Design and Development of New Nanocarriers

By Alexandru Mihai Grumezescu

Design and Development of New Nanocarriers focuses on the design and development of new nanocarriers used in pharmaceutical applications that have emerged in recent years. In particular, the pharmaceutical uses of microfluidic techniques, supramolecular design of nanocapsules, smart hydrogels, polymeric micelles, exosomes and metal nanoparticles are discussed in detail. Written by a diverse group of international researchers, this book is a valuable reference resource for those working in both biomaterials science and the pharmaceutical industry.

• Shows how nanomanufacturing techniques can help to create more effective, cheaper pharmaceutical products
• Explores how nanofabrication techniques developed in the lab have been translated to commercial applications in recent years
• Explains safety and regulatory aspects of the use of nanomanufacturing processes in the pharmaceutical industry

• Language: English
• Publisher: Elsevier Science
• Release date: Nov 30, 2017
• ISBN: 9780128136287

Book preview

Russia

Series Preface: Pharmaceutical Nanotechnology

Alina M. Holban, University of Bucharest, Bucharest, Romania

Due to its immense applicative potential, nanotechnology is considered the leading technology of the 21st century. The science and engineering of nanometer-sized materials is currently employed for the development of numerous scientific, industrial, ecological, and technological fields. Biology, medicine, chemistry, pharmacy, agriculture, food industry, and material science are the main fields which have benefited from the great technological progress developed in nanoscience.

In the pharmaceutical field, nanotechnology has revolutionized traditional drug-design concept and the art of drug delivery. The idea of a highly specific nanoscale drug for the targeted therapy of diseases is now considered a feasible treatment for severe health conditions.

Some scientists believe that the pharmaceutical domain has been reborn by the important contribution of nanotechnology. The field of pharmaceutical nanotechnology has the potential to offer innovative solutions for all diagnosis, therapy, and prophylaxis domains. Application of nanotechnology tools in pharmaceutical research and design is likely to result in moving the industry from a blockbuster drug model to personalized medicine. The current main focus of clinicians is to treat patients individually, not their general diagnosed diseases, which are usually difficult to diagnose or incorrectly diagnosed. There are compelling applications in the pharmaceutical industry where suitable nanotechnology tools can be successfully utilized. By designing and modifying drugs at nanoscale, pharmaceutical nanotechnology could be useful not only for the development of completely new therapeutic solutions, but also to add value to existing products. This possibility opens perspectives of success for pharmaceutical companies in existing markets, but also for new markets.

Scientists have manifested an impressive interest on the field of pharmaceutical nanotechnology research in recent years. However, we face today a true dilemma of data unavailability, due to the multitude of existing information which can be highly inaccurate and contradictory. This is because of the lack of an efficient model for sorting the plethora of nanotechnology tools and information that exists, and strategically correlate those with potential opportunities into different segments of pharmaceutical research and design.

This series is trying to cover the most relevant aspects regarding the great progress of nanotechnology in the pharmaceutical field and to highlight the currently emerging trend of pharmaceutical nanotechnology towards the personalized medicine concept.

The 10 volumes of this series are structured to wisely offer relevant information regarding basic concepts and also to reveal the newest approaches and perspectives in pharmaceutical nanotechnology.

Nanoscale Fabrication, Optimization, Scale-Up and Biological Aspects of Pharmaceutical Nanotechnology, introduces the readers into the amazing field of nanoscale design. Also, this volume facilitate understanding of the biological requirements of nanostructured pharmaceutical formulations for advanced drugs.

In Design and Development of New Nanocarriers, the most recent progress made on the field of nano-delivery is discussed. Modern nanostructured drug carriers employ innovative solutions for the detection and treatment of various diseases in a personalized and efficient manner.

Design of Nanostructures for Theranostics Applications, highlights the impressive impact of nanotechnology in the development of combined diagnosis and therapy concept: theranostics.

Design of Nanostructures for Antimicrobial, Antioxidant and Nutraceutical Applications, offers a dynamic solution for immune modulation, treatment of diseases by natural-based products and infection control, while employing nanostructured solutions to achieve top results.

Nanostructures for the Engineering of Cells, Tissues and Organs: From Design to Applications, is a highly investigated and debated field; tissue engineering, is dissected through this volume. Here is shown how nanotechnology has advanced research and applications in the manipulation and engineering of cells and tissues in vitro.

Organic Materials as Smart Nanocarriers for Drug Delivery, deals with the specific world of organic nanomaterials, revealing their wide applications, types, and advantages in drug delivery.

In the volume entitled: Inorganic Frameworks as Smart Nanomedicines, the main focus is to discuss the variety and properties of inorganic nanostructures for therapy and drug delivery in the context of improved personalized medicine.

Lipid Nanocarriers for Drug Targeting, deals with recently developed lipid nanostructures and the advances made in drug targeting.

Drug Targeting and Stimuli-Sensitive Drug Delivery Systems, dissects smart stimuli-responsive nanosystems employed to specifically detect various biochemical conditions and control the release of drugs.

Fullerens, Graphenes and Nanotubes: A Pharmaceutical Approach, reveals major findings made on widely applied drug-design nanosystems, namely fullerens, graphenes and nanotubes. The impact of these nanostructures in pharmaceutical research is highlighted.

All 10 volumes are nicely illustrated and chapters are organized into a logical manner to be accessible to a wide audience. The series is a valuable resource of new and comprehensive scientific proof on the intriguing and emerging field of pharmaceutical nanotechnology, which could be of a great use for scientists, engineers, pharmaceutical representatives, clinicians, and any non-specialist interested user.

Preface

Alexandru M. Grumezescu, University Politehnica of Bucharest, Bucharest, Romania

The aim of this book is to present the novel progress performed in recent years in the field of design and development of new nanocarriers used in pharmaceutical nanotechnology. Special attention is assigned to microfluidic techniques, supramolecular design of nanocapsules, smart hydrogels, polymeric micelles, polymerosomes, niosomes, exosomes, smart micelleplexes, therapeutic proteins, and drug-imprinted nanostructures, layer-by-layer platforms, and many others.

The book, entitled Design and Development of New Nanocarriers, contain 18 chapters, prepared by outstanding researchers from Portugal, Russia, Japan, Spain, United States, Turkey, India, Pakistan, Poland, Brazil, India, Iran, Italy, and Greece.

Chapter 1, Vesicle-based drug carriers: Liposomes, polymersomes, and niosomes, prepared by Nily Dan, gives an up to date overview about the characteristics of vesicle-based formulations and their biomedical applications, ranging from cancer therapies to delivery across the blood–brain barrier.

Chapter 2, Recent advances in micellar-like polyelectrolyte/protein complexes: Design and development of biopharmaceutical vehicles, prepared by Natassa Pippa et al., discusses in depth the challenges associated with the development of polyelectrolyte-protein complexes that are related to complexation process, their physicochemical properties, as well as protein stability, release kinetics, and the conditions under which the system is delivered to the human body.

Chapter 3, Calixarene-based micelles: Properties and applications, prepared by Grazia Maria Letizia Consoli et al. reports on the methods used for the physicochemical characterization of calixarene micelles, and the role played by structural conformation of the calix[n]arene skeleton, length of hydrophobic chains, and structure of polar head groups in the aggregation behavior. The micelles are subdivided into anionic, cationic, zwitterionic, non-ionic, reversed, stimulus-responsive, and bicomponent. Potential applications of micellar calixarenes in pharmaceutical nanotechnology are also described.

Chapter 4, Preparation of Janus nanoparticles and their application in drug delivery, prepared by Sepideh Khoee and Akram Nouri, presents recent progress of various methodologies for synthesis of Janus nanoparticles, such as masking, phase separation or pickering emulsions. Preparation of Janus nanoparticles with different morphologies, i.e., cylindrical, spherical or disk-like was also investigated. At the end, the authors discussed their possibilities in preparation of novel materials as drug delivery systems, due to their widespread applications in the future.

Chapter 5, Supramolecular design of hydrophobic and hydrophilic polymeric nanoparticles, by Leonardo M.B. Ferreira et al., reviews fundamental aspects of supramolecular interactions, focusing on their importance to the fabrication of hydrophobic and hydrophilic nanoparticles. Specific discussion explores the self-assembly process of polymer building blocks in polymeric nanoparticles, and the main tools and techniques used to monitor synthetic procedures and characterize physicochemical properties. The chapter ends with the major challenges that should be addressed to optimize self-assembly for rational design of polymer-based nanoparticles.

Chapter 6, Cationic polyelectrolyte–biopolymer complex hydrogel particles for drug delivery, prepared by Sabyasachi Maiti et al. highlights the factors influencing the formation of polyelectrolyte complexes and recent developments in the field of polysaccharide-based polyelectrolyte complex hydrogels.

Chapter 7, Smart micelleplexes: An overview of a promising and potential nanocarrier for alternative therapies, prepared by Mariana Magalhães et al. gives an up to date overview about smart micelleplexes as non-viral vectors to be used in new therapeutic strategies based on gene therapy, as well as, its advantages, structure and capability to perform an efficient and specific delivery to the target cells, tissue or organ. Additionally, it also referred to studies performed in the past few years using smart micelleplexes as nanocarriers.

Chapter 8, Polymeric micelles as a versatile tool in oral chemotherapy, prepared by Andreza Maria Ribeiro et al. discusses nanocarriers tailored to fit the physicochemical properties of anticancer agents and the therapeutic peculiarities of tumor management for improving the effectiveness/toxicity ratio of the current treatments.

Chapter 9, In Mixed micelles, Jan Sobczyński and Beata Chudzik-Rząd discuss and review the recent development and future perspectives in the field of mixed micellar drug delivery systems. Mixed micellar formulations showed improved efficacy in the fields of anticancer therapeutics and improved oral delivery of poorly water soluble agents. The versatility of this approach allows for the concomitant integration of different features into a single system; a task that is synthetically challenging to accomplish by using just one polymeric or copolymeric block.

Chapter 10, Amphiphilic block copolymers based micelles for drug delivery, prepared by Muhammad Imran et al. presents the recent progress related to the amphiphilic block copolymers-based micelles in terms of their formation, building blocks, characterization, factors affecting their properties, and their applications for targeted drug delivery.

Chapter 11, Synthesis and evolution of polymeric nanoparticles: Development of an improved gene delivery system, by R. Mankamna Kumari et al. focuses on polymeric nanoparticles that have been widely explored in the field of pharmaceutics and biotechnology since its inception. A number of pre-clinical reports regarding polymeric nanoparticles-assisted gene delivery provides clear evidence that polymeric nanoparticles can find their way into clinic and will have a significant impact on therapeutics.

Chapter 12, Therapeutic protein and drug imprinted nanostructures as controlled delivery tools, prepared by Handan Yavuz et al., reports recent progress in nanoparticulate drug delivery systems with an enhanced permeation that allow targeted delivery of intended therapeutic drug to the site of action with predetermined time and activity. The use of molecular imprinting technology to arrange the drug release properties from the polymer matrix in the context of the affinity of imprinted polymer for the drug molecule, providing high loading capacity and sustained release profiles are also discussed.

Chapter 13, Application of complex coacervates in controlled delivery, prepared by Merve Deniz Köse et al. presents the recent progress related to complex coacervation of natural polyelectrolytes and their applications in delivery systems. There are various applications of complex coacervates, especially in controlled delivery systems, and mostly in the form of either micro/nanocapsules/beads or stimuli-responsive membranes. Complex coacervation constitutes a promising route for the preparation of a wide variety of tools, such as capsules and membranes used in controlled delivery systems.

Chapter 14, Hydrogels: Biomedical uses, prepared by María-Dolores Veiga et al. gives an up to date overview about the hydrogels that have been widely exploited for biomedical uses due to their rheological and physical properties, their biocompatibility, and their ability to control the release of drugs. These nanocarriers have proven to be versatile tools for therapy, thanks to their various properties and the fact that they can be oriented to be stimuli-sensitive. These possibilities make them useful for drug delivery systems, as they can target a specific organ and release the drug in a sustained manner.

Chapter 15, Technologies that generate and modify virus-like particles for medical diagnostic and therapy purposes, prepared by Masaaki Kawano et al. described how virus-like particles (VLPs) derived from various viruses can be used for a wide range of human and animal diagnostics and therapies. The authors focused in particular on SV40 VP1-based VLPs and described in detail the technologies that can be used to modify them. These include the preparation of VLPs from insect cells, in vitro VLP assembly, encapsulation of bioactive materials, chemical or genetic modification of the VLP surface, and coating of artificial beads with capsid subunits. They also discussed how these techniques can be used for targeted cell delivery of encapsulated materials in vitro and in vivo, and for generating vaccines that induce cytotoxic T lymphocytes via various immunization routes, including the nasal mucosal route.

Chapter 16, Layer-by-layer coated drug-core nanoparticles as versatile delivery platforms, prepared by Ana Cláudia Santos et al. gives a comprehensive revision of recent layer-by-layer (LbL) deposition applications on drug nanocores as strategic nanocarriers for drug delivery. The different methods and coating materials used in the preparation of the LbL nanocapsules, as well as recent advances in releasing encapsulated drugs through LbL nanocapsules are exposed and discussed. In addition to these topics, the LbL pharmaceutical applications, their advantages, and challenges (with particular emphasis on the advances of recent years by in vitro and in vivo studies), as well as the toxicological issues of their use and future perspectives, are addressed.

Chapter 17, Effect of α-dextrin nanoparticles on the structure of iodine complexes with polypeptides and alkali metal halogenides, and on the mechanisms of their anti-human immunodeficiency virus and anticancer activity, prepared by G.A. Yuldasheva et al. presents an up to date overview about the study of the structure of the active centers of drugs containing molecular iodine, located inside the nanoparticles of α-dextrin, and coordinated by lithium halogenides and polypeptides. The probable mechanism of the drug’s cytotoxic activity is explained by the results of molecular modeling process and DFT calculations.

Chapter 18, Nanocarriers for the delivery of temozolomide in the treatment of glioblastoma: A review, prepared by Maria João Ramalho et al. provides a systematic review on the current progress in nanodelivery of temozolomide focusing on glioblastoma therapy.

www.grumezescu.com

Chapter 1

Vesicle-based drug carriers

Liposomes, polymersomes, and niosomes

Nily Dan,    Drexel University, Philadelphia, PA, United States

Abstract

Phospholipids, block copolymers and surfactants are surface-active molecules that assemble into bilayer structures in aqueous solutions, forming nanoscale vesicles: phospholipid-based liposomes, polymer-based polymersomes, or surfactant-based niosomes. These vesicles are highly versatile and can be used for the delivery of pharmaceutical agents—enabling targeted and controlled release while shielding the encapsulated drugs from environmental degradation agents and the immune system. The performance of vesicle-based drug carriers has been shown to depend on both their constituent chemistry (e.g., surface moieties) and their physical properties (size, shape). This review examines the characteristics of vesicle-based formulations and their application to biomedical applications ranging from cancer therapies to delivery across the blood–brain barrier (BBB).

Keywords

Drug delivery; encapsulation; vesicles; nano; micro

Chapter Outline

1.1 Introduction 1

1.2 Amphiphilic Bilayers 4

1.3 Liposomal Drug Carriers 7

1.4 Polymersome Drug Carriers 11

1.5 Niosome Drug Carriers 14

1.6 Biomedical Applications 15

1.7 Discussion 22

1.8 Conclusion 24

References 25

1.1 Introduction

The development of nanoparticle drug carriers is driven by the need to diversify administration and delivery methods, and shield sensitive pharmaceutical compounds from harsh environments.

Traditional, noninvasive drug delivery routes such as oral or transdermal are limited in their utility. The digestive system’s harsh conditions denature or degrade many biochemical formulations. As a result, therapeutics such as insulin (degraded by proteolytic enzymes in the gastrointestinal tract) must be administered using invasive methods that reduce patient compliance and quality of life (Hamman et al., 2005; Khafagy et al., 2007; Morishita and Peppas, 2006). Topical (transdermal) formulations are user-friendly, but yield limited dosages and localized distribution, due to the low permeability of the dermal barrier (Anissimov et al., 2013; Schaefer, 1993; Singh and Roberts, 1994).

Once delivered into the body, effective therapies must overcome other challenges. A significant barrier is clearance by the immune system, which not only reduces the amount of therapeutic that can reach the target tissue, but also concentrates drugs in the liver and spleen at potentially toxic levels (Sherwood, 2015). Drug molecules that do reach tissue are distributed indiscriminately, with potentially harmful side effects to nondiseased cells (Dodd, 1987).

In recent decades, research efforts focused on the development of novel drug delivery carriers that encapsulate and protect pharmaceutics from environmental degradation agents in order to increase the effective drug lifetime and enhance therapeutic outcomes by targeting specific tissue types (Andresen et al., 2005; Beduneau et al., 2007; Chilkoti et al., 2002; Immordino et al., 2006; Jain and Jain, 2015; Kedar et al., 2010; Maeda et al., 2009; Myerson et al., 2015; Narang and Mahato, 2010; Olivier, 2005; Perez-Herrero and Fernandez-Medarde, 2015; Samad et al., 2007; Sudimack and Lee, 2000; Thanh-Huyen and Amiji, 2015; Toporkiewicz et al., 2015; Torchilin, 2004; Torchilin, 2005). These also enable control over the drug release rate (Freiberg and Zhu, 2004; Gabizon, 1995; McCoy et al., 2010; Mohammadi-Samani and Taghipour, 2015; Robinson, 1978; Timko and Kohane, 2012; Toan et al., 2015; Varde and Pack, 2004; Vasir et al., 2003; Zhu and Chen, 2015) and offer functionalities such as theranostics (Bardhan et al., 2011; Janib et al., 2010; Li et al., 2012; McCarthy and Weissleder, 2008; Xie et al., 2010).

Drug carriers also allow new methods for the delivery of degradable drugs. They can shield drugs during transport in the gastrointestinal tract, thereby allowing oral delivery, and enhance the rate of transdermal transport. In intravenous (parenteral) administration, carriers reduce clearance and prolong circulation time, and can be designed for inhalation/pulmonary delivery (Antosova et al., 2009; Breitkreutz and Boos, 2011; Chetty and Chien, 1998; Cleland et al., 2001; Fogueri and Singh, 2009; Kumar et al., 2006; Leucuta, 2012; Mathias and Hussain, 2010; Swaminathan and Ehrhardt, 2012; Ulrich, 2002).

Oncology demonstrates the need, and roles, of drug carriers. Traditional chemotherapy agents affect many tissue types, resulting in temporary and reversible side effects such as hair loss or fatigue, as well as irreversible ones such as heart and kidney damage, even failure (Aronson, 2010). Managing these side effects therefore requires limiting the maximal dosage to levels that may be less than optimal. In addition, drug effects last for a relatively short time, requiring repeated administration cycles (Barpe et al., 2010). A well-designed drug carrier can circulate for long periods in vivo, prolonging the effective release time, as well as target specific cell types so that high dosages are concentrated at tumor cells only (Amoozgar and Goldberg, 2015; Bahrami et al., 2015; Brannon-Peppas and Blanchette, 2004; Brigger et al., 2002; Byrne et al., 2008; Cho et al., 2008; Danhier et al., 2010; Davis et al., 2008; Fernandes et al., 2015; Maeda et al., 2009; Nie et al., 2007; Peer et al., 2007; Tong and Langer, 2015; Topete et al., 2015).

Drug carrying nanoparticles can be defined by various criteria, including the type of drugs they carry (hydrophobic or hydrophilic) and the carrier structure, solid/uniform or core-shell (Allen and Cullis, 2004; Anderson and Shive, 1997; Cheng et al., 2014; Cho et al., 2008; Kawaguchi, 2000; Langer, 1998; Lavan et al., 2003; Lawrence and Rees, 2000; Moghimi et al., 2001; Muller et al., 2000; Peppas et al., 2006; Torchilin, 2005). The delivery efficiency and release rates are set by the carrier physical parameters, such as the size or shape, and chemistry, in particular surface characteristics and charge. This applies to both in vitro (Elnaggar et al., 2014; Eloy et al., 2014; Kumar, 2000; Ma et al., 2013; Mallick and Choi, 2014; Manjunath et al., 2005; Mundargi et al., 2008; Nkabinde et al., 2012; Qing et al., 2013; Sinha et al., 2004; Sonam et al., 2013) and in vivo applications (Froehlich, 2012; Gill et al., 2007; Kettler et al., 2014; Mailaender and Landfester, 2009; Salatin et al., 2015; Truong et al., 2015; Win and Feng, 2005).

Amphiphilic molecules are ones that are composed of at least one hydrophilic and one hydrophobic region (Stokes and Evans, 1997). Amphiphiles whose hydrophobic and hydrophilic regions are similar in size tend to aggregate into bilayer sheets in aqueous environments, which then close to form vesicles; core/shell nanoparticles with an aqueous core surrounded by the bilayer shell (see Fig. 1.1). Their unique structure is optimize for encapsulation of hydrophilic drugs in the core, but some hydrophobic drugs can be incorporated into their shell as well (Ahmed et al., 2006). Functionalizing agents can be integrated into the shell to enhance circulation time in vivo, or to and bind selectively to specific cell types. Release can be triggered by bilayer degradation under specific conditions (e.g., temperature or pH) for control of delivery site and time.

Figure 1.1 Nanoparticles for the encapsulation of active ingredients in aqueous media may be oil-based or water-based. Oil-core systems include emulsions, stabilized by emulsifier molecules, Pickering emulsions stabilized by colloidal particles (CPs), solid-lipid nanoparticles (SLN) whose core is a solid oil phase, or nanostructured lipid carriers (NLC) whose core is a mix of solid and liquid domains. Water-cores may be stabilized by a lipid bilayer (liposomes) or a polymeric bilayer (polymersomes), or by a shell of CPs.

Bilayer-based carriers can also be used to deliver nondrug therapeutics, in particular nucleic acids (Dan and Danino, 2014). Several of these carriers have been approved by the US Food and Drug Administration (FDA) for use starting in the mid 1990s, and more are in phase II or III clinical trials (Lytton-Jean et al., 2015; Marchal et al., 2015).

Vesicle drug carriers are divided into three categories, depending on the type of molecules that compose their bilayer shells: liposomes—where the shell consists of phospholipids, polymersomes—where it is formed by synthetic block copolymers, and niosomes—formed by nonionic bilayers (Fig. 1.1).

1.2 Amphiphilic Bilayers

Molecules composed of a hydrophilic head region and hydrophobic tail region are called amphiphiles. The head contain groups that are soluble in water such as ionic, polar, or hydrogen-bond forming moieties, while the tail regions consist of saturated or unsaturated hydrocarbon chains.

Amphiphiles are categorized in various ways. Surface-active agents, or surfactants, are small molecules with one hydrocarbon tail and a small ionic or nonionic headgroup. Phospholipids are also small molecules, but have two hydrocarbon chains (Israelachvili, 2011), while polymeric amphiphiles contain hydrophobic tail polymeric block(s) bound to hydrophilic head block(s). Most polymeric amphiphiles are diblocks (i.e., the macromolecular equivalent of surfactants), although triblock and multiblock architecture also show amphiphilic behavior (Israelachvili, 2011).

All amphiphilic molecules display similar phase behavior in a single solvent, such as water, driven by a balance between molecular interactions and mixing entropy. The former is associated with the (unfavorable) interactions between the solvent and the insoluble region (e.g., water/hydrophobic tails), while the latter decreases with increasing concentration of amphiphiles. At low concentrations, entropy dominates and compensates for the unfavorable interactions, so that the molecules are individually solubilized. Once the concentration increases above a critical value, defined as the critical micelle concentration (CMC), amphiphiles assemble into aggregates where the insoluble group is shielded from the solution by the soluble group (Israelachvili, 2011). The CMC of amphiphiles depends on the solvent; in water, it increases with the size of the head group and decreases with the hydrophobic tail molecular weight (MW) (Israelachvili, 2011).

Amphiphilic aggregates display a rich range of morphologies. Classic aggregates that are formed by all types of amphiphiles include spherical micelles, wormlike micelles, and bilayer sheets (Israelachvili, 2011). However, other geometries are also found such as perforated lamellae or cubic phases (Garti et al., 2012; Wang et al., 2016). The morphology of the aggregates is set by the ratio between the (effective) volume of the head and tail groups. Molecules with tails that are small compared to the headgroups, or headgroups with a large effective size (typically due to ionic charges) favor spherical micelles. If the size of the hydrophobic and hydrophilic groups is similar, aggregates are likely to be flat locally, thereby forming bilayer sheets that close, to reduce exposed edges, into vesicles (Israelachvili, 2011).

Bilayers, which are also called membranes or lamellae, are composed of an inner layer of the hydrophobic tails, and two external layers of the hydrophilic head groups (Fig. 1.1) (Israelachvili, 2011). Bilayers can be characterized by the density of molecular packing at the water interface, defined by the interface area per molecule, Σ, and the thickness of the bilayer, h.

In lipid bilayers, the area per molecule is dominated by the head group. As a result, the thickness h increases linearly with the tail MW, while the interfacial area per molecule Σ is relatively insensitive to the tail length, or weakly decreases with it (Kučerka et al., 2011). In contrast, in diblock copolymer bilayers the interfacial area is set by a balance between the head–head interactions and the tail–tail interaction so that h varies with the square root of the hydrophobic block MW (Bermudez et al., 2002).

Although bilayers are locally flat, their self-assembled nature allows them to be bendable and stretchable. The application of an external force can significantly decrease the area per molecule (e.g., in a Langmuir trough), or increase it (e.g., in micropipette aspiration). Bending can occur through natural, thermal fluctuations or as a response to induced deformations (Israelachvili, 2011). The energetic penalties associated with such perturbations are defined by membrane moduli. The area expansion modulus, KA, accounts for the energy associated with changes in the area per molecule (Iglic, 2012), and can be defined as:

(1.1)

Eexp is the energy penalty incurred by changing the membrane area per molecule from the optimal value, given by Σ0, to a new value Σ. KA is typically in the range of 0.1–0.3 J/m² for pure phospholipid bilayers (Bouvrais, 2012). For polymeric systems, KA depends on the MW of the amphiphile, as discussed below.

Bending modes in bilayers are described by two moduli: the mean bending modulus κ, and the Gaussian modulus κG (Helfrich, 1973; Bouvrais, 2012). Combining the effects of all perturbation modes yields the elastic free energy per molecule

(1.2)

where C1 and C2 are the principal curvatures (inversely proportional to the radii). C0 is the preferred curvature of the layer, which is zero for bilayer-forming systems (Israelachvili, 2011). Experiments find that κ for phospholipid bilayers is of order 10−19 J, or 10–100 kBT (Iglic, 2012).

The exact values of the bilayer moduli vary with numerous amphiphile properties, including the tail MW, tail degree of saturation, head group structure and interactions, as well as the system conditions (ionic strength, temperature) (Bouvrais, 2012; Bouvrais et al., 2014; Claessens et al., 2004). In small-molecule systems such as surfactants or lipids, predicting the moduli values is complicated. However, in polymeric bilayers the moduli and the polymer MW can be linked. If the packing density of the polymers is fixed at Σ chains per area thickness, and the MW of the hydrophobic block is Nt, then the core thickness is linearly proportional to Nt (Milner et al., 1988) and the moduli are given by (Ball et al., 1991)

(1.3)

a1, a2 and a3 are constants of order unity. While this model strictly applies to high MW systems, it has been shown to also fit small-molecule systems and lipid bilayers (Rawicz et al., 2000).

In lipid and nonionic surfactant bilayers, the surface density is largely set by the head–head interactions, although mixing or, in lipid systems, tail chain asymmetry can significantly affect the bilayer moduli Illya et al. (2005). In diblock copolymer systems, however, the surface density is set by a balance between the head and tail block MW (Halperin et al., 1992; Israelachvili, 2011), so that:

(1.4)

where dcore is the thickness of the bilayer core (the hydrophobic region), and Lh is the thickness of the hydrophilic head region. Nh is the MW of the hydrophilic block, and γ is the surface tension between the hydrophobic tail block and the aqueous solution (numerical prefactors of order 1 were omitted).

1.3 Liposomal Drug Carriers

The earliest studies of vesicle drug carriers focused on phospholipid-based systems. The inherent biocompatibility of the phospholipids makes these liposomes nontoxic, biologically inert, biodegradable and (relatively) immunogenic—optimal characteristics for in vivo applications (Allen and Cullis, 2013; Andresen et al., 2005; Barenholz, 2001; Drummond and Fong, 1999; Drummond et al., 2000; Gregoriadis, 1995; Langer, 1998; Lasic and Papahadjopoulos, 1998; Lian and Ho, 2001; Malam et al., 2009; Senior, 1987; Sharma and Sharma, 1997; Torchilin, 2005).

Numerous types of phospholipids have been studied for drug delivery applications, with all types of head groups: nonionic, cationic or zwitterionic, i.e., contain both an anionic and cationic charges. Most often, they contain phosphatidylcholine (PC), phosphatidylethanolamine (PE) or sphingomyelin. Anionic phospholipids are rarely used in drug delivery applications, however, since their negative charge inhibits uptake through the anionic cell membrane (Hillaireau and Couvreur, 2009; Opanasopit et al., 2002; Semple et al., 1998). Optimal performance frequently requires mixing different species of lipids with additional components, such as cholesterol, to control bilayer fluidity and moduli (Henri et al., 2015; Patel et al., 2015; van Hoogevest and Wendel, 2014; Giddam et al., 2012), or with fusogenic peptides to facilitate uptake and release (Shete et al., 2014).

Drug delivery performance depends not only on the lipid bilayer composition, but also on the properties of the liposome. Unilamellar vesicles are coated by a single bilayer, and may be small unilamellar vesicles (SUV), on order 100 nm, or giant unilamellar vesicles (GUVs) that are of order 10 µm. The latter are too large for clinical applications, but useful for in vitro studies (Bae and Park, 2011). Multilamellar vesicles (MLV), which are coated by a several bilayers, are typically 1 μm, slightly larger than the cutoff limits for cellular uptake but below the limit for rigid particles’ circulation in capillaries (Bae and Park, 2011). Multilamellar vesicles are more stable than unilamellar ones, due to the multilayered shell, and have a larger volume available to incorporate hydrophobic drugs, but tend to be more polydisperse in size.

One of the main issues in drug delivery is prolonging circulation time so that carriers can reach their target tissue. Longer circulation times may also reduce the need for frequent administrations. Although composed of biological components, liposomes can trigger phagocytosis and clearance, with typical blood circulation times of order hours (Blume and Cevc, 1990; Ishida et al., 2002; Senior et al., 1985). Liposome clearance rate was found to decrease with vesicle size (Juliano and Stamp, 1975; Senior et al., 1985; Senior et al., 1983) and depends on surface charge and bilayer fluidity (Gabizon and Papahadjopoulos, 1992; Levchenko et al., 2002; Nishikawa et al., 1990).

Understanding of the immunoclearance process allowed for the development of methodologies to prolong circulation. The process is triggered by adsorption of serum opsonin proteins (e.g., beta 2-macroglobulin, immunoglobulins, fibronectin) onto the particle surface (Immordino et al., 2006), so that retardation of protein adsorption increases circulation time (Chonn et al., 1992). In principle, controlling liposome surface charge or chemistry could achieve that. However, any such strategy is limited to a narrow selection of the rich variety of immune-proteins. For example, cationic surface charge would repel cationic proteins, but attract anionic ones and vice versa. A more effective and universal reduction in protein adsorption is obtained by incorporation of steric inhibitors, polymeric chains that mask the surface and prevent adsorption to form so-called stealth liposomes (Gref et al., 2000; Allen, 1994; Barenholz, 2001; Immordino et al., 2006; Lasic and Needham, 1995; Lian and Ho, 2001; Moghimi and Szebeni, 2003; Sharma and Sharma, 1997; Nag and Awasthi, 2013). This is done by incorporation of PEG-lipids (polyethylene glycol chains attached to a lipid molecule) into the bilayer, as shown in Fig. 1.1. However, the PEG slows down, but does not eliminate, protein adsorption so that clearance does occur, albeit more slowly than in non-PEGylated liposomes (Halperin, 1999).

Stealth liposomes were found to have longer circulation times than traditional ones, with resulting differences in their tissue distribution (Allen and Hansen, 1991; Allen et al., 1995; Cattel et al., 2004; Gabizon and Martin, 1997; Oku and Namba, 2005; Woodle and Lasic, 1992). Circulation time increases with the PEG-lipid content and with the PEG MW (Mosqueira et al., 2001; Gref et al., 2000; Moghimi and Szebeni, 2003) until reaching a plateau (Mosqueira et al., 2001; Gref et al., 2000; Moghimi and Szebeni, 2003). This is due to the fact that increasing the PEG MW increases the effective head-group size, which can cause a transition from bilayers to spherical micelles (Hristova et al., 1995) or limit the maximal amount of PEG-lipids soluble in the bilayer. It is important to note that, due to the self-assembly process, excess PEG-lipids form micelles so their concentration in the bilayer is not necessarily the same as their overall concentration in solution.

Although PEG is the most commonly used inhibitor of protein adsorption in liposome formulations, other types of polymers have also been used in stealth liposomes, with similar degrees of success (Amoozgar and Yeo, 2012; Cao and Jiang, 2012; Hu et al., 2014; Kierstead et al., 2015; Kohli et al., 2014; Mohr et al., 2014; Nag et al., 2015; Reibel et al., 2015; Sun et al., 2015).

In addition to PEGylation, liposomes have been formulated with a variety of components that enable targeting of specific cells types, in particular cancer cells (Bazak et al., 2015; Caracciolo, 2015; Fernandes et al., 2015; Zhang et al., 2015; Khare et al., 2014; Lin et al., 2015; Ran et al., 2016; Sakurai et al., 2015; Werengowska-Ciecwierz et al., 2015). However, despite demonstrations of enhanced uptake (Akhtari et al., 2016; Bazak et al., 2015; Hayward et al., 2016; Tuo et al., 2016; Ye et al., 2016) to date such formulations have not been approved for clinical use (Fathi and Oyelere, 2016). In general, incorporation of targeting ligands can disrupt the aggregate structure by changing the preferred packing (Israelachvili, 2011), which limits degree of bilayer functionalization and therefore the efficacy of targeting strategies.

To protect their cargo, drug carriers need to remain stable during administration and circulation, a role that liposomes perform well. In addition, due to their hydrophobic/hydrophilic layer structure and dense packing, there is little or no drug leakage from liposomes over prolonged periods (or order months). This advantage becomes a disadvantage, however, once the liposomes reach their target and the drug must be released. Thus, liposomal bilayers often contain some mechanism that will allow their controlled destabilization, namely, response to some stimulus (Blenke et al., 2013; Felber et al., 2012; Kneidl et al., 2014; Moussa et al., 2015; Movahedi et al., 2015; Paliwal et al., 2015). The trigger may be an altered conditions in diseased tissue such as pH (Zhu et al., 2011) or external fields such as ultrasound (Paliwal et al., 2015; Bibi et al., 2012; Boissenot et al., 2016; Moussa et al., 2015; Movahedi et al., 2015; Ta and Porter, 2013; Tila et al., 2015; Torchilin, 2014; Zhao and Rodriguez, 2013), or heat (Al-Ahmady and Kostarelos, 2016; Dabbagh et al., 2015; Dicheva and Koning, 2014; Haeri et al., 2016; Song et al., 2016).

The most common environmental trigger for liposomal degradation in specific tissues is pH, since inflamed tissue and tumors are more acidic than healthy tissue (Drulis-Kawa and Dorotkiewicz-Jach, 2010; Shi et al., 2002). Such liposomes are therefore designed to be stable under physiological conditions, but destabilize in more acidic conditions—thereby allowing release of the encapsulated drug. pH sensitivity may be obtained by using unsaturated phosphatidylethanolamine (e.g., dioleoyl phosphatidyl ethanolamine—DOPE, diacetylenic-phosphatidyl-ethanolamine-DAPE) (Karanth and Murthy, 2007) or pH-sensitive peptides or proteins (Simões et al., 2004).

Thermally sensitive liposomes disintegrate and release their encapsulated drug when a temperature jump occurs. This may be take place as the result of the inherent hyperthermia of diseased tissue, or through external application of heat to a particular region (Kono et al., 2015; Dabbagh et al., 2015; Kneidl et al., 2014). Thermosensitive liposomes often incorporate dipalmitoylphosphatidylcholine (DPPC), since its gel-to-liquid crystalline phase transition, which disrupts the bilayer packing and allows leakage, occurs at 41°C (Kim et al., 2014; Lee et al., 2014; May et al., 2013).

Potentially the most effective field for triggered release is ultrasound (US) in the clinically approved diagnostic fields of 1–18 MHz (Novell et al., 2015; Wallace and Wrenn, 2015; Ahmed et al., 2015; Husseini et al., 2014; Liu et al., 2015; Moussa et al., 2015; Sirsi and Borden, 2014; Udroiu, 2015). Ultrasound has been proven to be benign in decades of clinical imaging, and can penetrate into deep tissue not accessible to other fields (Husseini et al., 2014). Liposomal disintegration by US occurs through two mechanisms: (1) Local heating due to absorption of acoustic energy, that can be used to obtain release from thermosensitive liposomes, or (2) cavitation that occurs when gas filled bubbles respond to the pressure caused by the US field by oscillating, leading to bilayer rupture (Dijkmans et al., 2004; Schroeder et al., 2009; Wrenn et al., 2012). Unfortunately, the similarity between liposomes and cell membranes means that cellular membranes may also rupture under US (Krasovitski et al., 2011), limiting the strength and duration of application. Recent formulations incorporate gas bubbles in the liposomes, so that US application causes the bubbles to undergo volume changes, yielding shear forces that porate or rupture the bilayer at durations of field strength that do not harm cellular membranes (Geers et al., 2012; Husseini et al., 2014; Schroeder et al., 2009; Stride and Coussios, 2010; Wrenn et al., 2012; Wrenn et al., 2013).

Recent studies find that low frequency US (LFUS), at frequencies >1 MHz, can also effectively and transiently increases the permeability of lipid bilayers even when they are not thermo-functionalized nor incorporate bubbles (Enden and Schroeder, 2009; Schroeder et al., 2007; Small et al., 2012; Small et al., 2011; Pong et al., 2006; Pong et al., 2005). It has been suggested that the poration mechanism in this case differs from those known to occur in the high frequency case. Analysis shows that partitioning of dissolved gases from water into the hydrophobic bilayer can lead to the formation of microbubbles in the bilayer core. The response of these to the field ruptures the bilayer and allows drug leakage (Wrenn et al., 2013).

Magnetic fields are also benign and penetrate deep into tissue. Magnetically-doped liposomes have been synthesized by inclusion of magnetic nanoparticles in the bilayer, so that release is triggered either as the result of heating or as through field-induced of stresses developing under an applied magnetic field (Bi et al., 2016; Bixner and Reimhult, 2016; Salvatore et al., 2016; Shah et al., 2016; Yang et al., 2016).

Commercial production of drug carriers requires the ability to synthesize them in a controlled and sterile manner, with high drug loading efficiency. While liposomes form spontaneously when phospholipids are dissolved in water, the result is polydisperse, multilamellar vesicles (MLVs) that cannot be used unless they are homogenized by sonication or high-pressure value homogenization. Rehydration produces unilamellar liposomes (Taylor et al., 2005). Loading the drugs by dissolving them in the aqueous solution during the liposomal formulation stage results in low loading efficiency (McClements, 2015), which may be improved for a limited number of compounds by using external fields. Advances in microfluidics methods promise the most effective and consistent way to synthesize liposomes, yielding monodisperse, unilamellar vesicles of any target size and with 100% loading efficiency (Hassan et al., 2014; Jahn et al., 2004; Jahn et al., 2010; Matosevic, 2012; Yu et al., 2009; Vladisavljevic et al., 2012).

1.4 Polymersome Drug Carriers

Amphiphilic block copolymers are synthetic polymers composed of hydrophobic blocks—high MW sequences that are water soluble, and hydrophobic blocks—long sequences that are immiscible in water. Diblock copolymers contain one hydrophobic and one hydrophilic block, a macromolecular equivalent of small molecule surfactants. Indeed, their phase behavior in aqueous solutions largely follows that of surfactants, where the geometry of the aggregates is set by the MW ratio between the two blocks, in a similar manner to the ratio between head and tail volume in small molecule amphiphiles (Hamley, 2004). Multiblock copolymers also form aggregates in water, displaying a much richer range of phase geometries.

As in the case of small-molecule surfactants, diblock copolymers with similar hydrophobic and hydrophilic block volumes (which translate to similar MWs) form bilayers in water, with the hydrophobic block shielded from the solvent by the hydrophilic one (see Fig. 1.1). However, the bilayer thickness may be an order of magnitude or larger than that of lipid bilayers, depending on the copolymer MW. Polymersomes, the polymeric equivalent of liposomes, are vesicles whose shell is composed of a block copolymer bilayer, although the term is also sometimes used to vesicles whose shell is composed of a multiblock copolymer (Discher et al., 1999; Discher and Ahmed, 2006; Guan et al., 2015; Krishnamoorthy et al., 2014; Lee and Feijen, 2012; Letchford and Burt, 2007; Meng et al., 2009b; Messager et al., 2014).

Polymersomes for drug delivery applications usually use copolymers with a poly(ethylene glycol) (PEG) hydrophilic block, since it has been shown to be highly biocompatible in general, and effective in stealth liposomes in particular. Drug delivery polymersomes have been formed from PEG copolymers bound to hydrophobic blocks such as poly(ethyl ethylene) (PEE) (Bermudez et al., 2002), poly(butadiene) (PBD) (Bermudez et al., 2002, Li et al., 2007), or poly(styrene) (PS) (Kabanov et al., 1998). Another commonly-used hydrophilic block is polyethylene oxide (PEO), bound to polycaprolactone (PCL) (Levine et al., 2008), PBD (Li et al., 2007; Demirgoez et al., 2009), or poly(γ-methyl-ε-caprolactone) (Petersen et al., 2013). More recently, interest focused on the use of copolymers with biomolecular blocks such as folic acid-poly(L-glutamic acid)-block-poly(ε-caprolactone) (Liu et al., 2017), dextran-block-poly lactide-co-glycolide (Alibolandi et al., 2016), poly(glutamic acid)-b-polyphenylalanine (Vlakh et al., 2016) or poly(γ-benzyl-L-glutamate)-block-hyaluronan (Upadhyay et al., 2012). Polymersomes based on triblock copolymers also use similar chemistries (Cao and Jiang, 2012; Gallon et al., 2015; Nardin et al., 2000).

As noted, liposomes cannot support a high concentration of functionalized moieties since those tend to disrupt the bilayer structure. Polymeric bilayers are much more robust due to the high MW of the blocks, so that relatively high concentrations of targeting or degradation groups can be included without affecting the bilayer stability (Felber et al., 2012; Guan et al., 2015; He et al., 2012; Krishnamoorthy et al., 2014; Liu et al., 2013; Meng et al., 2009a; Onaca et al., 2009; Pawar et al., 2013; Petersen et al., 2013; Rijcken et al., 2007; Thevenot et al., 2013). Targeting moieties used in polymersomes include antibodies, peptides, or small molecules such as folate or biotin (Egli et al., 2011; Pawar et al., 2013; Tanner et al., 2011), with demonstrated success (Licciardi et al., 2010; Pang et al., 2010; Prabhu et al., 2015; Upadhyay et al., 2010b; Upadhyay et al., 2012; Zhu et al., 2015a).

In stealth liposomes, the incorporation of water-soluble polymer chains (PEG-lipids) in bilayer reduces immuno-protein adsorption, thereby increasing circulation times. However, the amount of PEG-lipid that can be incorporated is limited. In polymersomes, the entire hydrophilic layer is composed of a water-soluble polymer, providing a highly effective barrier for immunoprotein adsorption and yielding circulation times that exceed those of PEGylated liposomes (Photos et al., 2003; Lee and Feijen, 2012). As may be expected, the circulation time varies with the MW of the hydrophilic block (Photos et al., 2003), since longer chains provide a thicker barrier to adsorption. As in the case of stealth liposomes, the hydrophilic polymer layer retards, but does not suppress, immunoclearance. Phagocytosis does occur eventually, leading to polymersome accumulation in the liver (Photos et al., 2003; Lee et al., 2011b).

In Fig. 1.2A, the measured thickness, of the polymersome bilayer core (using data from Photos et al., 2003) is plotted as a function of the predicted thickness based on Eq. 1.4, showing the expected linear correlation. However, plotting the circulation time (as defined by a characteristic timescale, T1/2) as a function of the PEG-layer thickness shows larger variations (Fig. 1.2B). These may be due to the inherent randomness of complex biological systems. However, it should be noted that other factors, in addition to the PEG layer thickness, may affect circulation time. Quite surprisingly, the polymersome size also affects circulation time, showing a relatively sharp transition from short (~hours) to long (~day) circulation time at vesicle sizes of order 100 nm (Brinkhuis et al., 2012).

Figure 1.2 Effect of liposome radius on calcein release from liposomes composed of 1-palmitoyl-2-oleoyl-sn-glycero-3-phosphocholine (POPC), dipalmitoyl-phosphocholine (DPPC), and cholesterol. Blue circles: R=86.7 nm, brown diamonds: R=113.8 nm. M(t)/M0 is the fraction of encapsulated marker remaining in the nanoparticle at time (t). The release profiles differ significantly when plotted versus time, but collapse to the same plot when t is reduced by R² as predicted by Eqs. (1.2) and (1.4), as shown in the inset (). M(t)/M0 is the fraction of encapsulated marker remaining in the nanoparticle at time (t). Based on data from Small et al., 2012.

The high MW of the copolymers forming polymersome bilayers yields thicker shells whose area and bending moduli are higher than those of liposomes (see Eq. 1.3), resulting in better mechanical stability and shear resistance (Discher et al., 1999; Discher and Ahmed, 2006; Letchford and Burt, 2007; Meng et al., 2009a). Mechanical stability may be further enhanced by crosslinking of the hydrophobic layer (Cheng et al., 2011; Gaitzsch et al., 2011; Chevalier and Bolzinger, 2013; Thibault et al., 2006; Xu et al., 2009). In addition, the thick hydrophobic core allows solubilization of hydrophobic drugs in appreciable amounts so that polymersomes can deliver a combination of hydrophilic and hydrophobic drugs (Ahmed et al., 2006; Pramod et al., 2014; Surnar and Jayakannan, 2013; Thambi et al., 2012; Wu et al., 2014).

The enhanced stability of polymersomes is highly useful during storage and circulation, but becomes a disadvantage during the drug release stage. Bilayer disintegration—which causes a burst like release profile—can be achieved by using pH or temperature to degrade the hydrophobic block. The decrease in the insoluble block’s MW disrupts the bilayer structure and leads to vesicle disintegration (Gillies and Fréchet, 2005; Jeong et al., 2015; Jeong et al., 2013; Kim et al., 2012; Rodríguez-Hernández and Lecommandoux, 2005; Qing et al., 2013; Qiu et al., 2013). More sophisticated chemistries may be used to break the bond between the hydrophobic and hydrophilic blocks, e.g., by disulphide cleavage (Cerritelli et al., 2007), which also causes bilayer destruction and vesicle disintegration. Light-induced cleavage, has also been used for this purpose (Cabane et al., 2010; Li et al., 2013), but it is not effective for in vivo applications. Fields, in particular magnetic ones, have also been used to break polymersomes (Bixner et al., 2016; Oliveira et al., 2013b; Oliveira et al., 2013a). Using biodegradable polymers such as PLGA to form the polymersome bilayer allows for a slow release profile, where the rate at which the drug diffuses out may be controlled through the degradation rate (Kim et al., 2006; Meng et al., 2005; Meng et al., 2009b).

Polymersomes synthesis is similar to that of liposomes. As a result, although they form spontaneously in water, the encapsulation efficiency is low (Messager et al., 2014). One methodology that can yield high encapsulation efficiency is that of double emulsion, where the inner, aqueous solution contains the hydrophilic drug, and is coated by a shell of an oil phase stabilized by the copolymer (the oil phase may also contain a hydrophobic drug) in another, continuous aqueous solution. The intermediate oil phase is removed, leaving the inner phase coated by a polymeric bilayer (Lorenceau et al., 2005). This method is especially effective when using microfluidic devices, allowing 100% encapsulation efficiency and yielding highly monodisperse polymersomes, with sizes that can be exactly controlled over a large range (~10 nm–100 µm) (Brown et al., 2010; Duncanson et al., 2012; Kim et al., 2011; Kim et al., 2013; Liao et al., 2012; Shum et al., 2008; Shum et al., 2011; Thiele et al., 2010). While effective, this method must be conducted carefully to reduce complications such as bilayer inhomogeneities (Hayward et al., 2006).

Despite the synthetic nature of the polymersome copolymers, numerous studies have demonstrated their safety in vivo, mostly in murine models (Chen et al., 2013; Jain et al., 2011; Petersen et al., 2013; Petkar et al., 2011; Upadhyay et al., 2010a; Upadhyay et al., 2012; Xu et al., 2014). However, there may be some concern regarding toxicity due to their accumulation in different organs (e.g., liver, spleen).

1.5 Niosome Drug Carriers

Like lipids and diblock copolymers, nonionic surfactants can assemble into bilayers, thereby forming niosomes—the small-molecule synthetic analogue of liposomes and polymersomes (Schreier and Bouwstra, 1994; Uchegbu and Florence, 1995; Uchegbu and Vyas, 1998). While the low cost of niosomes led to interest in their application for uses that require bulk quantities, such as cosmetics (Aljuffali et al., 2015; Mazda et al., 1997; Tavano et al., 2014; Wu and Guy, 2009), niosomes have the potential to be an effective drug delivery vehicle as well (Drummond and Fong, 1999; Fang et al., 2001; Gaudana et al., 2009; Kaur et al., 2004; Uchegbu and Vyas, 1998).

Unlike the majority of phospholipids that favor bilayer aggregates, nonionic surfactants display a rich range of aggregate geometries (Israelachvili, 2011). In contrast to diblock copolymers, where the MW ratio of the blocks can be tailored to yield bilayers, the molecular structure of nonionic surfactants is fixed. This limits the types of nonionic surfactants that can be used to form niosomes (Salim et al., 2014). One category of vesicle-forming surfactants is based on natural or synthetic sugar-based molecules, with head groups of mono or poly saccharide: glucose, lactose, maltose, or sucrose (El-Laithy et al., 2011a,b; Faivre and Rosilio, 2010; Muzzalupo et al., 2013). Sorbitan esters (Spans) are fatty esters of the cyclized derivatives of sugar alcohol sorbitol-sorbitan, and have also been used to form niosomes (Balakrishnana et al., 2009; Guinedi et al., 2005; Hao et al., 2002; Uchegbu and Florence, 1995; Varshosaz et al., 2003; Yoshioka et al., 1994), as were polyethoxylated sorbitan esters (tween) mixtures with cholesterol (Alsarra et al., 2005; Bayindir and Yuksel, 2010; Di Marzio et al., 2011; Manosroi et al., 2003). In many systems, cholesterol is mixed into the bilayer, usually forming hydrogen bonds with the head-groups and increasing the mechanical stability of the vesicles(Seleci et al., 2016).

Niosomes can be synthesized in various methods, including thin film hydration and microfluidics (Seleci et al., 2016). The encapsulation efficiency in niosomes is commonly in the range of 10%–40% (Abd-Elbary et al., 2008; Guinedi et al., 2005; Yoshioka et al., 1994), and depends on the type of surfactant, mixture composition, and synthesis methodology, although in some cases in can increase to ~75%–90% (Abdelbary and El-Gendy, 2008; Waddad et al., 2013).

Although the thickness of niosome bilayers may be similar or even larger than that of lipid bilayers (Pozzi et al., 2010), studies show that the drug leaks from niosomes, with a rate of release that is characteristic of diffusive processes (Abd-Elbary et al., 2008; Abdelbary and El-Gendy, 2008; Akbari et al., 2015; Guinedi et al., 2005; Gopinath et al., 2004; Waddad et al., 2013). In addition, potential interactions between the encapsulated therapeutics and the surfactant shell may lead to drug incorporation in the bilayer, or disruption of the head–head interactions (Uchegbu and Vyas, 1998).

1.6 Biomedical Applications

Drug carriers can deliver their cargo via several routes (Fig. 1.3). The most common one is intravenous. To date, several uncoated or PEGylated (stealth) liposome-based therapeutics have been approved by the FDA or other agencies, and are now commercially available (see Table 1.1). Many others are in clinical trials for conditions ranging from fungal infections to hemophilia (Zou et al., 2015; Sercombe et al., 2015; Bozzuto and Molinari, 2015).

Figure 1.3 Drug administration routes: Ocular, through the eye; Oral, through the gastrointestinal system; Transdermal (topical), through the skin; Intravenous, injected directly into the bloodstream. Data from R.T. Rosenberg and N. Dan, unpublished.

Table 1.1

Comparing the Properties of Nanoparticles for the Encapsulation of Hydrophobic Compounds in Aqueous-Based Foods

aCan vary greatly, depending on the specific formulation.

Cancer therapy and tumor targeting are the most widely studied application of vesicle-based drug delivery systems. The low specificity of direct drug administration (whether oral, transdermal, or intravenous) means that healthy tissue, as well as disease sites, is affected by the cancer chemotherapy agents. Some side effects of these toxic compounds, such as fatigue or hair loss, are usually transient. However, others such as cardiotoxicity, may be life threatening. To minimize risks, chemotherapy dosages are therefore kept at values lower than the optimal level (Ewer and Yeh, 2006; Bocci and Francia, 2014).

Drug encapsulation can address many of these issues. Encapsulation shields toxic therapeutics from interactions with tissues before release is triggered. Long circulation times allow effective biodistribution of the carriers and their cargo, and can also reduce the need for frequent administration. Indeed, the first liposome-based systems to receive FDA approval and commercial distribution in the 1990s were introduced for treatment of Kaposi’s sarcoma, ovarian cancer, multiple myeloma, and metastatic breast cancer. Doxil/Caelyx (Janssen Pharmaceutica NV, Beerse, Belgium), which utilize stealth-liposomes, and Myocet (GP Pharm SA, Barcelona, Spain/Teva Pharmaceutical Industries Ltd, Krakow, Poland), an uncoated-liposome formulation. Patients treated with liposomal drug-carrying formulations show response rates that are comparable or higher than either unmodified drugs or traditional therapy methods, in breast, advanced or recurrent ovarian, or prostate cancer, as well as other types of tumors (Gabizon et al., 2003; Singla et al., 2002; Allen and Cullis, 2013; Batist et al., 2001; Muggia et al., 1997; Harris et al., 2002; Lyass et al., 2000; Ranson et al., 1997; Hubert et al., 2000; Beedassy and Cardi, 1999; Gokhale et al., 2001; Tejada-Berges et al., 2002; Marchal et al., 2015; Tila et al., 2015; Zuccari et al., 2015). Furthermore, liposomal therapies were shown to reduce toxicity (when compared to free drugs), although some side effects such as skin and hypersensitivity reactions or nausea were found (Iwamoto, 2013). Similarly, polymersomes composed of biodegradable polymers were found to increase the tolerated dosage and shrink tumors compared to either the free drugs (Ahmed et al., 2006; Bakalova et al., 2015) or other types of nanocarriers, such as polymeric micelles (Chen et al., 2010). A number of studies also examined the efficacy of niosomes for treatment of such cancers as melanoma (Dwivedi et al., 2015; Gude et al., 2002), breast, and ovarian cancer (Han et al., 2013; Shaker et al., 2015; Uchegbu et al., 1996), demonstrating their efficacy.

Drug carriers can be synthesized to target specific cell types—in particular cancer cells—thereby yielding localized high drug concentrations in the diseased tissue without compromising patient safety. Carrier targeting to tumors can be passive or active. Passive tumor targeting is based on the so-called leaky vasculature effect, namely, the enhanced permeability and retention effect (EPR) in tumors. The rapid growth of tumors leads to hyperpermeability associated with the poor histological organization of cancer cells’ limited lymphatic drainage (Brigger et al., 2002; Cho et al., 2008; Danhier et al., 2010; Rapoport, 2007). Particles of order 380–780 nm tend to accumulate selectively, therefore, in tumors (Hobbs et al., 1998). Conveniently, the appropriate particle size range is obtainable in vesicles of all types (lipid, polymer or nonionic surfactant based), either by separating a polydisperse mixture (as in those formed via film rehydration, for example) or by using microfluidics methods that produce monodisperse particles.

While the EPR effect enables high concentrations of drug-carrying particles in tumors, it cannot completely eliminate toxic side effects. Smaller particles (>10 nm) are quickly cleared via the renal system, but drug carrying particles of order 100–1000 nm accumulate in the liver, spleen and bone marrow (Almeida et al., 2011; Simon and Sabliov, 2014; De Jong et al., 2008; Sonavane et al., 2008), with associated adverse side effects (Bae and Park, 2011). Thus, while size-based passive targeting enables delivery of high drug concentrations to the target tumors, it still requires a delicate optimization of dosages.

An alternative to passive targeting is active targeting, where ligands that bind to overexpressed surface molecules or receptors unique to cancer cells are incorporated into the vesicle surface. Binding between the vesicle and the target cell can either cause vesicle disintegration (and therefore drug release outside the cell), or, more often, initiate receptor-mediated endocytosis (Katsogiannou et al., 2011; Trapani et al., 2012). Targeting of CD44 through such molecules as hyaluronic acid (Mattheolabakis et al., 2015; Tripodo et al., 2015) was found to be an effective method for liposomal binding and targeted delivery (Dalla Pozza et al., 2013; Dufay Wojcicki et al., 2012; Platt and Szoka, 2008; Qhattal et al., 2014; Wojcicki et al., 2012), but other targets such as the scavenger receptor type B1 (SR-B1) (Yuan et al., 2013) or folate receptors (Alibolandi et al., 2016; Lu et al., 2016; Ran et al., 2016) may be promising routes. Functionalized polymersomes are a more promising system for active delivery, since the polymeric shell can carry a larger density of ligands without destabilization (Guan et al., 2015; Krishnamoorthy et al., 2014; Oltra et al., 2014; Prabhu et al., 2015), utilizing such targeting moieties as anti-EGFR antibodies (Lee et al., 2011a; Li et al., 2012) or a PR_b integrin peptide (Demirgoez et al., 2009; Petersen et al., 2013). Niosomes with active targeting have also been studied (Kong et al., 2013; Seleci et al., 2016). However, despite successes in the lab, no vesicle-based targeted delivery systems has approved for clinical use to date (Fathi and Oyelere, 2016).

The utility of vesicle-based carriers is not limited to oncology. Some capitalize on a weakness of drug carriers, i.e., their uptake by macrophages, to target these cells. Liposomes carrying antiparasitic or antimicrobial drugs were shown to be effective for treatment of mononuclear phagocytic infections (Karlowsky and Zhanel, 1992; Pinto-Alphandary et al., 2000), while polymersomes were used to target leukocytes (Robbins et al., 2010) and macrophages (Broz et al., 2005). Liposome uptake by macrophages introduced antiretroviral drugs to HIV infected immune and central nervous systems cells without affecting other cell types, though this method is unlikely to become widespread due to limited drug-loading capacity of the hydrophobic drugs, short shelf life and liposome cost (Desormeaux and Bergeron, 1998; Gupta and Jain, 2010; Mallipeddi and Rohan, 2010; Ojewole et al., 2008; Ramana et al., 2014; Sagar et al., 2014).

Vesicle uptake and clearance by macrophages also allows targeting of the liver, spleen, and bone marrow (Maesaki, 2002; Vyas and Gupta, 2006; Nobili et al., 2006), or inflamed tissue such as rheumatoid arthritis (van Rooijen and Sanders, 1997). Functionalization of the liposomes provides additional tissue specificity (Varghese et al., 2014; Nogueira et al., 2015). Studies examined liposomes as drug delivery vehicles to unresolved inflammation following myocardial infarction, thereby delaying heart failure onset in mice (Kain et al., 2015), and to attenuate atherosclerosis and reduce atherosclerosis-dependent myocardial infarction and stroke (Hosseini et al., 2015), as well as to treat neuropathic pain.(Kobayashi et al., 2015).

Development of drug carriers that can cross the blood–brain barrier (BBB) is a challenge, since the barrier effectively suppresses transport of many drugs directly into the brain. It is estimated that the BBB allows less than 1% of molecules in the circulatory system to cross into the brain (Soni et al., 2016). Carriers that can cross the BBB will therefore permit treating cerebral diseases, from brain tumors to Alzheimer’s and Parkinson’s.

Nanoparticle access to the brain can be obtained using surface receptors on the brain capillary endothelial cell (BCECs), such as transferrin, lipoprotein, or insulin-like growth factor receptors. These are present at higher concentrations than in other tissue, to compensate for the inhibited transport of larger molecules (~400 Da) through the tight junctions between the BBB cells. Different methodologies have been used to tailor nanoparticles to cross the BBB, typically by functionalizing or coating them in molecules such as polysorbates. Several studies examined liposomes (Chen and Liu, 2012; Tiwari and Amiji, 2006; Jain and Jain, 2015; Larsen et al., 2014; Tajes et al., 2014; van Tellingen et al., 2015), as well as functionalized polymersomes (Larsen et al., 2014; Stojanov et al., 2012; Pang et al., 2008; Xin et al., 2011; Yu et al., 2012; Chen et al., 2014) and niosomes (Bragagni et al., 2014; Dufes et al., 2004; Jain and Jain, 2015) for this purpose. Although some results seem promising, it is not clear that many formulations actually increase the delivery rate (Johnsen and Moos, 2016). In particular, the development of effective therapeutic carriers is hindered by a lack of consistent quantitative data that