
proposed that the lung barrier was best fitted to a three pore size model, 
including a small number (2%) of large-sized pores (400 nm pore radius), 30% of 
medium-sized pores (40 nm radius) and 68% of small-sized pores (1.3 nm).80 
The rate of protein absorption from the alveoli is size dependent. Effros and 
Mason demonstrated an inverse relationship between alveolar permeability and 
molecular weight.42 In rats, after intratracheal instillation of DDAVP (1-desamino- 
8-D-arginine vasopressin) (raw = 1.1 kDa), peak serum DDAVP levels occurred 
at 1 hr compared with 16 to 24hrs after the intratracheal instillation of albumin 
(mw = 67 kDa).43 However, some proteins are cleared from the lung more rapidly 
than expected for their size. After intratracheal instillation or aerosolization of 
human growth hormone (mw = 22 kDa), peak serum levels were observed between 
0.5 to 4 hrs, indicating a rapid, saturable clearance from the lung that is suggestive 
of receptor-mediated endocytosis.65 Vasoactive intestinal polypeptide (VIP) 
is believed to be completely degraded during the passage across the pulmonary 
epithelium and into the bloodstream.81 
Aerosols as Drug Carriers 381 
Nanoparticles can pass rapidly into the systemic circulation. The distribution 
of radioactivity, after the inhalation of a 99mTechnetium (Tc)-labeled ultrafine carbon 
particles (5 to 10 nm), was detected in the blood one min post-inhalation and 
peaked between 10 and 20 min. This blood radioactivity level was sustained up to 
60 min. 8% of the initial lung radioactivity was measured in the liver 5 min postadministration 
and remained stable over time. The rapidity of the appearance of 
radioactivity systemically makes the translocation from the lung unlikely due to 
phagocytosis, by macrophages or endocytosis by epithelial and endothelial cells, 
but by passive diffusion.82 
6. Nanoparticle Formulations for Inhalation 
Delivery of nano-sized aerosols to the lung may result in very little drug being 
deposited in the lung. The majority of particles <500nm inhaled will not have 
enough residence time in the lung to deposit, and therefore will be exhaled (Fig. 1). 
However, if the nanoparticles were delivered in larger carrier particles, they could 
be sufficiently deposited in the lung. The carrier particle would dissolve after contact 
with the lung surface fluid, releasing the nanoparticle at the target tissue or cells. 
Sham and colleagues demonstrated that nanoparticles (173 to 242 nm) could 
be delivered into the lung in larger respirable lactose carrier particles produced 
by spray-drying.83 The dry powder containing the nanoparticles had a MMAD of 
3.0 /xm. pMDI formulations are typically micronized drugs in the 2 to 3 /xm range 
suspended in a hydrofluoroalkane (HFA) propellant. Solution pMDI such as QVAR 
produce smaller drug particles on propellant evaporation, resulting in better deposition 
and distribution than a micronized formulation.84 However, for insoluble 
drug particles in the propellant, the efficiency of pMDI is limited. A study by Dickinson 
et al. proposed the use of nanoparticles suspended in propellant as a method 
of increasing the delivery efficiency of insoluble drugs in pMDIs.85 They produced 
hydrophilic nanoparticles using a reverse phase microemulsion technique that captures 
nanoparticles by snap freezing, followed by freeze-drying. The nanoparticles 
of pure drug (salbutamol) and the drug in a non-polymer matrix (lecithin-based), 
with and without lactose, were dispersed in HFA-227 and in aerosol performance 
assessed by cascade impaction. The size of the salbutamol nanoparticles ranged 
from 34 to 216 nm. Dispersion of the nanoparticles in a HFA-227:hexane (95:5 v/v) 
blend resulted in a homogeneous fine suspension that showed no signs of sedimentation 
or creaming over several months. Rapid release of salbutamol from 
the nanoparticle was observed (approximately 4 min) as expected from the large 
surface area of the particles and the high water solubility of the drug. A high 
fine particle fraction (ex-device, % < 5.8 /xm) of 58.3% to 65.5% and a low MMAD 
382 Labiris, Bosco & Dolovich 
(1.2 to 1.5 /u.m) were observed with the nanoparticle formulations. This data suggests 
that a high fraction of the nanoparticles would be distributed in the alveolar 
region of the lung and represents the best aerosol that can be produced using 
a pMDI. 
Budesonide is a potent corticosteroid used as an inhaled anti inflammatory 
agent to treat asthma. It is available as a dry powder inhaler and as a suspension for 
inhalation with a nebulizer. A new formulation for nebulization has been developed 
that contains nanocrystals of budesonide that give the suspension solution-like 
qualities.86 The particles are 75 to 300 nm in diameter, compared with 4400 nm 
for the marketed budesonide suspension (Pulmicort Respules, AstraZeneca). In 
a randomized crossover study, 16 healthy volunteers were given the nanocrystal 
budesonide formulation (0.5 mg and 1.0 mg doses), Pulmicort respules and placebo 
via nebulization using a Pari LC jet nebulizer. Nebulization times were shorter 
for the nanocrystal formulation, compared with Pulmicort respules (~7.1 min vs. 
8.7 min). Similar AUCs were observed with the formulations, suggesting similar 
pulmonary absorption. However, a higher Cmax (1212pg/mL vs. 662pg/mL) and 
shorter Tmax (8.4 min vs. 14.4 min) for nanocrystal budesonide compared with the 
same dose of Pulmicort, suggests a more rapid drug delivery or absorption with 
the nanocrystal formulation. 
6.1. Diagnostic imaging 
Radiolabeled nanoparticles have been used for many years in pulmonary ventilation 
studies.87 Ultrafine 99mTc labeled carbon particles (Technegas) is a relatively 
new advance in ventilation scintigraphy.88 Technegas (Vita Medical Ltd., Sydney 
Australia) consists of nanoparticles of carbon with a diameter of approximately 
5 nm, that behaves more like a 0.2 /tm particle.89 Technegas is generated by the 
electrostatic heating of a graphite crucible to 2500C in which a saline solution 
of 99mTc-pertechnetate had been placed and dried. The aerosol is dispersed in a 
lead-lined chamber in an atmosphere of 100% argon gas that is then inhaled by 
the patient. It is deposited in alveoli by inhalation and distributes similarly as the 
inert gas radioisotopes. Once they are inhaled, the particles adhere to the alveolar 
structures without appreciable movement for at least 40 min.88 
Pulmonary delivery of nanoparticles is also being investigated for lymphoscintigraphy 
to assess the spread of or the staging of lung cancer. Lung cancer 
usually exhibits metastasis proliferation, spreading through the lymphatic system 
and the blood circulation. Lymphatic drainage is responsible for the alveolar 
clearance of the deposited particulates and drugs up to a certain particle diameter 
(500 nm).90 Thus, radiolabeled nanoparticles could be used to visualize the lymph 
nodes to determine the presence of tumors. 
Aerosols as Drug Carriers 383 
The lymphatic uptake of solid lipid nanoparticles has also been studied as 
an imaging method to stage lung cancer. The lipid nanoparticles were radiolabeled 
with the lipophilic tracer, D,L-hexamethylpropylene amine oxime (HMPAO), 
tagged with 99 m-Tc. The lipid nanoparticles were prepared by the melted homogenization 
method and had a mean diameter of 200 nm.90 The radiolabeled nanoparticles 
were aerosolized using an ultrasonic nebulizer and delivered to rats until 
200,000 cpm was achieved over the lung. After inhalation, the total activity in the 
lung was observed, followed by a fast clearance rate (ti/2 = lOmin) that decreases 
activity in the lung to 25% of the total dose. Asignificant uptake (16.7%) was detected 
in the regional lymph nodes during the first 45 to 60 min, suggesting that aerosol 
delivery to the lungs of solid lipid nanoparticles could be used as an effective colloidal 
carrier for lymphoscintigraphy. 
Drainage into the lymph nodes following the lung instillation of nanoparticles 
of insoluble iodinated CT x-ray contrast agents was studied in beagle dogs.91 
Nanoparticles of the contrast agent were prepared by microfluidization. A particle 
size of 150 to 200 nm was achieved. The nanoparticles were suspended in 2 different 
surfactant solutions. 1.5 mL of the suspension was instilled using a fiber optic bronchoscope 
at specific sites in the small airways and alveoli. The nanoparticles were 
transported from the lung to the draining lymph nodes, 6 to 9 days post instillation 
as visible on the CT radiographs. No adverse clinical signs were observed in the 
dogs. However, microscopic lung lesions were observed at the instillation sites for 
both formulations and vehicle. The lesions consisted of inflammatory infiltrates, 
mainly macrophages, in intra-alveolar, interstitial and perivascular locations. A 
few small sites had fibrosis and granulomatous nodules with the destruction of the 
lung parenchyma. The presence of foamy macrophages was observed in the lymph 
nodes. The microscopic findings suggest that instillation of these nanoparticles of 
contrast agent may be harmful to the lung. The authors suggested that administering 
the nanoparticles as an aerosol, rather than by instillation, would prevent high 
concentrations in focal areas believed to be responsible for these lesions. 
6.2. Vaccine delivery 
Mucosal vaccine administration is an attractive method of inducing an immune 
response, since many pathogens invade the body through mucosal surfaces in the 
nose, lung and gut. As it is the first contact point, the mucosa has developed barriers 
to protect the body. The mucosa associated lymphoid tissue (MALT) is one of these 
barriers. It contributes 80% of the immunocytes and secretes more immunoglobulins 
than any other organs in the body.92 Antigens are delivered locally in the respiratory 
tract to nasal-associated and bronchus-associated lymphoid tissues (NALT 
and BALT, respectively) and a mucosal immunity is induced. Using nanoparticles, 
384 Labiris, Bosco & Dolovich 
systemic immunity may also be induced. Several studies have investigated the use 
of nanoparticles as carriers for the nasal delivery of vaccines. Using tetanus toxoid as 
a model antigen, Vila and colleagues have studied the use of chitosan nanoparticles 
as well as polyethyleneglycol and polylactic acid (PEG-PLA) nanoparticles as nasal 
vaccine carriers.93,94 They compared PEG-PLA nanoparticles with PLA alone.94 
Tetanus toxoid was entrapped in the hydrophobic PLA core and protected from 
interacting with enzymes such as lysozymes, by a hydrophilic PEG coating. Upon 
incubation with lysozymes in vitro, PLA particles aggregate and do not reach the 
epithelium, whereas PEG-PLA nanoparticles remain stable and size unmodified. 
The nanoparticles were produced using a double emulsion technique. PEG-PLA 
tetanus toxoid nanoparticles had a similar diameter to the PLA particles (196 nm 
vs. 188 nm), but had a lower loading efficiency of 33.4% compared with 48.1 % with 
PLA. The IgG antibody response induced by PEG-PLA was superior at weeks 2 to 
24, after intranasal instillation of 30 fig of tetanus toxoid (10 fil per nostril) on days 1, 
8 and 15 in male BALB/c mice. In a similar study, the same group compared radiolabeled 
PEG-PLA, PEG-PLA with gelatin stabilizer to radiolabeled PLA encapsulated 
tetanus toxoid. They reported that 1 hr after intranasal administration, PEG-PLA 
nanoparticles produced a radioactivity level 10-fold higher in the blood than PLA 
which remained constant for 24 hrs. The radioactivity detected in the lymph nodes, 
lungs, liver and spleen was 3 to 6 fold higher for PEG-PLA than PLA nanoparticles 
24 hrs post instillation. The results of this work suggest that the PEG-PLA 
nanoparticles are partially taken up by the M cells of the NALT, as well as being 
transported to the submucosa and drained into the lymphatic system and blood 
stream.95 Recent work by the same group has investigated the potential use of chitosan 
nanoparticles for nasal administration of vaccines.93 Chitosan is a hydrophilic 
natural polysaccharide that is biodegradable and has mucoadhesive properties. The 
nanoparticles are formed spontaneously by adding the counter anion sodium TPP 
into the chitosan solution, without the use of energy sources or organic solvents 
required for the production of PEG-PLA nanoparticles. Again, using tetanus toxoid 
as the model antigen, the investigators studied the effect of chitosan dose (200 fig 
and 70 fig) and molecular weight (23,38 or 70 kDa) on the efficacy of the nanoparticles. 
The nanoparticles produced were 300 to 350 nm and had a positive surface 
charge (+40 mV). The loading efficiency of tetanus toxoid was 50 to 60%, irrespective 
of the molecular weight of chitosan. In vitro, the formulations exhibited a rapid 
release over the initial 2 hrs followed by a slow release for 16 days, with the greater 
initial release at lower molecular weights of chitosan. 30 or 10 fig of antigen (associated 
with 200 and 70 fig of chitosan) was given intranasally to BALB/c mice on 
days 1, 8 and 15. The IgG levels induced by the nanoparticles were significantly 
higher than those elicited by free tetanus toxoid. The response lasted for the 24 
weeks studied with the IgG titres increasing over time. Anti-tetanus IgA titers were 
detected in the saliva, bronchoalveolar and intestinal lavage fluids 24 weeks post 
Aerosols as Drug Carriers 385 
administration. The results were independent of the administered dose and were 
significantly higher for the nanoparticle than the free tetanus toxoid. 
Jung and colleagues evaluated tetanus toxoid-loaded polymer nanoparticles as 
potential nasal vaccine carriers in mice.96 The nanoparticles were produced with 
various diameters (100 nm, 500 nm) using a novel polyester, sulfobutylated poly 
(vinyl alcohol)-graft-poly(lactide-co-glycolide), SB(43)-PVAL-g-PLGA. The surface 
charge was 43 to 59 mV. Mice were immunized with tetanus toxoid nanoparticles 
or free toxoid in solution at weeks 1, 2 and 3, either by oral, intranasal or 
intraperitoneal administration. Four weeks after the first intranasal immunization, 
IgG and IgA titers were significantly higher than baseline. Oral immunization with 
the nanoparticles produced a weak IgG antibody response. Only 10% of the oral 
dose was administered to the nose (2.89 vs. 28.9 /u,g), however, intranasal immunization 
appeared to be more effective in inducting an immune response. Particle size 
had an effect on the titer levels. Particles > 1 /i.m did not induce an immune response, 
but no difference was observed between the 500 nm and 100 nm nanoparticles which 
both induced significantly levels of IgG and IgA. 
These studies suggest that nasal delivery of vaccines using biodegradable 
nanoparticles are a promising method of inducing mucosal and systemic immunity. 
6.3. Anti Tuberculosis therapy 
Intracellular bacterial infections caused by pathogens such as Mycobacterium tuberculosis 
are difficult to eradicate because they are generally inaccessible to free antibiotics. 
By loading antibiotics into nanoparticles, it is expected that delivery to the 
infected cells would improve since nanoparticles have been shown to localize preferentially 
in organs with high phagocytic activity and in circulating macrophages 
as well.97 The encapsulation of antibiotics has several advantages: (1) It modifies 
their pharmacokinetic characteristics by prolonging the antibiotics half-life and 
increasing the area under the concentration time curve (AUC), while decreasing its 
apparent volume of distribution. (2) It improves the targeting of the drug to the 
phagocytic cells. (3) It reduces toxicity of the antibiotics, such as the hepatotoxicity 
of anti tuberculosis drugs and the nephrotoxicity of aminoglycosides. Antibiotics 
encapsulated in nanoparticles have been shown to be superior at treating intracellular 
infections when administered intravenously. However, the pulmonary delivery 
of these nanoparticles have only been investigated recently. 
Although effective therapy for tuberculosis is available, treatment failure and 
drug resistance is typically the result of patient's noncompliance. To improve compliance, 
investigators have been studying ways to reduce the dosing frequency of 
the drugs. Poly (lactide-co-glycolide) (PLG) nanoparticles as an aerosolized sustained 
release formulation for anti tuberculosis drugs, isoniazid, rifampicin and 
pyrazinamide, has been investigated since pulmonary tuberculosis is the most 
386 Labiris, Bosco & Dolovich 
common form of the infection.98 The majority of the nanoparticles were 186 to 
290 nm in diameter. Drug encapsulation efficiency was 56.9% to 68%. Aerosolized 
nanoparticles had a MMAD of 1.88 //.m, with 96% of the particles in the respirable 
range (<6 /u,m). A single nebulization to guinea pigs resulted in sustained plasma 
drug concentrations for 6 to 8 days and in the lung for 11 days. The half-life and 
mean residence time of the drugs was significantly prolonged, compared with the 
oral free drugs. Nebulizing the nanoparticles every 10 days to guinea pigs infected 
with Mycobacterium tuberculosis resulted in no detectable bacilli in the lung after 5 
doses of treatment, compared with 46 daily doses of orally administered drug to 
achieve the equivalent efficacy. 
The use of lectin-based PLG nanoparticles as an aerosolized sustained release 
formulation of isoniazid, rifampicin and pyrazinamide has also been studied in 
guinea pigs." Mucoadhesive drug delivery systems such as chitosan have been 
previously investigated as a method of prolonging residence at a site of absorption. 
The main drawback of mucoadhesive systems is that its residence time is limited 
by the turnover time of the mucous gel layer, which is only a few hrs. Attaching the 
polymeric nanoparticles to cytoadhesive ligands such as lectins could prolong the 
duration of adhesion, thereby prolonging residence time. Lectins bind to epithelial 
surfaces via specific receptors. Wheat germ agglutinin (WGA) is the least immunogenic 
lectin and has known receptors on the alveolar epithelium as well as the 
intestinal wall. WGA lectin-PLG nanoparticles were prepared by a two-step carbodiimide 
procedure. Their size ranged from 350 to 400 nm with drug encapsulation 
efficiency between 54% and 66%. The nanoparticles were delivered via nebulization 
to guinea pigs. 88% of the aerosol was in the respirable range (<6/U,m) with 
a MMAD of 2.8/zm (GSD of 2.1). Three doses of nanoparticles were administered 
every 15 days for 45 days. The WGA-PLG nanoparticles resulted in a prolonged 
Tmax/ increased AUC and mean residence time after inhaled delivery. All three drugs 
were present in the lungs, liver and spleen at concentrations above the minimum 
inhibitory concentration 15 days post dosing, compared with orally-administered 
free drug. Chemotherapeutic studies in guinea pigs infected with Mycobacterium 
tuberculosis showed that 3 doses administered every 15 days for 45 days yielded 
undetectable mycobacterial colony forming units, which was only achievable w