Review

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Therapeutic Delivery

Polymer-based vehicles for therapeutic peptide delivery

During the last decades increasing attention has been paid to peptides as potential therapeutics. However, clinical applications of peptide drugs suffer from susceptibility to degradation, rather short circulation half-life, limited ability to cross physiological barriers and potential immunogenicity. These challenges can be addressed by using polymeric materials as peptide delivery systems, owing to their versatile structures and properties. A number of polymer-based vehicles have been developed to stabilize the peptides and to control their release rates. Unfortunately, no single polymer or formulation strategy has been considered ideal for all types of peptide drugs. In this review, currently used and potential polymer-based systems for the peptide delivery will be discussed.

Peptides with a length of 50 or less amino acids truly fill a niche between small molecules and large proteins in terms of size and structure complexity [1] . Peptide drugs offer certain advantages over other drug candidates. As compared with the small molecules, they demonstrate higher selectivity and efficacy, along with lower risks of systemic toxicity with amino acids as degradation products  [2,3] ; while, peptides typically exhibit higher stability, better tissue penetration and lower immunogenicity than proteins and antibodies [4] . Therefore, the interest in developing new peptide therapeutics is steadily growing. The commercial market of therapeutic peptides emerged in the 1970s when Novartis launched Lypressin, a vasopressin analog [5] . As of 2015, US FDA has approved more than 60 peptide drugs which are currently available on the market, and this is expected to grow significantly, with approximately 140 peptide drugs undergoing clinical trials now and more than 500 therapeutic peptides in preclinical development stage [6,7] . Previously, the applicability of peptide drugs was mainly limited to metabolic disorders (diabetes, obesity, etc.) and oncology, but now their therapeutic use has been extended to a wide assortment of medical disorders

10.4155/tde.15.71 © 2015 Future Science Ltd

Jinjin Zhang‡,1, Swapnil S Desale‡,1 & Tatiana K Bronich*,1 1 Department of Pharmaceutical Sciences & Center for Drug Delivery & Nanomedicine, University of Nebraska Medical Center, Omaha, NE 68198, USA *Author for correspondence: Tel.: +1 402 559 9351 Fax: +1 402 559 9365 [email protected] ‡ Authors contributed equally

including rare diseases, infection, inflammation, cardiovascular and neurodegenerative diseases [8–10] . However, therapeutic peptides pose multiple challenges for their delivery to the site of action  [11] . First, the chemical and p­hysical instability of therapeutic peptides remain the major obstacles [12] . Chemical instability is defined as the modification of the peptide by formation or cleavage of covalent bonds, generating new chemical entities, via hydrolysis, oxidation, racemization, isomerization, deamidation, disulfide exchange and β-elimination. Physical instability normally refers to conformation change (dimerization and further aggregation), which could potentially result in denaturation, adsorption to surfaces, aggregation, precipitation, etc. [13] . Second, the successful delivery of peptide drugs to the therapeutic target also suffers from the rapid renal clearance, reticuloendothelial system recognition and off-target binding. Third, restricted membrane permeability and degradation in the gastrointestinal tract lead to poor oral bioavailability of most peptide drugs, which explains why the most peptide therapeutics (∼75%) are injectables [14,15] . In an attempt to overcome these restraints, polymeric materials have been

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Key terms Physical instability: Peptide conformation changes including aggregation, precipitation and adsorption. Chemical instability: The peptide covalent modification via bond formation or cleavage, including hydrolysis, deamidation, oxidation, disulfide exchange, β-elimination and racemization. Peptide PEGylation: Covalent attachment of polyethylene glycol (PEG) polymer chains to peptides

extensively exploited during the past decade for peptide delivery. As was first demonstrated by Langer and Folkman, the controlled release of proteins was proved to be feasible for more than 100 days via their incorporation into synthetic polymers [16] . Since then, various polymer-based platforms have been explored as delivery vehicles for peptides and proteins including chemical modification (e.g., peptide PEGylation)  [17] , nano-/micro-particulate encapsulation  [18–20] , etc. One classical example is insulin, which perhaps is the most studied peptide drug. To achieve the noninvasive delivery, a variety of polymers (e.g., chitosan, dextran, poly[glutamic acid], hyaluronic acid, poly[lactic acid], poly[lactide-coglycolic acid]), polycaprolactone, acrylic polymers and polyallylamine have been exploited to reduce the enzymatic degradation and increase the intestinal permeability of insulin [21–23] . Polymeric materials with highly tunable structures and physicochemical properties offer multiple remarkable advantages for the delivery of peptides, including improved peptide stability, controlled and sustained release and improved pharmacokinetics/biodistribution profiles [24–27] . Overall, polymer-based vehicles are attractive carriers for peptide delivery (Figure 1) . Aim of this review is to provide an overview of different types of polymer-based nanocarriers that have been developed for peptide delivery. Particulate systems Numerous studies have shown that the micro-/nano-particulate systems protect peptides from enzymatic degradation, prolong the plasma half-life and reduce the dosing frequency. As opposed to the polymer–peptide conjugate, they do not require the chemical modification of peptide structure for encapsulation. To select the appropriate particulate system, the properties of peptide drugs such as their solubility, size and structure need to be considered. The administration route of peptide drugs also dictates the choice of the type of the carriers. This part of the review covers several of the polymeric particulate systems and their application for peptide delivery via different administration pathways (Figure 2) .

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PLGA-based particles

Poly(d,l-lactic-co-glycolic acid) (PLGA)-based nano-/ micro-particles perhaps are the most extensively explored systems for peptides delivery as these polymers are FDA approved, biodegradable and biocompatible (Figure 1A)  [28–30] . The degradation of PLGA copolymers occurs mainly through the hydrolysis of ester bonds, and the rate of degradation is known to be dependent on the molar ratio of the lactic acid to glycolic acid, polymer molecular weight and end groups as well as particle porosity [31–34] . Therefore, the release kinetics of peptides from PLGA particles can be tuned by adjusting these parameters [35] . On the other side, initial rapid release, incomplete release of native proteins/peptides due to partial degradation and irreversible adsorption of proteins/peptides to the polymer matrix have been recognized as major issues with this system [33,36,37] . The hydrolysis-induced acidic microenvironment during the storage impairs the stability of the encapsulated peptides [24] , and this could be addressed by blending with other stabilizer agents including polymers (e.g., cellulose, alginate, polyethylenimine, poly[ethylene glycol] [PEG]) or antacid salts (e.g., Mg(OH)2, MgCO3, ZnCO3) during encapsulation  [38,39] . Another limitation is the recognition of injected hydrophobic PLGA particles by the reticuloendothelial system and their elimination from the blood stream. Hydrophilic polymers (e.g., PEG) can be introduced to the particle surface to confer stealthiness and extended circulation half-life. Furthermore, the modification of PLGA particles with ligands capable of recognition of cell-specific surface receptors can be utilized for targeted delivery of bioactive molecules to specific organs [40,41] . In general, three approaches are used to produce peptides-loaded PLGA particles: double emulsion (w/o/w) technique, phase separation methods and spray drying methods  [42] . Notably, peptide drugs are less affected by harsh manufacturing conditions compared with large proteins. Various peptide hormones have been loaded in PLGA microparticles, which act as a depot displaying sustained release over an extended period of time, thereby reducing the dose and the frequency of administration. This is evidenced by numerous PLGA microparticulate products on the market (Table 1) [43] , which are usually administered subcutaneously or intramuscularly. Lupron Depot® is the first approved injectable PLGA microparticles for sustained delivery of leuprolide. To treat advanced prostrate cancer, this product is available in 7.5 mg (1-month release), 22.5 mg (3-month release), 30 mg (4-month release) and 45 mg (6-month release). The 14 kDa PLGA (lactic acid to glycolic acid ratio 75:25) is used for the 1-month release product, and the release period is further

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Polymer-based vehicles for therapeutic peptide delivery 

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Injectable implants

Increased stability Controlled release longer half-life Decreased toxicity Targeted delivery

Improved pharmacokinetics and biodistribution profiles Increased therapeutic activity

Polymer–peptide conjugates Figure 1. Overview of polymer-based vehicles (particulate systems, injectable implants, and polymer–peptide conjugates) for peptide delivery and their advantages.

extended to 3, 4 and 6 months by using the poly(lactic acid) with the molecular weight of 12–18 kDa [44] . The product consists of lyophilized microparticles, which are reconstituted and administered as a single intramuscular injection. The diluent for reconstitution is composed of suspending agent (carboxymethylcellulose sodium), along with d-mannitol, polysorbate 80, water for injection USP and acetic acid to control the pH [45] . Compared with microparticles, nanosized PLGA particles display relatively higher cellular uptake and preferential penetration into tissues (e.g., liver, colon, tumor vasculature) [46,47] , and are typically administered through the intravenous route. It has been demonstrated that PLGA nanoparticles can act as a reservoir for antigenic peptides or their combination with adjuvants and, therefore, are very attractive platforms as vaccines or cancer immunotherapeutics with longlasting effect [48,49] . Multiple advantages have been reported, including enhanced immune response due to efficient uptake by antigen-presenting cells, sustained release of entrapped antigenic peptide, simultaneous delivery of antigen and adjuvant, etc. [47] . Zhang et al. prepared the PLGA nanoparticles carrying murine melanoma antigenic peptide and toll-like receptor 4 agonist as adjuvant for cancer immunotherapy  [50] . As compared with the antigen peptide mixed with Freund’s adjuvant, the intradermal injection of

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PLGA nanoparticle-based formulation into mice led to stronger activation of functional cytotoxic T lymphocytes specific to tumor-associated self-antigens. More importantly, in vivo, vaccination showed significantly delayed melanoma tumor growth in a prophylactic setting. The dendritic cells-targeted delivery of antigenic peptide (830–844 region of tetanus toxoid) can be achieved by coating the PLGA nanoparticles (200 nm in diameter) with the humanized targeting antibody hD1. As a result, the similar immune responses were achieved from targeted nanoparticles at 10–100-fold lower concentrations than non-targeted nanoparticles  [40] . Insulin-loaded PLGA nanoparticles have also been extensively explored for noninvasive routes of insulin administration for treating diabetes [51–53] . For example, after administration of inhalable PLGA-formulated insulin (weight mean diameter of 400 nm) via a sieve-type ultrasonic nebulizer to the lung of guinea pigs, the blood glucose level was reduced significantly and the hypoglycemia was prolonged over 48 h, compared with a nebulized aqueous solution of insulin as a reference (6 h) [54] . This result was attributed to the sustained release as well as the wide spread of insulin in the whole lung. The oral delivery of insulin has been also explored using PLGA nanoparticles as carrier systems that are able to protect the drug from enzymatic degradation and increase its permeability through intestinal epithelium. Various strategies have been

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Figure 2. Particulate carrier systems for peptide delivery. (A) PLGA-based particles. (B) Polymeric capsules. (C) Polymeric micelles. (D) Polyelectrolyte complexes (left) and block ionomer complexes (right) of cationic peptide with anionic polymer. (E) Hydrogels. PEG: Poly(ethylene glycol); PLGA: Poly( d,l-lactic-co-glycolic acid).

investigated for this purpose: the encapsulation efficiency of hydrophilic insulin into hydrophobic PLGA nanoparticles was markedly improved by its complexation with various compounds like lipids and hypromellose phthalate [55–57] ; favorable encapsulation of insulin was observed in star-branched β-cyclodextrinPLGA nanoparticles over linear PEG-PLGA, which led to retention of insulin stability and more sustained release [58] . Polymeric capsules

In contrast to the polymeric nano-/micro-particles described above with a matrix-type structure, polymeric capsules have inner liquid cores (oil or aqueous solution, depending on the synthetic strategy) surrounded by single-/multilayered polymeric shells (Figure 1B) . Generally, polymeric capsules can be fabricated by template-free or template-assisted techniques  [59] . The most frequently applied strategies are the self-assembly of block copolymers (template-free),

and the layer-by-layer technique employing a sacrificial template. The major issue with a template-free method is the size polydispersity of the resulting capsules, and this can be addressed by extrusion, sonication, freeze–thaw cycles, etc. Comparatively, the layer-bylayer technique provides a more precise control over the shell properties. The therapeutic agents can be either entrapped in the inner core of the capsule or absorbed on the surface of polymeric shell using various strategies such as drug loading in preformed capsules by temporarily changing shell permeability; drug pre-encapsulation into sacrificial core template upon capsule formation, which will leave the drug molecules in the inner cavity when template is removed; and incorporating drugs in specific domains of the capsules. In addition, the entrapment can be managed by modulating the core material (oil or aqueous phase) according to peptide properties. The composition and the thickness of polymeric shells determine the capsules permeability and the release rate of entrapped drugs [60,61] . The

Table 1. Peptide-loaded poly( d,l-lactic-co-glycolic acid) microparticles on the market. Trade name

Peptide drug

Indication

Lupron Depot  

Leuprolide acetate

Prostate cancer

1, 3, 4 or 6 months

Abbott

TrelstarTM depot

Triptorelin pamoate Prostate cancer

1 or 6 months

Pfizer

Triptorelin pamoate Prostate cancer

1 month

Ipsen

®

Decapeptyl

Company

Sandostatin LAR depot

Octreotide

Acromegaly

1 month

Novartis

Somatuline® LA

Lanreotide

Acromegaly

2 weeks

Ipsen

Suprecur® MP

Buserelin

Endometriosis

1 month

Aventis

Bydureon

Exenatide

Diabetes

1 week

AstraZeneca

®

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Polymer-based vehicles for therapeutic peptide delivery 

triggered drug release can also be achieved by incorporating stimuli-responsive (e.g., pH-, redox-sensitive, enzyme-cleavable) ‘building blocks’. A wide range of different polymers can be selected to prepare polymeric capsules with desired particle size, composition, morphology, shell thickness and functionality. Nanocapsules composed of bioadhesive polymers have been widely investigated as carriers for transmucosal delivery. Poly(alkylcyanoacrylate)-based (PACA) nanocapsules have been investigated for oral delivery of various peptides including calcitonin, insulin, octreotide and cyclosporine A (CsA) [62,63] . They were prepared by interfacial emulsion polymerization. The peptide encapsulation efficiency as well as release rates were modulated by adjusting the alkyl chain length, molecular weight, monomer concentration and polymerization kinetics [64] . In addition to the protective effect against enzymatic degradation and controlled release of peptide, the bioadhesive properties of PACA nanocapsules facilitate peptide translocation through the intestinal mucosa resulting in improved oral bioavailability [65] . Different chitosan-based nanocapsules were reported to be effective for the oral administration of salmon calcitonin, a potent hypocalcemic agent and allow achieving a significant response (∼30% peak reduction in serum calcium levels), which was maintained for more than 24 h [66] . These type of nanocapsules were also effective for nasal delivery of the calcitonin  [67] . Interestingly, it was demonstrated that chitosan-based nanocapsules fabricated using solvent displacement technique are also useful for loading and delivery a combination of therapeutic peptides, for example they can be loaded with a combination of both hydrophobic peptide such as CsA and the hydrophilic peptide (e.g. insulin, calcitonin) [68] . Polymeric micelles

Polymeric micelles are self-assembled nanoconstructs formed from amphiphilic polymers. They have unique core-shell architectures with hydrophobic polymer chains segregating into the core surrounded by a shell of hydrophilic chains such as PEG (Figure 1C) [69] . Hydrophilic shell inhibits protein binding and opsonization during systemic administration, which allows them to circulate in the blood for extended periods of time by evading the mononuclear phagocytic system. Also, modification of the shell with various ligands using different surface chemistries enable the micelles to be targeted to a specific site. Compared with other drug delivery systems, small size (10–100 nm) and low polydispersity of micelles enable better penetration into various tissues, for example, by extravasation [70] . The inner hydrophobic core as well as the palisade layer allow the encapsulation of hydrophobic or amphipa-

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Review

Key terms Polymeric capsules: The structure with an inner liquid core surrounded by a polymeric shell. Polymeric micelles: Self-assembled nanoconstructs of amphiphilic polymers with a core-shell architecture.

thic peptides [71] . A wide range of hydrophobic blocks have been explored as drug-loading cores. Examples include polyesters, polyanhydrides, poly(amino acids), phospholipids/long-chain fatty acids and polypropylene oxide (in pluronics/poloxamers). An important consideration for drug delivery is the relative stability of polymeric micelles: micelles must be stable enough to retain drug cargo upon administration and remain intact long enough to accumulate in sufficient concentrations at the target site. The thermodynamic tendency for micelles to dissociate is primarily controlled by the length of the hydrophobic block while the kinetic (rate of dissociation) stability depends on many factors, including the size of a hydrophobic block, the mass ratio of hydrophilic to hydrophobic blocks and physical state of the micelle core and can be further reinforced by formation of crosslinks between the polymer chains [72,73] . Polymeric micelles have been shown to overcome many challenges related to therapeutic use of peptides including, but not limited to, susceptibility to proteolytic degradation, aggregation in aqueous media, unfavorable release profile, poor biodistribution and pharmacokinetics, limited therapeutic efficacy in vivo, etc. Micelles formed by PEG-phospholipid conjugate (PEG-DSPE) have been utilized to incorporate various therapeutic peptides with the size of 11–42 residues  [74–76] . For example, the PEG(2 kDa)-DSPE micelles loaded with vasoactive intestinal peptide (VIP) with effective diameter of 15 nm were developed as a potential therapeutic modality for rheumatoid arthritis [77] . Interestingly, the association of VIP with PEG-DSPE micelles led to peptide conformation change from random coil to α-helix, which was shown to be beneficial to its stability in aqueous medium  [78,79] . Following intravenous injection, VIP incorporated into the micelles displayed increased half-life in circulation (10.9 h) compared with free VIP (22.6 min), leading to a 13-fold higher drug accumulation at diseased site. However, it should be pointed out that there are certain limitations of this particular system: the peptide loading was strongly dependent on size, secondary structure and hydrophobicity of the peptide [78] . Zeng et al. designed the cationic micelles prepared from branched polyethylenimine (2 kDa)-stearic acid conjugate (effective diameter of 20–45 nm), which were loaded with melanoma antigen peptide (Trp2) [80] . With cationic polymer micelles alone acting as a robust adjuvant as well as a vector

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Review  Zhang, Desale & Bronich for Trp2 antigen, Trp2-loaded micelles promoted the recruitment of immature dendritic cells as well as their proliferation, maturation and homing to the draining lymph nodes. Immunizing mice with these micelles enhanced Trp2-specific cytotoxic T lymphocyte activity and significantly inhibited tumor growth in mouse model of advanced melanoma. CsA is a cyclic lipophilic endecapeptide exhibiting strong immunosuppressive activity, however, its application suffers from very low water solubility (23 μg/ml). Solubilization of CsA into polymeric micelles based on amphiphilic block copolymers [81,82] and graft copolymers [83] have been explored to achieve clinically relevant CsA concentrations in aqueous media and to improve unfavorable absorption characteristics of CsA. For example, CsA was effectively encapsulated into PEG5k-b-poly(εcaprolactone) (PEG-b-PCL) reaching aqueous levels of more than 1 mg/ml for CsA [84] . Notably, CsA encapsulation in PEG-b-PCL micelles resulted in elevated CsA levels in blood after intravenous administration while limiting the CsA distribution to kidney, liver and spleen. Xiong et al. investigated the feasibility of using poly(lactic acid)-b-Pluronic F127-b-poly(lactic acid) (Mn 29 kDa) micelles for the oral delivery of insulin [85] . According to their study, this micellar formulation was able to maintain a prolonged hypoglycemic effect after oral administration which was attributed to the delayed GI transit of the insulin-loaded micelles due to their small size and strong interactions of PEG shell with the intestinal wall. Polyelectrolyte complexes

As compared with the delivery platforms based on PLGA or polymeric micelles, which are more favorable to hydrophobic molecules, hydrophilic polymers offer some advantages for peptide delivery as most of the peptides are hydrophilic or amphipathic in nature. Therefore, another promising class of delivery vehicles for peptides is the polyelectrolyte complexes (PECs), which are composed of polyelectrolyte polymer and peptide/protein of opposite charge ( Figure 1D, left). The PECs assembly is predominantly driven by the electrostatic interactions. However, other intermolecular interactions including hydrogen bonding, van der Waal forces and hydrophobic interactions may also play a role in the complexation process [86] . PECs are an attractive alternative to conventional peptide/protein carrier systems as they can be formulated with high loading efficiency (up to 100% of peptide) under benign conditions that minimize the chances of peptide damage and aggregation. The formation and stability of PECs is dictated by system pH, ionic strength, molecular parameters of polymers (length, structure, charge distribution and density)

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and peptide/polymer stoichiometric ratios and concentrations  [87] . PECs formation leads to particles in the size range of 200–500 nm. The stability of PECs can be further improved by introducing the additional low molecular weight ionic crosslinking agents (e.g., tripolyphosphate, zinc sulfate, calcium chloride) or the oppositely charged polymers [88–90] . Kissel and co-workers utilized cationic polysaccharide chitosan and its derivatives with the molecular weight higher than 25 kDa for the synthesis of PEC-based carriers for insulin which is negatively charged above its isoelectric point (pI 5.3) [91] . It was demonstrated that chitosan methylation and PEGylation significantly improved the stability of insulin in the PECs. Moreover, the PECs could protect insulin from degradation even at 50°C and in the presence of trypsin. All complexes could be lyophilized without influencing the particle size, complex concentration and stability of insulin. The interchain electrostatic bonds can be further strengthened by adding the third component like anionic polymers (e.g., alginate, dextran sulfate, hyaluronate) [92] or crosslinkers (e.g., tripolyphosphate, carboximethyl-β-cyclodextrin) [93–95] . A very special class of polyion complexes based on doubly hydrophilic copolymers containing ionic and nonionic blocks (‘block ionomers’) was proposed in the mid 90s as promising carrier systems for peptide and proteins [96,97] . Such complexes are spontaneously formed in aqueous solutions upon electrostatic binding of block ionomers with oppositely charged macromolecules and termed ‘polyion complex micelles’, or ‘block ionomer complexes’ (BICs) [97–99] . BICs form core-shell nanoparticles with complex cores of neutralized polyions and proteins and hydrophilic polymer shells ( Figure 1D, right). Ionic block lengths, charge density and ionic strength of solution affect the formation of stable BICs and, therefore, control the amount of the protein that can be incorporated within the micelles  [98,100] . It was demonstrated that proteins/ enzymes incorporated in BIC displayed extended circulation time, are stable against proteolytic degradation in blood and inside cells and preserved their activity [101,102] . Our laboratory has developed and characterized several BIC formulations of an amphipathic α-helical peptide, a positively charged analog of C5A peptide derived from the HCV NS5A protein, with a reported virocidal activity. Anionic PEG (5 kDa)b-poly(l-aspartic acid) or PEG (5 kDa)-b-poly(l-glutamic acid) block copolymers with the anionic segment around 10–20 monomer units were used for peptide immobilization into BIC. The self-assembled antiviral peptide nanocomplexes were ca. 35 nm in size, stable at physiological pH and ionic strength and retained in vitro antiviral activity against HCV and HIV. More-

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Polymer-based vehicles for therapeutic peptide delivery 

over, incorporation of the peptide into BIC attenuated its cytotoxicity associated with the positive charge. We demonstrated that these BIC were able to decrease the viral load in mice transplanted with human lymphocytes and HIV-1-infected without signs of toxicity to the animals [103] . Hydrogels

Hydrogels are generally defined as the 3D and crosslinked polymeric networks that, due to their hydrophilic nature, can absorb large amount of water [104–106]. Based on the size, hydrogels can be in the form of macroscopic networks or confined to smaller dimensions like microgels or nanogels (Figure 1E)  [107] . They represent a distinct class of drug carriers exhibiting high loading capacity, biocompatibility, responsiveness to external factors (e.g., ionic strength, pH and temperature) and potential for controlled release. Besides, the functionalization of the hydrogel particles with targeting ligands can further facilitate their selective accumulation in the target tissue or cells [108,109] . An incorporation of peptides into such hydrophilic systems makes them less susceptible to conformation change and potential aggregation, which tend to happen at the hydrophobic interfaces [110] . A number of hydrogel particulate systems have been synthesized and characterized with respect to proteins/peptides delivery. Among polymers of natural origin, polysaccharides (e.g. chitosan and its derivatives, hyaluronic acid, poly[γ-glutamic acid], pullulan) are most commonly used for the preparation of hydrogels. Synthetic biocompatible polymers such as poly(vinyl alcohol), poly(vinyl pyrrolidone) and poly(N-isopropylacrylamide) have frequently been used in the development of potential delivery systems [111] . Diverse synthetic strategies for hydrogels preparation involving physical or chemical methods of crosslinking have also been elaborated [107,112–113] . Physical self-assembly of hydrogels involves electrostatic interactions [114] , hydrophobic interactions [115] and/or hydrogen bonding of the interactive polymers [116] . The more stable networks can be fabricated using covalent crosslinking. The coupling of reactive functional groups can take place during polymerization of monomers or can be used for crosslinking of macromolecular precursors. The most commonly used covalent crosslinking strategies include, but are not limited to radical polymerization, amide-based crosslinking [117] , photo-induced crosslinking [118] , quaternization of amino groups [119] , thiol-disulfide exchange reaction [120] and click chemistry [121] . The porosity, charge density and other characteristics of the gel particles can be systemically varied using covalent crosslinking. Cleavable linkages and degradable polymer backbones or pendant groups are

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Review

Key terms Polyelectrolyte complexes: The charge-driven self-assembly of oppositely charged polymers and macromolecules. Hydrogels: 3D and crosslinked polymeric networks that can swell in aqueous medium.

often used to make hydrogels degradable and trigger the drug release. The majority of micro-/nano-gels that have been investigated for peptide delivery are based on physically crosslinked polymers [122–124] . The entrapment of peptides into such hydrogels is predominantly driven by electrostatic interactions, and the loading capacity of these carriers depends on hydrogel and peptides net charges and charge densities [125] . Besides, the binding and release of peptides from these carriers are also determined by network properties (crosslinking density, network homogeneity, dimensions), peptide size/length and conformation, hydrophobicity, and so on [105] . The delivery of insulin has been extensively studied using hydrogel particles based on polysaccharides (e.g., pullulan, chitosan, dextrin) [124,126] , poly(methacrylic acid) [127] , poly-γ-glutamic acid [128] , etc. Insulin loading of up to 80% was achieved using positively charged chitosan nanogels prepared based on the ionotropic gelation of chitosan with tripolyphosphate containing poloxamer 188 (1.0% w/v) as the surfactant. [129] . Chitosan gel particles with the size of 250–400 nm enhanced the intestinal absorption of insulin compared with free insulin after oral administration in alloxan-induced diabetic rats that led to a prolonged hypoglycemia effect. Akiyoshi et al. have developed self-assembled nanogels composed of cholesterol-bearing pullulans (CHP) (pullulan [Mw 55 kDa] substituted with 2.1 cholesterol molecules per 100 glucose units) with a particular focus on encapsulation of various proteins and peptides [116,130–132] . They demonstrated that up to 10 insulin molecules can be complexed with CHP nanogels (20–30 nm in diameter). The thermal denaturation and subsequent aggregation of insulin were effectively suppressed upon complexation. The complexed insulin was also significantly protected from enzymatic degradation. However, the decrease in blood glucose levels after intravenous injection of insulin-loaded CHP nanogels in rats was not significantly different from that observed for free insulin that was attributed to the relatively fast release of insulin from the complex [116,133] . These challenges were further addressed by incorporating CHP nanogels into hybrid macrogels [134,135] , wherein the sustained release of insulin was achieved. With insulin immobilized into the nanoaggregates of α,β-poly(N-

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Key term Injectable implants: The liquid formulation composed of biodegradable polymer and drug that, once administered, undergoes solidification in situ to form a depot for controlled release of the drug.

hydroxyethyl)-dl-aspartamide modified with hydrophobic segments (C4), anionic groups (COOH) and cysteine residues, its oral administration was proved to be feasible due to the elevated stability against the gastric degradation [136] . Moreoever, Sonaje et al. reported the insulin-loaded pH-sensitive and self-assembled nanogels, which are composed of the mixture of positively charged chitosan (80 kDa) and negatively charged poly-γ-glutamic acid (60 kDa), showing high encapsulation efficiency of 71.8 ± 1.1% [128,137] . In diabetic rat models, effective penetration of the insulin-loaded nanogels through the mucus layer of intestinal epithelium was observed due to the mucoadhesive property of chitosan. Subsequently, insulin was released from the disintegrated nanogel into the systemic circulation due to pH-dependent dissociation of chitosan/polyγ-glutamic acid complex matrix. Overall, the unique characteristics of hydrogels such as hydrophilicity, tunable size, interior network for the incorporation of various biomolecules and large surface area available for functionalization make them very attractive platform for the drug delivery applications. Injectable implants In terms of parenteral controlled-release formulations, injectable implants are an alternative to microparticles and preformed implants. They are liquid formulations comprised of biodegradable polymers and therapeutic agents. After intramuscular or subcutaneous injection via a syringe, the low-viscosity solution is delivered to the malignant site. Triggered by environ-

mental factors, it subsequently transformed in situ to solid or semisolid drug depot that releases entrapped drugs over an extended period of time [138] . Simultaneously, the polymer matrix begins to degrade and gradually becomes metabolized and excreted, eliminating the need for surgical removal of the implants. Generally, injectable implants are applicable when the diseased site is accessible via injection. As compared with micro-/nano-particulate drug delivery systems, the manufacturing process is less complex and mainly requires dissolution and dispersion. Additionally, lower administration frequency, smaller gauge needle and tissue biocompatibility potentially can improve patient compliance. With many advantages listed above, injectable implants have emerged as a tangible delivery platform for therapeutic peptides [139] . Generally, the mechanisms of implants formation can be divided into three categories: in situ polymer precipitation (phase separation), which can be initiated by solvent removal, temperature change or pH change; in situ crosslinking, wherein polymers form networks by a variety of ways including photoirradiation, chemical crosslinking or physical crosslinking of specific monomers; and in situ solidification, which are either hot melts that solidify upon cooling to physiological temperature or liquid crystalline polymers, which self-assemble in aqueous solutions (Figure 3) [140] . So far, the majority of the marketed products are developed based on in situ polymer precipitation systems. Atrigel® technology is one of the earliest in situ forming depots developed by Dunn et al.  [141] . It contains the biodegradable carrier PLGA, which is hydrophobic and dissolved in water-miscible and biocompatible organic solvent (e.g., N-methyl-2-pyrrolidone). Both PLGA concentration and molecular weight affect the viscosity of the liquid formulation. Upon subcutaneous injection, the water-miscible organic solvent

In situ precipitation

In situ crosslinking

In situ solidification

Figure 3. The injectable implants formation mechanisms. Adapted with permission from [139] .

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Polymer-based vehicles for therapeutic peptide delivery 

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Figure 4. First generation of PEGylation of amine-reactive poly(ethylene glycol) derivatives. PEG: Poly(ethylene glycol).

migrates into surrounding tissues and water penetrates into the organic phase, which results in a fast phase inversion and precipitation of the polymer at the injection site. The properties of organic solvent mainly influence the dynamics of this process. When drugs are blended into the polymer solutions, they also remained entrapped in the polymer matrix upon precipitation and their release from the implant is highly dependent on the PLGA molecular weight, composition and concentration. Based on this platform, injectable implant systems that release leuprolide acetate over 1, 3, 4 or 6 months (Eligard®, Sanofi Aventis, USA) are currently available for the management of advanced prostate cancer. For 1-month release, the formulation is composed of 7.5 mg of leuprolide acetate and PLGA copolymer (lactic acid to glycolic acid ratio 50:50) dissolved in N-methyl-2-pyrrolidone (NMP) (NMP to polymer ratio 1.94). The drug release rate can be further delayed (3, 4 and 6 months) by increasing the lactic acid to glycolic acid ratio and by decreasing the ratio of NMP to polymer [142] . Another marketed product (ReGel®, BTG, UK) is based on low molecular weight PLGA (1.5 kDa)-PEG (1 kDa)-PLGA (1.5 kDa) triblock copolymer. Once injected, it undergoes sol-gel transition at physiological temperature (above lower critical transition temperature) [143] . Notably, this technology does not use organic solvents and may be more suitable for incorporation of peptides/proteins.

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Indeed, injectable implants based on thermosensitive gelation were investigated for delivering various peptide drugs including insulin [144] , GLP-1 [145] , CsA  [146] , etc. For example, the constant release rate of insulin from ReGel® was achieved for over 2 weeks after single subcutaneous injection. However, the complete release was hampered by random peptide aggregation, which can be suppressed by using zinc–insulin complexes [144] . Apart from the commercial products, a lot of efforts have been made to develop in situ-forming injectable implants based on a diverse range of polymers as well as gelling mechanisms. However, some challenges remain to be addressed for their successful clinical translation such as initial burst release and potential peptide instability caused by protein–polymer inter­ actions. Furthermore, control over the drug release and polymer matrix degradation rates is limited due to the highly variable process of in situ implant formation. Polymer–peptide conjugates Polymer–peptide conjugates are hybrid materials that exhibit properties of both the biomolecules and synthetic polymers. Covalent attachment of polymers to peptides/proteins is known to improve their efficacy by increasing lifetimes in vivo and decreasing immunogenicity  [147–149] . The peptide modification points, linkers, modification degree and the conjugation

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O H N

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Figure 5. Second generation of PEGylation with thiol-reactive poly(ethylene glycol) derivatives. PEG: Poly(ethylene glycol).

chemistry are all important design considerations having a dramatic impact on the properties of resulting conjugates and their in vivo performance [147,150] . Recent advances in synthetic polymer chemistry have facilitated the possibility of site-specific conjugation of polymers to the proteins/peptides without compromising their activity. Lysine and cysteine residues of a peptide as well as N- and C-terminus modifications and disulfide bridge insertion are commonly exploited for direct-covalent attachment of the polymers. Among various polymers exploited for attachment to peptides/ proteins, PEG is the most successfully used polymer for development of protein therapeutics. Readers are referred to several excellent and more detailed reviews on various peptide/protein conjugation chemistries including PEGylation [125,149,151–154] . For the purposes of this review, we will highlight the major PEGylation strategies and focus on some examples of PEG–peptide conjugates and their applications, followed by some other polymer–peptide conjugates in preclinical developments. Types of PEGylation chemistries are broadly classified into first generation and second generation  [150] . Examples of first-generation PEG derivatives include amine reactive PEG that can be easily attached to the amino groups of lysine residues of a peptide (Figure 4)  [153] . Most first-generation PEG chemistries used acylation (Figure 4), accompanied by loss of charge. Alkylation, on the other hand, using PEG-aldehyde maintains a positively charged amine. Unfortunately, these PEGylation methods have lacked the ability to produce pure monofunctional PEG derivatives [154] . In addition, it was shown that unstable bonds formed between the biomolecule and PEG can lead to degradation of the conjugate during manufacturing and injection [155] . Among second-generation PEGylation strategies, thiol-reactive PEGs have been explored most extensively due to more selective polymer attachment without interfering with the biological

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function of the peptide, while maximally decreasing its immunogenicity [156–158] . PEG-maleimide, PEGvinylsulfone, PEG-iodoacetamide and PEG orthopyridyl disulfide are thiol reactive PEGs that have been used for cysteine modification (Figure 5) . The enzymatic (e.g., glucose oxidase) or chemical (e.g., sodium periodate) oxidation of N-terminal serine or threonine residues generates reactive aldehyde groups, which can be further conjugated with PEG hydrazide or amine derivatives [148] . Another popular second-generation PEG is the heterobifunctional PEG having two different terminal groups, which can be used as a spacer to link two functional groups, or to introduce targeting ligands  [150] . In addition to the linear PEG, branched polymers (Figure 6) have also been reported and proven useful for protein modification (e.g. PEG-IFN α-2a, Pegasys®) [159,160] . Collectively, PEGylation technology has proven its pharmacological advantages and acceptability, but the technology still lags in providing a commercially attractive, generic process to produce highly specific PEGylated therapeutic products with a high yield. With growing interest from emerging biotechnology, there is great scientific and commercial interest in improving present methodologies and in introducing innovative PEGylation strategies for peptide modification. PEGylated peptides have already entered the clinic for the treatment of a range of diseases. Somavert (pegvisomant) is a PEGylated HGH antagonist marketed by Pfizer to treat acromegaly, a rare hormonal disease, in which excessive IGF-1 causes soft-tissue enlargement  [161] . Pegvisomant consists of five linear chains of 5 kDa PEG covalently linked to peptide through lysine and phenylalanine residues. Omontys (peginesatide), developed by Affymax and Takeda, is an erythropoietic agent, a functional analog of EPO. It was approved by the US FDA for treatment of anemia associated with chronic kidney disease in adult

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Polymer-based vehicles for therapeutic peptide delivery 

patients on dialysis [162] . Two 20 kDa PEG are conjugated to the serine terminals of peptide. Clinical trials have demonstrated that peginesatide, administered monthly, was as effective as epoetin, administered one to three-times per week, in maintaining hemoglobin levels in patients undergoing hemodialysis. A number of PEGylated peptides are in clinical trials (Table 2) . Peglispro is insulin PEGylated with the 20 kDa polymer at lysine B28 via a urethane linkage developed by Lilly Research Laboratories. In a recently completed Phase III clinical trial, Peglispro demonstrated superiority in maintaining the blood glucose level compared with Lantus (a commercial insulin analogue) in patients with Type 1 diabetes [163] . Peglispro is currently under Phase III development for the treatment of both Type 1 and Type 2 diabetes. CBX129801, developed by Cebix Inc., is a PEG-conjugate of human C-peptide that can be administered subcutaneously once weekly. Currently it is in Phase II clinical trial for the treatment of Type 1 diabetes patients with mild-tomoderate diabetic peripheral neuropathy. Thymalfasin pegol developed by Jiangsu Hansoh Pharmaceutical is a PEGylated Thymalfasin for the treatment of Hepatitis C, which is currently in Phase III clinical trial in China. These examples clearly demonstrate the usefulness of PEGylation in the improvement of therapeutic value of peptide drugs. Apart from PEGylation, there have been other attempts to conjugate polymers to the therapeutic peptides in the preclinical development stage [164–167] . The extended circulation time and the sustained release of the cytotoxic EGFR-lytic peptide was achieved after its conjugation to the carboxymethyl dextran (15–20 kDa) through a formation of disulfide bond in pancreatic cancer model [165] . Yamamoto et al. demonstrated the 15-fold increased plasma half-life of YIGSR peptide (inhibitor of adhesion and invasion of tumor cells to extracellular matrix) when it was conjugated to poly(vinylpyrrolidone) (PVP) through an amide bond  [166] . In another study, YIGSR was conjugated to poly(styrene-co-maleic anhydride) which led to

Review

the decrease of degradation and increase of antimetastatic effect on lung metastasis of B16–BL6 melanoma cells  [167] . Conjugation was achieved through anhydride group of styrene-co-maleic anhydride and free amine group of peptide. Conclusion The therapeutic potential and applicability of peptide drugs is now well established. Its successful bench-tobedside translation can further be achieved with the help of diverse delivery systems. Among all, the polymer-based vehicles as reviewed here have demonstrated excellent potential for the delivery of therapeutic peptides both in preclinical and clinical settings. The feasibility to control the composition, architecture, and physicochemical properties make polymeric materials extremely promising approach for the peptide delivery. This review is our attempt to highlight several important and widely-used polymer-based carriers for the efficient delivery of peptide drugs. Due to the tremendous advances in biotechnology field, more peptide drugs will be introduced in the coming years and polymeric materials with well-defined characteristics are anticipated to meet its delivery need. Future perspective Notably, each of these polymer-based vehicles has their own pros and cons, and their applicability for peptide delivery is determined by the nature of peptides. Until now, the injectable sustained-release formulations based on biodegradable polymer microparticles or implants remain a major part of the peptide-based products, however, many efforts have been dedicated to develop noninvasive delivery systems. Therefore, we expect that in the future there will be more products on the market for oral, transdermal or nasal administration. Moreover, the recent advances in polymer chemistry have allowed the development of a diverse range of nanomaterials of various sizes, shapes, surface chemistries and targeting properties, which offer exciting prospects for novel formulations, providing noninvasive delivery, stabilization

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Figure 6. Two examples of branched poly(ethylene glycol)s (PEG2). PEG: Poly(ethylene glycol).

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Table 2. Representative PEGylated peptides on market or in clinical trials. Name

Description

Status

Company

Somavert (pegvisomant)

PEGylated HGH antagonist

Approved

Pfizer

Omontys (peginesatide)

PEGylated EPO 

Approved

Affymax and Takeda

Peglispro

PEGylated insulin

Phase III

Eli Lilly

CBX129801

PEGylated human C-peptide

Phase II

Cebix Inc.

Thymalfasin pegol

PEGylated thymalfasin

Phase III

Jiangsu Hansoh Pharmaceutical

and immunogenicity reduction for therapeutic peptides. However, there are still several hurdles that hamper bench-to-bedside transition of polymer-based peptide formulations. During the course of its development, the scale-up phase plays a crucial role because, very often, many of the process limitations that are not apparent on the small scale become significant on a larger scale, and may even lead to the failure of translating a product to industrial dimensions. From the point of view of the

polymeric particle production, the scale-up may significantly alter heat/mass transfer rates, which leads to different regimes for the nucleation and growth of nanoparticles. In this case, the properties of the nanomaterials synthesized (e.g., crystallite size, solid solubility, etc.) and stability of peptide could differ from those produced at smaller scales, which may be undesirable for an intended application. Safety of polymeric materials is another important consideration, which needs to be assessed

Executive summary Background • Peptide drugs bridge the gap between small molecules and large proteins. Compared with small molecules, they exhibit higher efficacy and selectivity, along with lower toxicity. Besides, the lack of tertiary structure and smaller size imparts higher stability, better tissue penetration and lower immunogenicity than proteins and antibodies. Delivery of peptide to the site of action is recognized as a challenge, mainly due to its poor stability. • Polymeric materials with highly tunable structures and physicochemical properties offer multiple remarkable advantages for the delivery of peptides, including improved peptide stability, sustained release and improved pharmacokinetics/biodistribution profiles.

Particulate systems • Biodegradable poly( d,l-lactic-co-glycolic acid) (PLGA)-based nano-/micro-particles are the most extensively used systems for encapsulation of peptide drugs. The incorporation of peptide in PLGA particles allows it to improve stability and to control the release rate of peptide, thus reducing the dose and administration frequency. • The inner cavity of polymeric capsules can be loaded with peptide drugs. The composition and the thickness of the shells determine the capsules’ permeability and the release rates of entrapped drug. The noninvasive delivery of peptide drugs has been extensively studied using these nanocarriers. • Polymeric micelles are favorable nanocarriers for hydrophobic and amphipathic peptides. Apart from common advantages offered by particulate systems, the small size and narrow polydispersity of micelles enable better penetration into various tissues. • The loading of peptide in polyelectrolyte complexes offers several distinct advantages such as high loading efficiency (nearly 100%), simplicity of preparation and preservation of peptide bioactivity. • With 3D and crosslinked polymeric networks, hydrogels are attractive delivery systems for hydrophilic peptides due to their hydrophilicity, high loading capacity, biocompatibility and stimuli-responsive drug release.

Injectable implants • Injectable implants are liquid formulations comprised of biodegradable polymers and therapeutic agents. Upon injection, they transform in situ to solid or semisolid drug depots that release entrapped peptide drugs over an extended period of time.

PEGylation • PEGylation – covalent attachment of PEG polymer chains to peptides – is the most successful approach to improvement of bioavailability of peptides. PEGylation increases the water solubility of peptides, protects them from enzyme degradation and shields antigenic epitopes of the polypeptide. A few PEGylated peptides are already in clinic and several are in clinical trials.

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before proceeding to clinical study. More and more biodegradable polymers are in development, which would certainly address safety-related issues in coming years.

The authors have no relevant affiliations or financial involvement with any organization or entity with a financial

interest in or financial conflict with the subject matter or materials discussed in the manuscript. This includes employment, consultancies, honoraria, stock ownership or options, expert testimony, grants or patents received or pending, or royalties. No writing assistance was utilized in the production of this manuscript.

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Polymer-based vehicles for therapeutic peptide delivery.

During the last decades increasing attention has been paid to peptides as potential therapeutics. However, clinical applications of peptide drugs suff...
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