Polymeric Nanoparticles Explored for Drug-Delivery Applications

The integration of nanotechnology with medicine leads ... endothelial barrier, defective vascular architecture and an impaired lymphatic drainage syst...
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Polymeric Nanoparticles Explored for Drug-Delivery Applications Heba Asem and Eva Malmström* Department of Fibre and Polymer Technology, School of Chemical Science and Engineering, KTH Royal Institute of Technology, Teknikringen 56, SE-100 44, Stockholm, Sweden *E-mail: [email protected].

The main drawback of conventional chemotherapeutics is their non-specific distribution in the body which causes serious side effects to healthy cells. As a consequence, the drug concentration reaching the tumor is reduced, resulting in suboptimal therapeutic efficacy. The discovery that polymer-based nanomaterials can be used for controlled drug delivery systems offers well-defined reservoirs for a wide spectrum of pharmaceutical agents, with the ability to reduce the toxic response. The most widely explored polymeric nanocarriers, including biodegradable polymers, amphiphilic copolymers and polymers that form unimolecular micelles, are discussed in this brief chapter.

Introduction Personalized medicine is a rapidly growing and important research area which can lead to significantly more efficient and patient resilient treatment, giving societal benefits such as improved quality of life for patients and lower costs for care. Personalized medicine has stimulated the design of various delivery platforms for pharmaceuticals and chemotherapeutics, in particular to overcome the limitations with respect to toxicity encountered by the conventional delivery of drugs (1, 2). The integration of nanotechnology with medicine leads to the development of tailored nanomaterials facilitating a personalized and efficient therapy. In this respect, nanoparticles (NPs), typically with dimensions © 2018 American Chemical Society

below 200 nm, have gained increasing interest for the efficient delivery of active components (3). Such NPs can be based on inorganic materials or on soft matter, such as silica NPs and others or polymeric macromolecules (Figure 1). NPs have also been shown advantageous in solubilizing therapeutic cargos, substantially prolonging the circulation life times of drugs. Particle size and size distribution are the major factors affecting the circulation, extravasation through vasculature leakage, macrophages uptake and renal clearance of NPs upon intravenous administration (4). Biological barriers to deliver NPs to their destination prevent the delivery of drugs and hence decrease their therapeutic activity. These barriers include opsonization, renal clearance and mononuclear phagocyte system (MPS). Particles with a diameter greater than 200 nm are sequestered and accumulated in liver and spleen by MPS (4, 5). The renal filtration threshold of NPs is · 10 nm in diameter and the renal molecular weight cutoff size of polymers is · 48 kDa (6). NPs of intermediate sizes, 20-100 nm, thus have the optimal potential for in vivo applications, owing to their ability to circulate in the blood stream for long periods of time. These NPs are large enough to evade renal clearance and small enough to avoid rapid clearance from circulation. Due to the inherent size of NPs, they can be selectively accumulated in the tumor tissue through the enhanced permeation and retention (EPR) effect which is characterized by the presence of leaky vasculatures having an incomplete endothelial barrier, defective vascular architecture and an impaired lymphatic drainage system of the tumor (7). Polymers have attracted significant attention as therapeutic carriers due to their inherent physical, chemical, and biological properties and their highly beneficial size. Polymeric nanoparticles (PNPs) offer significant advantages over inorganic NPs with regard to design flexibility and biological fate. Inorganic NPs (gold, superparamagnetic iron oxide, silica NPs and quantum dots) are potential carriers for the cellular delivery of various drugs, but inorganic NPs have to be subjected to chemical and/or biological modification to meet the stringent requirements for cellular delivery, such as good biocompatibility (8). Moreover, it may be difficult to efficiently clear these NPs from the body due to lack of biodegradability, and they are therefore exposed to various intracellular degradative enzymes to improve their clearance from the body (9). PNPs offer great versatility in terms of composition and functionalization when platforms for effective therapeutic delivery are being designed (10). Biodegradable polymers which can be degraded chemically or enzymatically and excreted safely from the body through the renal system have emerged as promising materials for drug delivery systems (DDS) due to their biodegradability and biocompatibility (11). Most therapeutic drugs are released rapidly after intravenous administration and this may lead to an increased drug concentration in the blood, leading to toxic levels. Moreover, many drugs such as anticancer drugs, are hydrophobic in nature and can be subjected to physiological degradation. Controlled DDS can inhibit systemic drug toxicity with a relatively high drug concentration at the site of treatment. The drug(s) can be entrapped in the DDS via physisorption to form polymer-drug dispersions or via chemisorption to form polymer-drug conjugates (12). The advances in polymer-based drug carriers allow the spatiotemporal release of therapeutics in both pulsatile dose delivery products and implanted 316

reservoir systems. Sustained drug release from nanocarriers maintains the plasma drug concentration in the therapeutic window and minimizes the adverse effects of the drug (11).

Figure 1. Schematic representation of different nanocarrier systems for drug delivery.

Polymeric Nanocarriers for Drug Delivery PNPs represent a fascinating class of materials, and they are extensively utilized for biomedical applications owing to their small size, design flexibility based on functionalization, macromolecular synthesis methods and polymer diversity. PNP-based drug carriers have been developed to improve the diagnosis and treatment of a wide range of diseases including cardiovascular diseases, respiratory diseases, neurodegenerative diseases, viral infections and cancer. PNPs are defined as colloidal particles, 3-200 nm in size, designed for drug delivery applications (3). Due to their large surface area, a pharmaceutical agent of interest can be adsorbed onto the surface of the PNPs, or encapsulated or conjugated to the PNPs permitting a high loading capacity. PNPs can host a variety of active constituents including chemotherapeutics, contrast agents, and DNA as well as proteins for cancer and gene delivery. The main purpose of any DDS is to deliver the active agent (drug) to the intended pathological site in a selective manner and in a sufficient amount to increase efficacy and reduce cytotoxicity towards peripheral healthy tissues. Two strategies of drug targeting 317

can be applied to the site of action, active or passive targeting. Active targeting requires cell-specific ligands, attached covalently to the surface of engineered PNPs. Various target ligands can be attached to the drug carrier such as antibodies, peptides, folate, sugar moieties and other ligands (13). For the potential targeting of liver cell carcinoma, glucosamine was conjugated with paclitaxel (PTX)-loaded-poly (γ-glutamic-co-lactic) acid NPs (14). Folate-modified PNPs are widely investigated as agents to target tumors, because the folate receptor is highly overexpressed in various types of human tumor cells (15–17). In passive targeting, tumor tissue is characterized by impaired lymphatic drainage facilitating the penetration of PNPs through leaky blood vessels and accumulation in tumor tissue, a phenomenon known as the EPR effect (7, 18). The decoration of PNPs with poly(ethylene glycol) (PEG) allows stealth effect of NPs and improves NPs clearance from the circulation. PEGylation involves coating the surface of NPs with PEG so that ethylene glycol (EG) units interact with water molecules by hydrogen bonds to form a hydrating layer (19). This hydrating layer in turn reduces protein adsorption and subsequent clearance by MPS. PEG is approved by FDA for food, cosmetics and pharmaceutical purposes. It is a hydrophilic non-toxic biocompatible polymer soluble in a wide range of solvents and lacks immunigenity (20). PEG is also a thermo-responsive polymer with a lower critical solution temperature (LCST) above 100 °C (21, 22). Thermo-responsive polymers display a phase transition in aqueous solution, switching from a hydrophilic to a hydrophobic character upon heating, due to inter- and intramolecular hydrogen bonds between water molecules and polymer chains. The high LCST of linear PEG is not appropriate for drug delivery applications, for this reason many research groups have studied copolymerization of PEG with hydrophobic monomers to reduce its LCST and make it more suitable for drug delivery applications. Poly(ethylene oxide-stat-propylene oxide) block copolymer has been investigated as a thermo-responsive polymer (23). Although, poly(N-isopropyl acrylamide) (PNIPAM) is a popular thermo-responsive polymer with a LCST of · 32°C extensively used in drug delivery applications, including hyperthermia (24). However, it has been reported that PNIPAM exhibits cytotoxicity at a physiological temperature, which could be due to the presence of unreacted monomers (25). In addition, secondary amide groups distributed pendant to the back-bone of PNIPAM form undesired hydrogen bonds with proteins (26). During the last decades, poly (oligo(ethylene glycol)) methacrylate (POEGMA)s have been explored as an alternative to PNIPAM (27). This non-linear PEG analogue exhibits thermo-responsive properties in water or physiological media. It has a carbon-carbon backbone connected to EG segments up to 85 % of its weight, and it therefore, displays both water solubility and biocompatibility (27–29). The LCST of these polymers can be tuned in the range of 26-90 °C depending on the length of EG units in the side chain. Several research groups have studied the polymerization and copolymerization of oligo(ethylene glycol) methyl ether methacrylate (OEGMA) monomer using different polymerization methods including atom transfer radical polymerization (ATRP) and reversible addition-fragmentation chain transfer (RAFT). Lutz and his coworkers studied the random copolymerization of di (ethylene glycol) methyl ether methacrylate (DEGMA) and OEGMA via ATRP with tunable LCST (30, 318

31). A library of homopolymers and copolymers of OEGMA with methacrylic acid (MAA) was investigated using RAFT polymerization with adjusted LCST (32). We conclude that the precise design of PNP-based drug carriers, in terms of particle size and chemistry, will influence their biological safety, fate in the body and drug pharmacokinetics. The chemistry of PNPs and their hosted payloads have a great impact on their biodegradability, tissue biodistribution and cellular fate. Different classes of polymers are able to form PNPs-based drug delivery systems, including biodegradable polymers, amphiphilic copolymers, dendrimers and hyperbranched polymers. In the following part of the chapter we discuss biodegradable polymers and amphiphilic copolymers. In the next section, we shall consider hyperbranched polymers forming unimolecular micelles. Biodegradable Polymers Biodegradable polymers have emerged as promising candidates for therapeutic devices such as temporary prostheses, three-dimensional porous structures as scaffolds for tissue engineering and drug delivery systems. The principal features of a biodegradable polymer is that it degrades in vitro and in vivo either into products that are normal metabolites in the body or into products that can be eliminated from the body with or without further metabolic transformation (33). Collagen and gelatin are natural biodegradable polymers that have been studied most as drug delivery system due to their nontoxicity, biocompatibility and ease of extraction and purification. However, collagen has shown an immunogenic response which may limit its implementation as a drug carrier. Moreover, most natural polymers are water-soluble and must be cross-linked to produce a water-insoluble polymer network, a DDS-scaffold. In a gelatin-based drug delivery system, gelatin is cross-linked with glutaraldehyde. However, glutaraldehyde as a non-specific cross-linking agent interferes with protein drugs such as interferon (IFN-α) leading to inactivation of these drugs and loss of efficacy (12, 34). Synthetic biodegradable polymers offer remarkable advantages over natural polymers owing to their flexibility, having a wide spectrum of properties with good reproducibility. Fine control of the degradation rate of these polymers is possible by varying their structure (35). Synthetic biodegradable polymers include poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(ε-caprolactone) (PCL), polyanhydride, polyphosphoesters and polyphosphazenes (35, 36). Among these, PLA, PGA, and their copolymer poly(D, L-lactide-co-glycolide) (PLGA) are the most well-defined and most frequently used polymers in drug delivery. The first FDA-approved drug delivery system, Lupron® Depot, based on a biodegradable polymer was released in 1989. Lupron® Depot consists of PLGA microspheres encapsulating leuprolide for the treatment of prostate cancer (37). The drug release rate from the biodegradable formulation is controlled by the biodegradation of PLGA, allowing a sustained drug release profile and thus minimizing toxic side effects and increase patient compliance. Khan et al. (38) synthesized PLGA NPs and encapsulated ormeloxifene for pancreatic carcinoma therapy. They showed that the prepared drug formulation, having a particle size 319

of 100 nm, displays significant anti-cancer activity toward pancreatic cancer cells with high accumulation in the cytosol and mitochondria. In another study, PLGA nanocarrier co-encapsulated two or more active agents, doxorubicin (DOX) and epidermal growth factor receptor (EGFR) siRNA in which the angiopep-2 (ANG) was conjugated for glioma therapy. They found that both DOX and siRNA were released in a controlled manner from PLGA NPs. The main role of ANG is to permeate the blood brain barrier that is known to restrict the passage of drugs to the brain (39). Theranostic is a wide medical concept in which both diagnosis and therapy are achieved using a single nanocarrier system. Schleich and his coworkers (40) developed nano-theranostic PLGA particles loaded with PTX and contrast agent to achieve simultaneous molecular imaging, drug delivery and real-time monitoring of the therapeutic response. A biodegradable amphiphilic linear-dendritic hybrid polymers based on PEG as the hydrophilic segments and a dendron branched PCL as the hydrophobic segments are synthesized. Theses polymers can form micelles with average size about 100 nm which were used to encapsulate DOX at loading efficiency up to 22 %. The cellular toxicity of these micelles were evaluated on two breast cancer cell lines and primary human macrophages. The pristine micelles showed no cell death below 35 µg/ml, while the Dox-loaded micelles displayed a significant decreases in the cell viability (41). Amphiphilic Copolymers Amphiphilic block copolymers can form polymeric micelles, leading to drug-loading systems with unique characteristics for drug delivery applications such as improved stability, enhanced drug solubility and bioavailability. Amphiphilic block copolymers can self-assemble above their critical micelle concentration (CMC) in aqueous solution forming core-shell micelle architectures. This created an integrated multifunctional system consisting of hydrophobic inner core to encapsulate water-insoluble drugs with a hydrophilic outer shell protecting the micelle and drug from the surrounding environment and providing colloidal stability. Stealth properties of polymeric micelles originate from the PEG forming corona giving a prolonged residence time in the blood circulation. The micelle-forming process starts when a number of amphiphilic polymer chains, amphiphiles or unimers, gather at the air/water interface. At a certain unimer concentration, the CMC, they associate in an entropy-driven process, by releasing water molecules to the bulk of the aqueous phase (42). The equilibrium between micelles and their individual amphiphiles is crucial for micellar stability. Upon dilution or other changes, thermodynamics may end to a disassembly of the micelles, resulting in a premature release of their cargo. A facile approach to increase the micellar stability in a highly diluted environment is to covalently connect the blocks that form either the core or the shell of micelles by crosslinking (43–45). Core cross-linking seems to be the more promising approach to preserve the loaded cargo in the core of the micelles. The core can be cross-linked by the addition of a difunctional reagent to a reactive core-forming block. Several cross-linkers are highly reactive to functional groups in the block copolymer such as diamine (46), bis-succinimidyl crosslinker (47) and bis-benzophenone photocross-linker (48). On the other 320

hand, the amphiphilic block copolymers produced by reversible-deactivation radical polymerization (RDRP) of monomers such as acrylates or methacrylates can self-assemble in aqueous solution and form micelles. These micelles can be stabilized and core cross-linked by chain extension in the presence of a difunctional monomer possibly containing acid-labile or -reducible disulfide functions for pH- or reduction-controlled drug release (49–51). In this context, micelles containing disulfide bonds in the core are exploited for intracellular drug delivery, since they are accessible to chemical cleavage under reduction conditions such as glutathione. Glutathione is found in cell cytosol and other organelles at a concentration of 0.5-10 mM, while its extracellular amount in plasma is 2-20 µM and that limits the dissociation of the disulfide bond (52, 53). Preservation of the micelles maintains the colloidal stability during blood circulation. When drug loaded-micelles penetrate a cell, they are subjected to intracellular disassembly and release the drug in a controlled fashion. The Lee and Lam research groups (54, 55) investigated the loading of anticancer drugs such as docetaxel and PTX into disulfide-cross-linked micelles; and showed that cross-linked micelles improve the therapeutic efficacy and enhance the tumor specificity in tumor-bearing mice compared to noncross-linked micelles. Wei et al. (56) developed reduction-sensitive reversibly core cross-linked micelles based on poly(ethylene glycol)-b-poly(N-2-hydroxypropyl methacrylamide)-lipoic acid (PEG-b-PHPMA-LA) conjugates. The PEG-b-PHPMA copolymer was obtained under control of RAFT, it was prone to conjugate with lipoic acid via esterification of the hydroxyl groups in the copolymer forming amphiphilic PEG-b-PHPMA-LA polymer. The micelles showed a high DOX loading efficiency of 90 %. However, the drug release profile at 37 °C showed that only 23 % of the DOX released from the cross-linked micelles in 12 hours while about 87 % of the DOX was released after treatment with 10 mM dithiothreitol (DTT) showing high anti-tumor activity in HeLa and HepG2 cells. In another study, a bioreducible moiety was introduced via click chemistry to produce cross-linked micelles in which azide-functionalized biodegradable polymeric micelles encapsulated methotrexate for breast cancer therapy. The micellar cores were cross-linked at their pendent azide groups with disulfide-containing bisalkyne. The drug-loaded-cross-linked micelles inhibited the metabolic activity of human breast cancer cells (MCF-7) more than drug-loaded-uncross-linked micelles or the free drug (57).

Unimolecular Micelles: Synthetic Aspects In the previous section conventional amphiphilic block copolymers forming-micelles, known as multimolecular micelles were discussed. There is always an equilibrium between the micellar form and their individual molecules, with a constant and rapid chain exchange, when they are present in concentrations above the CMC. If they are diluted to a concentration below the CMC, a change in solution conditions such as temperature, pH, ionic strength or shear forces results in the dissociation of micelles into free polymeric chains which may lead to undesired drug release and systemic toxicity (Figure 2). Unimolecular micelles have therefore attracted great interest owing to their unique properties where a 321

single molecular architecture exhibits micelle-like properties, and it is possible to stabilize the construct individually in extremely dilute conditions. Unimolecular micelles were first proposed by Newkome et al. (58) as an alternative way of designing stable polymeric micelles. Significant efforts have been devoted to design macromolecular architectures that mimic the spherical 3D structure typical for micelles, and are covalently linked to avoid any dynamic nature. Some parameters have been shown to be crucial to accomplish individual stabilization including balanced amphiphilicity, high molecular weight and an extensive number of hydrophilic, stabilizing arms. Two main strategies have been suggested to design unimolecular core-shell type micelles, which can be divided into the “arm-first” and the “core -first” approaches (59).

Figure 2. Different behaviors of unimolecular and multimolecular micelles upon dilution. (60). Adapted with permission from reference (60). Copyright © 2015 Elsevier Inc.

Arm-First In the “arm-first” approach, linear amphiphilic AB diblock copolymers are linked covalently to form the characteristic morphology of a micelle. This can be accomplished by taking advantage of the spontaneous self-assembly of amphiphilic and low molecular weight polymers to form conventional micelles, followed by post-crosslinking of the hydrophilic (A) or hydrophobic (B) domains to stabilize the assembly (43, 44, 61), as illustrated in Figure 3A. Another concept studied in the last decade is the formation of linear hydrophilic macroinitiators followed by a cross-linking during the polymerization of the hydrophobic domain in the presence of a small amount of bifunctional monomer (Figure 3B). In this way shape-directed monomer-to-particle synthesis has been shown to produce unimolecular micelles (62). Hawkett et al. (63) and Charleux et al. (64) utilized the polymerization-induced self-assembly (PISA) technique to 322

obtain NPs. In this technique, a hydrophilic macroinitiator is used to polymerize hydrophobic monomers in a surfactant-free emulsion polymerization typically conducted in water. Once the hydrophobic block is long enough, the diblock copolymer self-assembles into a well-defined NPs with a hydrophilic corona and a hydrophobic core. Davis et al. (65) demonstrated the use of PISA to construct NPs intended for drug delivery, where the hydrophobic core was subsequently cross-linked. Core-First In the “core-first” approach, the focus has been on structural hybrids based on dendritic segments and linear polymers (Figure 3C). Dendritic polymers, with their globular structure, their large number of functional groups in the periphery and their highly branched structure, are extremely suitable for subsequent chainextension by linear segments enabling steric stabilization as unimolecular micelles (66, 67). However, the tedious and costly synthesis of dendrimers as well as their limitation in size (typically ≤ 10 nm), has diverted attention to their less perfectly hyperbranched analogues (68, 69).

Figure 3. Schematic illustration of the different strategies to synthesize unimolecular micelles: A) self-assembly of amphiphilic linear block-co-polymers and subsequent crosslinking of core or shell (arm first), B) core crosslinking by using a small amount of bifunctional monomer (arm-first), C) chain-extension of dendritic polymers with linear segments to form dendritic-linear hybrids (core-first). 323

Amphiphilic Hyperbranched Polymers Hyperbranched polymers are highly branched macromolecules with 3D dendritic architectures, and these polymers have been used for the preparation of unimolecular micelles. They have attracted significant attention for the design of NPs for use in therapeutic delivery systems. Hyperbranched polymers are synthesized via the step-growth polymerization of ABx-type monomers (x ≥ 2) which was postulated by Flory (70) in 1952. Another approach to produce hyperbranched polymers developed by Fréchet et al. (71) is the radical polymerization of AB*-type monomers (A is a monomer vinyl group and B* is an initiator fragment), referred as self-condensing vinyl polymerization (SCVP). In both cases the polymer produced is characterized by a broad molecular weight distribution; i.e., high polydispersity (72, 73). The immense progress of RDRP now renders a unique structural control over the hyperbranched polymer structure possible, thereby narrowing down the molecular weight distribution.

Figure 4. Schematic illustration of the synthetic route to Hyperbranched Dendritic-Linear Polymers via SCV(C)P-ATRP and followed by ATRP polymerization techniques (74). Reprinted with permission from reference (74). Copyright © 2014 American Chemical Society. 324

The introduction of the SCVP technique also makes it possible to employ RDRP to synthesize hyperbranched polymers in a more controlled fashion (72). This gives great freedom of design, where the polymer composition and architecture can be readily tailored by a careful choice of starting materials. Previously, our group has, among others, developed the synthesis of high molecular weight amphiphilic polymers based on a hydrophobic hyperbranched core synthesized through a combination of ATRP and self-condensing vinyl homo- and copolymerization (SCV(C)P) (74) (Figure 4). This strategy makes it possible to create highly sophisticated hyperbranched dendritic-linear polymers (HBDLPs) via a straightforward two-step protocol, and this strategy makes it possible to finely-tune the inherent features of the unimolecular micelle such as; size, stealth, targeting, drug and selective delivery. In another study, an amphiphilic hyperbranched terpolymer was synthesized through SCVP of OEGMA and methyl methacrylate (MMA) monomers in the presence of a vinyl-functionalized chain transfer agent (C12-raft) under the control of RAFT polymerization. The hyperbranched polymer produced can encapsulate a hydrophobic molecule, epoxiconazol, which could make a complex with MMA segments forming unimolecular micelles in aqueous solution (75).

Figure 5. Schematic illustration of the synthetic route to unimolecular micelles with (a) backbone-cleavable disulfide bonds and (b) azide-functional cleavable pendant disulfide bonds (59). Reprinted with permission from reference (59). Copyright © 2015 American Chemical Society. 325

Smart Unimolecular Micelles Responsive, or smart, polymers have a great applicability in the field of controlled and self-regulated drug delivery for cancer therapy. Responsive polymers can undergo a rapid change in properties in response to endogenous stimuli, including redox potential, pH and enzymes, or to exogenous stimuli such as temperature, light and ultrasound (76). Unimolecular micelles based on these polymers can designed for delivery systems that closely resemble the physiological process as in cancer tissue and optimize drug release according to the physiological needs. Porsch et al. (59) showed how disulfide bonds can be conveniently introduced into HBDLPs using the SCVP technique that has previously been prepared by the same group (74). The disulfide bonds were attached to the backbone of HBDLPs to enable biodegradation or to the pendant groups to allow the triggered release of the cargo as shown in Figure 5. They showed successfully encapsulated chemotherapeutic DOX in the core of the micelles displaying a diffusion-controlled release profile in a reductive medium, including glutathione and dithiothreitol. In another study, pH-sensitive bonds were introduced to an amphiphilic hyperbranched polymer consisting of a hydrophobic hyperbranched polyacetal attached to hydrophilic PEG by hydrazine bonds. The micelles formed showed a high stability in pH 7.4 buffer solution but were quite fragile in a pH 5.0 buffer. The micelles loaded with DOX gave a faster drug release in pH 5.0 buffer than in a physiological pH solution (77).

Conclusions The rapid developments in nanotechnology and materials science are making it possible to improve the quality of human life. Polymeric nanoparticles are being used as versatile nanocarriers for a wide range of drugs to treat lethal diseases such as cancer. Essential characteristics of these polymeric drug carriers include size, design flexibility, functionality and biocompatibility, which are critical when designing macromolecules for use in drug delivery devices. Drug-nanocarriers based on polymeric platforms include biodegradable polymers, amphiphilic copolymers, and hyperbranched polymers that can form unimolecular micelles. Unimolecular micelles are promising for drug delivery systems due to their high stability. Their structures, properties and methods of preparation are also briefly discussed.

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