Diblock and triblock copolymers of polylactide and polyglycolide

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Diblock and triblock copolymers of polylactide and polyglycolide

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Divya Sharma, Lindsey Lipp, Sanjay Arora and Jagdish Singh Department of Pharmaceutical Sciences, College of Health Professions, North Dakota State University, Fargo, ND, United States

13.1 INTRODUCTION The term resorbable comes from the word resorb, which means to be absorbed again. In the context of drug delivery, resorbable means to be broken down and assimilated in the body. Diblock and triblock copolymers composed of polylactide (PLA) and polyglycolide (PGA) are bioresorbable, which means that the brokendown parts of these polymers will get absorbed or dissolved in the body. Block copolymers are a specific type of polymer that constitute different blocks or sections of polymerized monomers. A diblock copolymer is composed of two different chemical blocks, such as PLA-poly(D,L-lactic-co-glycolic acid) (PLGA) and a triblock copolymer is composed of three different chemical blocks where each block has at least one feature absent in the adjacent sections, such as PLA-PLGAPLA. The basic units (monomers) of PLA and PLGA are lactic acid and glycolic acid. PLA is generally synthesized by the ring opening polymerization of two monomers, lactic acid and the cyclic diester lactide using a metal catalyst (e.g., stannous octoate). The polymer PLA exists in an optically active form (L-PLA) which is semicrystalline in nature and an optically inactive racemic form (D, LPLA) which is an amorphous polymer because of irregularities in its polymer chain structure. D, L-PLA forms a more homogenous dispersion of drug within a polymer matrix and is, therefore, the preferred choice over L-PLA for controlled drug delivery systems. Polyglycolide or poly(glycolic acid) (PGA) is a polymer formed by the polycondensation of glycolic acid or most commonly by the ringopening polymerization of the cyclic diester of glycolic acid; glycolide. PGA is hydrolytically unstable and degrades rapidly by random hydrolysis and cellular enzymatic activity, owing to the ester linkage in the backbone, to form glycolic acid which is consumed by cells via the citric acid cycle. Expeditious degradation leading to low mechanical strength usually limits its application as a biomaterial. PLA as compared to glycolic acid is more hydrophobic due to the presence of an extra methyl group resulting in its resistance to hydrolysis and degradation. Consequently, to optimize its degradation rate and pattern, PLA is often Materials for Biomedical Engineering: Thermoset and Thermoplastic Polymers. DOI: https://doi.org/10.1016/B978-0-12-816874-5.00013-X © 2019 Elsevier Inc. All rights reserved.

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FIGURE 13.1 Schematic representation of diblock and triblock copolymers of different types.

copolymerized with other degradable polymers, such as PGA and polyethylene glycol (PEG), which are comparatively hydrophilic in nature. Additionally, PEG serves as a protective layer against the immune system (Laredj-Bourezg et al., 2015). Biodegradable copolymers of ABA and a BAB triblock were introduced by MacroMed, where A represents the hydrophobic polyester block (PLA or PLGA) and B represents the hydrophilic (PEG) block. Some examples of biodegradable triblock copolymers include PLGA-PEG-PLGA; PLA-PEG-PLA; and mPEGPLGA-mPEG (Fig. 13.1). Due to these hydrophobic and hydrophilic moieties, these polymers have the ability to form temperature sensitive polymeric micelles. These polymeric micelles resemble natural carriers (such as viruses and serum lipoproteins) owing to a hydrophilic shell allowing them to circulate in the blood stream for a longer period of time unharmed by the immune system with a small size that prevents their uptake by the reticuloendothelial system (RES), and a hydrophobic core enabling protective encapsulation of drugs/proteins/peptides (Bonacucina et al., 2011). The main objective of this chapter is to summarize the history, synthesis, and characterization of PLA and PGA based diblock and triblock copolymers along with the application of these copolymers as resorbable drug delivery systems.

13.1.1 HISTORY OF POLYLACTIDE PLA is synthesized using two main monomers; lactic acid and the cyclic diester, lactide. Lactic acid was discovered in 1780 by a Swedish chemist, Carl Wilhelm Scheele, who isolated it as an impure brown sirup from sour milk (Kompanje et al., 2007). By the early 1880s lactic acid was commercially produced in the United States, marking the first step toward the study of lactic acid polymers. In 1845, PLA was first synthesized by the condensation of lactic acid

13.1 Introduction

(Jimenez, 2015), and later using reversible polymerization of cyclic esters by heating lactic acid under vacuum (Carothers et al., 1932). Soon after that, PLA had started being used commercially as a fiber material for resorbable sutures (Auras et al., 2010). PLA produced by these methods was expensive and of low molecular weight. Breakthrough research by Cargill Inc. in the early 1990s made acquainted the production of high molecular weight PLA using a commercially viable lactide ring-opening polymerization reaction. The direct condensation route was an equilibrium reaction which made it difficult to remove traces of water at high conversion stages in order to drive the reaction to a higher molecular weight. Additionally, the polymerization method used by Cargill Inc. involved synthesizing both lactide and PLA in the melt thereby avoiding the use of costly and unfriendly solvents. Firstly, a low molecular weight PLA prepolymer was produced by the continuous condensation reaction of aqueous lactic acid, following which a high molecular weight PLA was synthesized using a tin-catalyzed ringopening lactide polymerization reaction. Unreacted lactide was recycled to the beginning of the process by vacuum distillation. The major advantage of this process was the selectivity of the intramolecular cyclization reaction to add from a mixture of lactide stereoisomers using tin catalysis involving coordinationinsertion mechanism with more than 90% conversion and extremely low rate of racemization (Gruber and O’Brien, 2005).

13.1.2 HISTORY OF POLYGLYCOLIDE PGA was one of the very first resorbable polymers to be investigated for use as a biomaterial. It is synthesized from glycolic acid, which is a colorless, odorless, hygroscopic, and crystalline solid with high water solubility. PGA has been known since 1954 as a biodegradable, tough fiber-forming polymer largely used for forming synthetic absorbable sutures (Dexon) of high strength and modulus as well as medical implants (Gilding and Reed, 1979). PGA is a highly crystalline polymer (45% 55%) with a glass transition temperature B35 C and a high melting point in the range 225 230 C. It is insoluble in water and most organic solvents. Fluorinated solvents, such as hexafluoroisopropanol and hexafluoroacetone sesquihydrate, are unique in their capability of dissolving PGA allowing for its spinning or molding into cast films (Schmitt et al., 1971). Several methods have been investigated for the synthesis of PGA. The polycondensation of glycolic acid is the simplest way of synthesizing PGA; involving the heating of glycolic acid at 175 185 C to distill off water followed by continued heating at reduced pressure for a few hours to obtain the low molecular weight byproduct, glycolide. A ring opening polymerization method to synthesize PGA was invented by heating pure glycolide under nitrogen atmosphere in the presence of antimony, zinc, or tin containing compounds as catalysts. At present, stannous octoate is the most commonly used catalyst for this reaction. The reaction is carried out at a temperature below the melting point of PGA and the reactants are allowed to react for about 30 minutes to obtain high molecular weight PGA (Lowe, 1954). In another

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study, solid state polycondensation of halogenoacetates (e.g., sodium chloroacetate) under nitrogen atmosphere in a round bottom flask have also been used effectively to synthesize high molecular weight PGA (Schwarz and Epple, 1999).

13.1.3 SYNTHESIS OF DIBLOCK AND TRIBLOCK COPOLYMERS OF POLYLACTIDE AND POLYGLYCOLIDE Ring-opening polymerization is the most widely used method for the synthesis of diblock and triblock copolymers of PLA, PLGA, and PEG, where the hydrophobic A block is covalently linked to the hydrophilic B block by an ester linkage (Bret et al., 2011). Various authors have used ring-opening polymerization to synthesize different polymers of varying copolymer compositions. The scheme of synthesis remains analogous. Diblock copolymer (PEO:D, L-PLA) synthesis can be achieved by taking equal molar ratios of PEO and D, L-lactide in a round bottom flask. The solvent toluene and a nitrogen atmosphere are used to obtain an anhydrous atmosphere for the reaction, while stannous octoate is used as a catalyst. The reaction is carried out under reflux to prepare the diblock copolymer. Diblock copolymers can be coupled to synthesize a triblock copolymer (PEO:D, L-PLA:IPDI:D,L-PLA:PEO) using isophorone diisocyanate (IPDI) (a coupling agent) dissolved in toluene and refluxed with the diblock copolymer. The obtained copolymer is purified using fractional precipitation from methylene chloride using diethyl ether (Singh et al., 2007). Similarly, mPEG-PLGA diblocks can be synthesized and coupled using IPDI to prepare an mPEG-PLGA-mPEG triblock copolymer (Tang and Singh, 2009). On the other hand, the synthesis of PLA-PEG-PLA triblock copolymers is a one-step process with no intermediate step for coupling. Briefly, calculated molar ratios of PEG (initiator) and D,L-lactide are taken, and D, L-lactide is charged into a three-necked flask containing predried PEG in anhydrous toluene under nitrogen atmosphere. Once all the reactants are in a molten state, stannous octoate is added as the catalyst and the reaction is carried out at 120 C for 12 hours to synthesize the triblock copolymer of the desired copolymer composition. The copolymer obtained by this method can be purified by dissolving the crude copolymer in ice cold water followed by precipitation by heating. This purification step is repeated 2 3 times to remove unreacted monomers and impurities. The final product is freeze-dried to remove the residual water (Al-Tahami et al., 2011; Oak, 2012). Using the mentioned ring-opening polymerization method, copolymers of different block lengths can be achieved by varying the feed ratio of the monomers and the initiator. Studies have suggested that a larger hydrophobic block leads to sustained degradation of the resorbable polymer matrix resulting in controlled delivery of the incorporated drug over a long period. The ratio of hydrophilic and hydrophobic block lengths also affects the aqueous solubility and sol gel transition temperature of the respective copolymer. In different articles, Singh et al. varied the block lengths of both the hydrophobic and hydrophilic blocks while

13.1 Introduction

conserving the water solubility of the polymer as well as its injectability at room temperature, sol gel transition ability, and stability of the gel at 37 C (Chen and Singh, 2005,a,b, 2008; Chen et al., 2005; Al-Tahami, 2007; Tang and Singh, 2009; Al-Tahami et al., 2011). A different method was employed by Wu et al., to synthesize an MPEG-bPLA diblock copolymer using monomers MPEG and LA and carrying out the copolymerization reaction in an oil bath at 140 C for 48 hours. The precipitate obtained was cooled to room temperature and purified by dissolving in anhydrous methylene chloride followed by precipitating out the copolymer using ethyl ether. The MPEG-b-PLA copolymer obtained was then dried under vacuum and used as a macroinitiator; owing to the hydroxyl groups present on its backbone which can initiate the ring opening polymerization of cyclic poly(ethyl ethylene phosphate) (EEP), to generate methoxypolyethylene glycol-poly(D,L-lactide)-poly(ethyl ethylene phosphate) (MPEG-b-PLA-b-PEEP) triblock copolymers, in an additional synthesis reaction (Wu et al., 2011). A transesterification reaction between PLA with PEGNH2 was also explored by a group of authors to synthesize a PLA-b-PEG copolymer (Ta¸skin et al., 2012). Powder form PLGA-PEG-PLGA triblock copolymers can be synthesized by first preparing PLGA in powder form using a direct melt polycondensation method. Then to synthesize the PLGA-PEG-PLGA triblock copolymer in powder form, different proportions of PLGA can be added to a fixed amount of PEG with stannous octoate as the catalyst and the mixture heated in a two necked round bottom flask under nitrogen atmosphere to obtain a crude brown colored product (Gajendiran et al., 2013). Another study demonstrates the synthesis of a PLA/ PEG triblock and multiblock copolymers using an acyl halide-terminated PLA (PLA-diCOCl) prepolymer and anhydrous pyridine (Zhao et al., 2012). Thermosensitive star shaped block copolymers have been investigated for their application in injectable copolymeric drug delivery systems (Lee et al., 2009). The copolymers constituted of fixed molecular weights of PEG, varied mole ratios of D,L-lactide to glycolide (PLGA block), and overall feed ratios of D,L-lactide, glycolide, and three- or four-arm PEG. The monomers are added into a round-bottom flask under nitrogen atmosphere where the three- or four-arm PEG acts as a multifunctional initiator and stannous chloride acts as a catalyst. Bulk ring-opening polymerization is performed and the copolymer product obtained is cooled to room temperature and precipitated for purification using diethyl ether several times (Lee et al., 2009). A three-step synthesis mechanism was also invented to synthesize three-arm star-shaped PLGA-mPEG (3sPLGA-mPEG) and four-arm star-shaped PLGA-mPEG (4sPLGA-mPEG) copolymers using an armfirst method (Zou et al., 2012). In this method, a linear chain hydroxyl-terminated PLGA-mPEG diblock copolymer (LPLGA-mPEG) is synthesized by bulk ringopening polymerization followed by the carboxylation of dried trimethylolpropane (TMP) or pentaerythritol (PTOL) to produce CTMP or CPTOL with three or four carboxyl acid terminal groups using excess amounts of succinic anhydride (SA). Finally, an esterification reaction of two reactive precursors, L-PLGA-mPEG and

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CTMP or CPTOL, is performed using 1,3-dicyclohexylcarbodiimide (DCC) as a dehydrating agent and 4-(dimethylamino) pyridine (DMAP) as a catalyst to obtain a three-arm or four-arm star-shaped PLGA-mPEG block copolymer.

13.1.4 CHARACTERIZATION OF COPOLYMERS OF POLYLACTIDE AND POLYGLYCOLIDE 13.1.4.1 Structural composition analysis The confirmation of the completion of the polymerization reaction of PLA-PEGPLA and PLGA-PEG-PLGA triblock copolymers along with the molecular structure can be determined using Fourier transform infrared (FTIR) spectrometer in the frequency range 4000 1000 cm21 in absorbance mode. The characteristic peak of the carboxylic acid of PLA from 1700 to 1725 cm21 disappears and a new peak appears in the region of 1730 1750 cm21 due to the newly formed ester groups in the FTIR spectra. The characteristic peak for isophorone diisocyanate, used as a coupling agent, at 2175 cm21 can be used as an indicator of the completion of the coupling reaction in PEO:D,L-PLA:IPDI:D,L-PLA:PEO type coupled diblock copolymers (Singh et al., 2007). The characteristic signals of the PEG ether band and PLA ester carbonyl band can also be seen at 1086 and 1755 cm21 respectively, in a such a diblock or triblock copolymer (Ta¸skin et al., 2012). Additionally, both proton (1H) and carbon (13C) nuclear magnetic resonance (NMR) are established techniques to determine the chemical structure and structural composition of PLA and PGA block copolymers. The analysis may be performed using an NMR spectrometer operating at 300 or 400 MHz. The copolymer is dissolved in an organic solvent, such as deuterated dimethyl sulfoxide (DMSOd6) or deuterated chloroform (CDCl3), with tetramethylsilane (TMS) signal as the internal reference standard. Resonances in the B5.2 5.0 ppm ( O CH) and B1.5 1.4 ppm (CH3) ranges belong to PLA blocks. The main chain methylene signals of PEG (in a PLA-PEG-PLA triblock copolymer) usually show in the 3.7 3.3 ppm range. The α-methylene protons (PLA-COO-CH2) and hydroxylated methine ( CH) protons of lactyl end units appear together in the range of B4.3 4.1 ppm. If the copolymerization did not take place effectively, carboxylated end units of lactyl and methine protons of free lactic acid appear in the B5.0 4.0 ppm range and 4.03 ppm respectively, in a 1H NMR spectrum (Rashkov et al., 1996). Similar characteristic peaks were reported by several other authors with slight variations complying with the change in the type of block copolymer and the monomers in the copolymer chain backbone (Jeong et al., 1999; Tang and Singh, 2009; Al-Tahami et al., 2011). The structural composition, graft ratio, and the number average molecular weight (Mn) of the polymers can then be calculated by analyzing the integrated signals corresponding to chemical groups CH and CH3 of LA and CH2 of EG in an 1H NMR spectrum (Ta¸skin et al., 2012; Oak, 2012). The spectrum of 13C NMR has also been used to confirm

13.1 Introduction

the presence of PEG and PLA blocks by the characteristic peak of the methylene ( CH2) group of the PEG block at B71 ppm and the carbonyl ( CQO), methine ( CH), and methyl ( CH3) groups of the PLA block at B170, B69.4, and B17 ppm respectively (Oak, 2012). Gel permeation chromatography (GPC) is used to further determine the number average molecular weight (Mn), weight average molecular weight (Mw), and the molecular weight distribution (polydispersity index, PDI) of the synthesized copolymer. Polystyrene standards are used for calibration and tetrahydrofuran is a popularly used carrier solvent for the GPC analysis of PLA/PEG copolymers (Zhao et al., 2012; Oak, 2012).

13.1.4.2 Aqueous solubility and injectability Aqueous solubility and injectability are two lucrative properties making the copolymers of PLA/PGA versatile for drug delivery use; the main advantage being avoidance of toxic organic solvents. The concentration at which the copolymers are soluble below the gelation temperature is called the functional concentration (Rathi et al., 1998). The copolymer dissolves in cold water due to the PEG blocks keeping the copolymer in solution and the hydrophobic PLGA/PLA segments forming associative crosslinks. This happens due to the hydrogen bonding between hydrophilic PEG blocks and water making the copolymer soluble. This effect is dominant at lower temperatures. The copolymer concentration in water can be varied to allow injectability at room temperature and also to tailor the drug release profile. As the temperature increases the hydrogen bonding gets weaker and the hydrophobic forces in the hydrophobic PLGA/PLA blocks get strengthened and become dominant. In this way, a change in temperature leads to the reversible sol to gel transition of an aqueous copolymer solution of copolymers of PLA/PGA/PLGA (Bret et al., 2011). The transition temperature can be varied by changing the hydrophobic and hydrophilic block lengths of these copolymers. For injectable drug delivery application, it is desired that the copolymer solution incorporating the desired therapeutic should be injectable (sol form) at room temperature and on administration at physiological temperature (37 C) transform into a stable gel depot at the injection site (Tang and Singh, 2009). The total molecular weight of PLA and PGA copolymers for optimum solubility and reversible thermogelation should lie between 3500 and 4100 Da for ABA-type and 4000 4600 Da for BAB-type copolymers. For both types, the average molecular weight of hydrophilic block B (preferably PEG) should fall between 600 and 2200, while the overall weight percentage of the hydrophobic block relative to the hydrophilic block should be preferably high, between 65% and 78%. In BAB-type copolymers it has been found that the copolymer composition (ratio of PLA/PEG) and the total molecular weight of the copolymer have a striking effect on the release profile, especially for hydrophilic drugs (Rathi et al., 1998).

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13.1.4.3 Phase transition Response to stimulus is an innate property of living systems. The ability to design a system to manifest this property has been a starting point to several sterling researches and inventions. Thermosensitive copolymers of PLA and PGA undergo reversible in situ sol-to-gel transition in response to temperature changes. The transition mechanism of such aqueous copolymeric solutions is related to the presence of both hydrophilic and hydrophobic parts in their structure and usually to a lower critical solution temperature (LCST) in the aqueous solution (Chen and Singh, 2008). As the temperature increases above the LCST the equilibrium shifts from unimers to spherical micelles. The copolymer water interactions become thermodynamically unfavorable in comparison to water water or copolymer copolymer interactions, leading to dehydration of the solvated copolymer chains and finally to a transition into gel state (micelle packing). On increasing the temperature above the upper critical transition temperature (UCST), the polymer precipitates (Oak, 2012). The transition from gel to sol is related to the shrinkage of the hydrophilic component’s corona in the micelles owing to the effect of temperature on its solubility and the interaction of its chains with the hydrophobic hard core (Jeong and Gutowska, 2002). In copolymers containing PEG, the PEG chains orient themselves to align, forming the outer hydrophilic shell of the micelles facing the external aqueous environment. This layer of PEG acts as a barrier by reducing interactions with foreign molecules resulting from steric and hydrated repulsion. This results in the increased stability and shelf life of such systems (Makadia and Siegel, 2011; Bret et al., 2011). It has also been noted that PEG-PLGA-PEG triblock copolymers in aqueous solutions show increased polymer-polymer interaction as compared to polymer-solvent interactions. In other words, it has also been suggested that with an increase in temperature the polymeric micelles grow by increasing their diameter and eventually aggregate, thus, driving the sol gel transition (Bonacucina et al., 2011).

13.1.4.4 Thermal properties Thermogravimetric and differential thermal analysis (TG-DTA) of copolymers provides useful data to assess the influence of the copolymer composition on the degradation behavior. A TGA analyzer instrument is used for thermal characterization. Gajendiran et al. (2013) quantitatively assessed the degradation behavior of a PLGA-PEG-PLGA triblock copolymer using a thermal characterization method. According to their study, PLGA-PEG-PLGA triblock copolymers degrade in two steps with the loss of PLGA at B300 C and PEG at B410 C. Various changes in the degradation patterns with changing copolymer compositions were also reported. In a different study differential scanning calorimeter (DSC) was used to investigate the thermal properties of PLA-PEO-PLA triblock copolymers. It was found that the melting temperature (Tm) of PLA/PEO copolymers was lower than that of PEG alone, which ultimately has significant

13.1 Introduction

implications on the glass transition temperature (Tg) and crystallization peak (Tc) of these copolymers (Rashkov et al., 1996).

13.1.4.5 Crystallization behavior The effect of different copolymer compositions and architectures on the thermal properties and crystal structures of block copolymers MPEG-b-PLLA; PLLA-bPEG-b-PLLA; and four-arm PEG-b-PLLA has been intensely investigated (Zhou et al., 2015). DSC was used to scrutinize the thermal properties of the copolymers. The instrument was calibrated with pure indium and experiments were performed under nitrogen flow. Heating, cooling, and second heating scans were recorded for different MPEG-PLLA block copolymer compositions and the crystallinity of PLLA and PEG were calculated. Alongside this, wide angle X-ray diffraction (WAXD) measurements were also taken to study the effect of the chain connectivity, composition, and architecture of these copolymers on the crystal structures of PLLA and PEG. It was reported that melting point and crystallinity were affected by increasing molecular weight, arm length, and number of arms of PLLA in the MPEG-PLLA and PEG-PLLA block copolymers. This probably happens due to the formation of PLLA crystallites which causes the internment of PEG, resulting in increased difficulty for PEG to be packed into the crystal lattice. This suggests that varying the architecture and molecular weights of PLA/PEG block copolymers alters the consequential properties and exploring those will open new doors for the application of branched PLA/PEG block copolymers for controlled drug delivery applications (Zhou et al., 2015). WAXD data were also shown to support the fact that the formation of crystalline hydrophobic domains in PLLA gels resulted in a higher stiffness while in racemic PLA resulted in the formation of easily degradable amorphous hydrophobic domains due to the stereo random structure (Sanabria-DeLong et al., 2006). Thus, changing a simple chemical parameter, that is, stereo-regularity, can help tailor PLA containing block copolymers, such as PLA-PEO-PLA, for controlled drug release. Overall, it has been well investigated and proved that PLA and PEG blocks in a diblock or triblock copolymer affect the crystallization behavior of each other (Yang et al., 2006), and the gradually the increasing confinement of PEG is dictated by the crystallization of the PLA block (Huang and Paul, 2007). Another study supported this fact by suggesting that the crystallizability of PEO blocks depends on their length and can be reduced by copolymerization with PLA blocks (Rashkov et al., 1996).

13.1.4.6 Biocompatibility, cytotoxicity, and biodegradability In general, a system containing a drug suspended in the aqueous solution of PLA/ PGA/PLGA based diblock and triblock copolymers causes minimum toxicity and mechanical irritation to the surrounding tissues due to their inherent mucomimetic property, pliability of the gel, and biodegradation into lactic acid, glycolic acid, and ethylene glycol, which naturally dissipate from the body (Rathi et al., 1998; Bonacucina et al., 2011). PEG forms the hydrophilic shell of these copolymeric

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micelles in aqueous media making them nonimmunogenic, biocompatible, and soluble in water. It has also been documented that PEG of a molecular weight below 30,000 is easily eliminated by the body (Rathi et al., 1998). A plethora of data are available from cytotoxicity and biocompatibility studies on PLA/PGA based copolymers. 3-(4,5-dimethylthiazol-2-yl)-2,5- diphenyltetrazolium bromide (MTT) assay is a popularly used method to test the biocompatibility of a system in vitro. The principle of this assay is based on the ability of mitochondrial succinate dehydrogenase present in living cells to reduce this MTT dye to water-insoluble purple formazan crystals. The formazan crystals are then dissolved using isopropanol and the absorbance is measured using a plate reader. MTT assay is useful for assessing the subtle toxicity of systems which may not kill cells rapidly (i.e., within 24 72 hours) but may affect the metabolic and other functions of the cells necessary to maintain viability. A high absorbance relates to a high viability of the cells and hence, the low cytotoxicity of the sample tested. The cytotoxicity testing of PLA and PGA constituting block copolymers has been reported with insignificant difference from the control (cells incubated with growth medium only) affirming their biocompatibility (Oak and Singh, 2012a; Chen and Singh, 2008; Tang and Singh, 2009; Al-Tahami et al., 2011; Zou et al., 2012). In vivo biocompatibility can be tested by injecting animal models with these copolymeric drug delivery systems and comparing the histology of the injection site skin tissue using hematoxylin-eosin (H and E) stain to test for inflammatory responses, Masson’s trichrome staining to examine the vascularization, and Gomori’s trichome stain to test for collagen deposition, after specific time intervals. The safety and in vivo biocompatibility of PLA/PEG based copolymeric systems have been published in various articles. The overall results indicate that up to a week following subcutaneous injection of this copolymeric system, the infiltration of neutrophils and macrophages to the injection site occurs, demonstrating clear incidence of an acute inflammatory response, which subsides to a milder chronic inflammatory response at 30 days post administration with the presence of a few inflammatory cells, and finally at about 90 days closely resembles the control, indicating restoration to normal tissue with no signs of necrosis and/or chronic inflammation (Ma et al., 2014; Chen and Singh, 2008; Oak and Singh, 2012a). Biocompatibility of PLA and its copolymers for orthopedic, ophthalmic, otologic, skin, central nervous system, pulmonary system, parotid glands, urinary tract, and cardiovascular applications has also been established with a good safety profile compared to conventionally used devices and implants (Ramot et al., 2015). The in vivo biocompatibility testing of star-shaped block copolymers has also been similarly performed (Zou et al., 2012). The histology after 30 days showed almost complete restoration to normal tissue and no significant tissue necrosis, hyperemia, edema, hemorrhaging, or muscle damage. A Masson’s staining experiment showed that vascularization took place after 15 days, suggesting that starshaped PLGA-mPEG copolymers supported vascular in-growth, and overall they have good biocompatibility.

13.2 Resorbable Thermosensitive Polymers

PLA/PEG constituting copolymers degrade by nonenzymatic hydrolysis of ester bonds to nontoxic products which are naturally eliminated by the body (Zhao et al., 2012; Oak, 2012). 1H NMR and GPC can be used to determine the reduction in molecular weight of these copolymers while undergoing hydrolytic degradation. Copolymer composition affects the degree of gel hydration affecting the degradation rate of the copolymer, which in turn affects the permeability coefficient of the incorporated drug through the gel matrix (Jeong et al., 2000; Ma et al., 2014). The in vitro degradation of block copolymer hydrogels happens similarly; by the hydrolysis of ester bonds accompanied by the erosion of the gel in PBS solution at physiological temperature (Zou et al., 2012).

13.2 RESORBABLE THERMOSENSITIVE POLYMERS Temperature sensitive or thermoresponsive polymers are the most widely studied type of stimuli-sensitive smart polymers for drug delivery owing to the ease and benefit of exploiting changes in their state in response to physiological temperature. Numerous research papers provide fortifying evidence that drug delivery using thermosensitive polymeric systems is progressing at a rapid rate. Poly(N-isopropylacrylamide) (poly-NIPAAM) was the first most extensively studied prototype of thermosensitive polymers (Tang and Singh, 2009). It was synthesized in the early 1950s by free radical polymerization of N-isopropylacrylamide (Schild, 1992). However, due to its toxicity as well as low mechanical strength, poly-NIPAAM did not succeed for drug delivery applications (Bae et al., 1987). Later, ABA type of nonionic triblock copolymers (poloxamers), containing poly(ethylene oxide) as the hydrophilic block B and poly(propylene oxide) as the hydrophobic block A (poly(ethylene oxide)-co-poly(propylene oxide)-copoly(ethylene oxide) (PEO PPO PEO) copolymer, also called Pluronics) received the US Food and Drug Administration (USFDA) approval as a pharmaceutical excipient, but could not progress further as a pharmaceutical drug delivery system due to the nonbiodegradability of the hydrophobic PPO block and related toxicity (Wasan et al., 2003). In 1997, MacroMed Inc. replaced the nonbiodegradable block in Pluronics with a biodegradable and biocompatible PLA block to develop PEO PLA PEO thermosensitive triblock copolymer (Jeong et al., 1997). In an effort to optimize the phase transition behavior, degradation pattern, and hence the drug release kinetics, several copolymer compositions have been investigated over the years with encouraging results. Broadly, a mixture of the biodegradable/biocompatible copolymer and drugs/ proteins/peptides can be prepared by simple mixing in aqueous copolymer solution below gelation temperature to form a partially dissolved (colloidal state dispersion, such as suspension or emulsion) or completely dissolved drug delivery system which could be injected parenterally, administered topically/transdermally,

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and/or inserted into a cavity (ocular, vaginal, transurethral, rectal, nasal, oral, or aural). On administration, the formulation would undergo thermal gelation at physiological temperature (typically body temperature would be above the transition temperature) forming a depot and entrapping the drug in the polymer matrix (Zentner and Shih, 2004). The release from these copolymeric delivery system follows two mechanisms acting simultaneously: diffusion of the incorporated drug and degradation of the polymer matrix. Mostly, the initial release is diffusion-controlled and the later stage is a combination of both with degradation being dominant (Jeong et al., 2000). Thermosensitive copolymers of PLA and PGA retain water equivalent to B10% of the total weight of the hydrogel, which allows for the swelling of the gel depot and a diffusion pathway for the incorporated drug molecules. Water retention has been observed to vary with the ratio of the hydrophilic/hydrophobic content in the copolymer (Al-Tahami, 2007; Oak, 2012). It has also been reported that during the erosion of the hydrogel matrix (in the later phase), there is a preferential loss of hydrophilic segments (PEG-rich) rendering the remaining gel matrix hydrophobic with reduced water retention and swelling resulting in decreased copolymer degradation, ultimately leading to reduced drug release (Bret et al., 2011). Additionally, the drug release profile can also be altered by varying the copolymer concentration (Oak, 2012). Concentrations between 10% and 30% w/w are most preferred for drug delivery as lower concentrations were found to transition to form a weak gel, and higher concentrations are too viscous to be injectable. Optimization is required to reach a balance between a strong gel network and the desired release rate (Rathi et al., 1998). A model hydrophilic drug (ketoprofen) and a model hydrophobic drug (spironolactone) were tested by Jeong et al. using a PEG-PLGA-PEG thermosensitive copolymer to assess the release model of such a copolymeric system (Jeong et al., 2000). A domain structure was assumed with the drugs partitioning between the hydrophilic shell domain and the hydrophobic core domain. Drug release from the hydrophilic shell can be explained by diffusion and that from the hydrophobic core by the modified Higuchi equation. Thermosensitive copolymers made of PLA and PGA have the advantages of being easy to manufacture, soluble in water, showing avoidance of toxic organic solvents, simple formulation, ease of administration, controlled release of the incorporated drug, and ability to adjust copolymer composition for controlling the release period by modifying the degradation rate, the permeability of the matrix, and hence the drug release profile. Drug delivery systems using thermosensitive diblock and triblock copolymers of PLA and PGA will be discussed in detail next.

13.2.1 THERMOSENSITIVE POLYMER-BASED DRUG DELIVERY SYSTEMS The use of amphiphilic block copolymers for drug delivery was first proposed in the early 1980s (Pratten et al., 1985). The innovation of using PLA/PGA/PLGA/

13.2 Resorbable Thermosensitive Polymers

PEG copolymers for drug delivery applications lies in the simplicity of using these copolymers to deliver a wide variety of drugs, hormones, as well as sensitive proteins and peptides with efficacy. Sustained delivery of various such therapeutics is highly desirable as conventional drug delivery methods are far from ideal. Frequent subcutaneous, intramuscular, or intravenous injections at short intervals, daily application of patches which adhere poorly and/or cause irritation, poor oral bioavailability, and short half-life after parenteral administration, confront the need for a better controlled delivery system without toxicity (Chen and Singh, 2005a). Thermosensitive copolymers of PLA and PGA have shown good results both in vitro and in vivo for a large number of such therapeutics, with some currently in the clinical testing phase discussed in Section 13.2.2. Succinic anhydride terminated diblock copolymer methoxypoly(ethylene glycol)-b-poly(lactide) (mPEG-PLA-SA) has been investigated to synthesize 7-Ethyl10-hydroxy camptothecin (SN38) drug conjugated polymeric micelles (Lu et al., 2016). SN38, an active metabolite of irinotecan is a potent topoisomerase I inhibitor. Its clinical applicability as an antineoplastic drug is limited by its hydrophobicity and the instability of the lactone ring in its structure at a physiological pH. Drug conjugates with amphiphilic diblock copolymers allow for the formation of polymeric micelles as drug carriers. Advantages of this micellar drug delivery system include passive accumulation of polymeric micelles in solid tumors via enhanced penetration and retention effect (EPR), increased therapeutic efficacy, reduced side effects, less frequent drug administrations, and improved patient compliance. The chain lengths of mPEG and PLA were shown to have a large effect on the particle size of the drug conjugated micelles as well as antitumor efficacy, both in vitro and in vivo. Polymer drug conjugate micelles were found to be less toxic and more efficacious drug delivery systems for cancer treatment. Enhanced controlled release properties were also depicted by this drug delivery system, which were mainly due to the shielding effect of the hydrophilic mPEG shell against plasma proteins, thereby reducing clearance via the mononuclear phagocyte system; while the hydrophobic core (due to PLA) showed the ability to incorporate hydrophobic drugs and allowing for their controlled release. Similarly, another study reported the solubilizing efficacy of typical amphiphilic block copolymers by studying the poorly water soluble drugs, paclitaxel and cyclosporin A (Rathi et al., 1998). In another study, the ocular pharmacokinetics of dexamethasone acetate was evaluated in rabbits using the microdialysis method with 20% w/w PLGA-PEGPLGA copolymer solution and compared to regular eye drops. A sevenfold higher maximum serum concentration (Cmax) and a 7.89-fold larger area under the curve (AUC) was obtained with the thermosensitive in situ gelling copolymer, thus, validating enhanced corneal permeability, prolonged precorneal retention, improved bioavailability, and higher drug efficacy (Gao et al., 2010). A group of researchers also suggested that various additives, such as sugars, surfactants, salts, amino acids, proteins, and other substances, can be readily incorporated in these block copolymers as and when required to modify the release characteristics and/or

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stability of the drug compound (Rathi et al., 1998). A long acting formulation of exendin-4 (exenatide, EXT) incorporated in a PLGA-PEG-PLGA triblock copolymer was tested with and without excipients (zinc acetate, PEG, and sucrose) to control the burst release of EXT, both in vitro and in vivo, with promising results (Li et al., 2013). EXT is an incretin-mimetic polypeptide established to enhance glucose-dependent insulin secretion for the treatment of type II diabetes. Due to the viscous environment of the gel, the hydrolytic instability of the polypeptide significantly decreased. The addition of excipients reduced the burst release with zinc acetate showing the best effect. In other studies, controlled release of insulin has been widely studied using this delivery system injected subcutaneously. Regel (PLGA-PEG-PLGA) polymer was used for the controlled delivery of recombinant human insulin for the basal requirement of insulin for up to 15 days (Kim et al., 2001). Meanwhile, Zentner et al. studied the release of paclitaxel, porcine growth hormone, glycosylated colony-stimulating factor, and recombinant hepatitis B surface antigen, using Regel polymer with positive influence on drug effectiveness and stability (Zentner et al., 2001). The controlled delivery of levonorgestrel, testosterone, and growth hormone has also been investigated from PLGA-PEGPLGA thermosensitive copolymer-based delivery systems. The effects of varying block lengths on release profiles were observed and conclusions were drawn for the effect of different drug types on the release profile and duration of drug release (Chen and Singh, 2005a,b, 2008). In order to simultaneously utilize the benefits of PLGA-PEG-PLGA, in terms of ease of formulation, localized administration, biodegradability, low systemic toxicity, and sustained drug delivery in combination with drug therapy, in an effort to improve the anticancer efficacy of drugs against osteosarcoma, a localized codelivery system of PLK1shRNA/PEI-Lys complexes and doxorubicin (DOX) suspended in a PLGA-PEG-PLGA thermosensitive hydrogel was developed (Ma et al., 2014). The delivery system allowed for the sustained codelivery of the incorporated drugs with no cytotoxicity, biocompatibility, and significant synergistic antitumor efficacy. Moreover, localized delivery to the tumor was beneficial in reducing systemic toxicity as observed by ex vivo histological analysis of major organs in Saos-2 xenograft models. Furthermore, the controlled delivery of proteins and peptides is a highly challenging effort owing to low half-life, implicit instability, and structural constraints. These are also some preeminent reasons that render basal level insulin delivery to type I diabetes patients a daunting task. Multiple frequent injections or round the clock insulin pumps are conventionally used currently in order to maintain normoglycemia. The delivery of sensitive proteins and peptides has been extensively studied using PLA/PLGA based triblock thermosensitive copolymers. PLGA-PEG-PLGA thermosensitive triblock copolymers showed a controlled release of different proteins for B2 weeks (Kim et al., 2001; Singh et al., 2007). This copolymer system demonstrated high burst release of hydrophilic drugs like insulin owing to the higher hydrophilic GA content in the copolymer backbone. PLA being more hydrophobic than PLGA was hypothesized to undergo a slower

13.2 Resorbable Thermosensitive Polymers

degradation owing to retarded hydration, swelling, and hydrolysis, and was further investigated for the controlled basal delivery of insulin. PLA-PEG-PLA triblock copolymers showed significantly lower burst release with desirable zeroorder release profiles over a period of 2 3 months (Al-Tahami et al., 2011). Later, by incorporation of chitosan zinc insulin complex into a PLA-PEG-PLA copolymer, a controlled basal insulin delivery of B84 days was obtained in vitro (Oak and Singh, 2012b). In an additional study, the biocompatibility of the delivery system and efficacy of the released insulin was successfully confirmed in vivo using a streptozotocin-induced diabetic rat model (Oak and Singh, 2012a). Simultaneously, an additional advantage observed with these block copolymers is the protection of the incorporated drugs from chemical degradation which is extremely helpful for easily chemically degraded drugs as well as sensitive protein and peptide-based drugs. Chen et al. (2005) studied the release profile of the model protein, lysozyme, using PLGA-PEG-PLGA thermosensitive copolymers of varying block lengths and aqueous copolymer concentrations. The controlled delivery of lysozyme was reported in a biologically active form, with significant lowering of burst release with increasing copolymer concentration (Chen et al., 2005). Similar studies were done with an mPEG-PLGA-mPEG copolymer, and the effect of extending the PLGA block resulting in decreased degradation and controlled release of the protein for a longer duration was reported (Tang and Singh, 2009). Controlled delivery of salmon calcitonin, a polypeptide hormone for the prevention and management of osteoporosis, was also investigated using an mPEG-PLGA-mPEG triblock copolymer in vitro in a female rat model. Calcitonin suspended in 40% w/v aqueous copolymer solution administered subcutaneously was seen to protect the rat from methylprednisolone acetate induced osteoporosis for up to 40 days (Tang and Singh, 2010). Representative examples of sustained release depot based drug delivery systems of PLA and PGA diblock and triblock copolymers are summarized in Table 13.1.

13.2.2 COMMERCIAL AND INVESTIGATIONAL EXAMPLES Long-term controlled delivery of hydrophilic and hydrophobic drug substances via the parenteral route is an attractive approach. PLA and PEG diblock and triblock copolymers are profoundly investigated for this purpose due to their myriad benefits. Ensuing a great deal of success in veterinary medication (Matschke et al., 2002), there is abundant appreciation of the potential for various applications of these copolymers. Substantial investigations are being carried out and several of them have made it to clinical trials. A sterile, lyophilized micellar formulation of paclitaxel Genexol-PM (Cynviloq) using a PLA-PEG diblock copolymer has been approved by the USFDA to be marketed in Europe and Korea (Pillai, 2014). In this copolymeric colloidal carrier, PEG served as a nonimmunogenic outer shell while the PLA in the hydrophobic core solubilized the hydrophobic drug. The maximum tolerated dose (MTD) of paclitaxel and its biodistribution in the liver, spleen, kidneys,

463

Table 13.1 Representative Examples of Depot-Based Drug Delivery Systems of Polylactide and Polyglycolide Diblock and Triblock Copolymers Copolymer

Drug or Active Ingredients

Major Effects

PLGA-PEGPLGA

Lysozyme

mPEG-PLA

7-Ethyl-10-hydroxy camptothecin

• • • •

PLGA-PEGPLGA

Cyclosporin, paclitaxel

PLGA-PEGPLGA

Dexamethasone acetate

PLGA-PEGPLGA

Exendin-4

PLGA-PEGPLGA PLGA-PEGPLGA (Regel)

Recombinant human insulin Paclitaxel, pGH, G-CSF, insulin, rHbsAg

• • • • • • • • • • • • •

PLGA-PEGPLGA

Levonorgestrel, testosterone, growth hormone

PLGA-PEGPLGA

PLK1shRNA/PEI-Lys complexes and doxorubicin

PLA-PEG-PLA

Insulin, zinc-insulin hexamers, chitosan-zinc-insulin complex

• • • • • • • •

mPEG-PLGAmPEG

Lysozyme, salmon calcitonin

• •

Increasing the PLGA block lengths of copolymers decreased initial burst release. Increasing copolymer concentration reduced the rate of drug release. Self-assembling micelles forming mPEG-PLA-SN38 conjugates were synthesized. Passive accumulation of polymeric micelles in solid tumors via enhanced penetration and retention effect was observed in vitro and in vivo. Reduced toxicity and increased anticancer efficacy of the system. Improved solubilization of poorly water-soluble drugs. Increased chemical stability. Enhanced corneal permeability and prolonged precorneal retention. Increased Cmax and AUC. Improved bioavailability, and higher drug efficacy. Increased stability. Possible addition of excipients reduced burst release. Controlled basal insulin release observed up to 15 days in vitro and in vivo after single s.c. injection. Reduced clearance of paclitaxel after direct intratumoral injection with minimal distribution into any organ. Controlled release of paclitaxel for B50 days. Controlled release of equivalent amount of pGH, insulin and G-CSF after single s.c. administration compared to daily i.v. conventional therapy. Regel/rHBsAg increased rHBsAg-specific antibody titers by 6 times compared to commercial vaccine Engerix-B. Increasing the hydrophobic PLGA block length of copolymers significantly decreased the release rate. Controlled zero-order in vitro release was observed. Enhanced absolute bioavailability of pGH compared to s.c. aqueous pGH solution. Synergistic antitumor efficacy of co-incorporated drugs. Reduced systemic toxicity owing to localized tumor delivery. Optimization of drug release rate by varying copolymer composition and aqueous copolymer concentration. Significantly lower burst release and controlled zero-order release profile of the system. Controlled basal insulin delivery in vitro and in vivo in chemically and structurally stable form. Controlled release of the protein for a longer duration by extending the PLGA block resulting from decreased degradation rate of the copolymer matrix. Protection of in vivo animal model from methylprednisolone acetate induced osteoporosis for up to 40 days.

References Chen et al. (2005) Lu et al. (2016)

Rathi et al. (1998) Gao et al. (2010)

Li et al. (2013) Kim et al. (2001) Zentner et al. (2001)

Chen and Singh (2005a,b, 2008)

Ma et al. (2014) Al-Tahami et al. (2011), Oak and Singh (2012a,b)

Tang and Singh (2009, 2010)

13.2 Resorbable Thermosensitive Polymers

lungs, heart, and in tumors were both found to be increased by two to threefold in preclinical studies. The antitumor efficacy was significantly improved compared to free paclitaxel. Clinical studies have demonstrated a better safety profile, higher efficacy, and better response rates of Genexol-PM in patients with metastatic breast cancer and advanced pancreatic cancer. Combination chemotherapy of Genexol-PM with cisplatin allowed for the administration of higher doses of paclitaxel, and showed significant results. Genexol-PM also increased the response rates for patients who were not responsive to conventional paclitaxel therapy. It is also considered a potentially effective treatment alternative for gemcitabineresistant pancreatic ductal adenocarcinoma based on promising in vivo data. Genexol-PM has completed phase I and phase II trials as a treatment strategy in metastatic breast cancer, nonsmall cell lung carcinoma, pancreatic cancer, ovarian cancer, and bladder cancer. Studies are underway for the treatment of several other diseases, as well as phase III and phase IV studies in recurring breast cancer patients (Jain et al., 2016). Additionally, a paclitaxel incorporated PLGA-PEG-PLGA triblock copolymeric formulation based on MacroMed’s proprietary ReGel technology, called OncoGel, was investigated for local tumor management. ReGel is a water soluble thermosensitive copolymer designed to undergo reversible phase transition from an injectable low viscosity solution incorporating a drug of choice (sol-state) between 2 C and 15 C, to a controlled release gel depot at physiological temperature (37 C) (Elstad and Fowers, 2009). Phase I clinical trials of OncoGel on patients with inoperable solid tumors showed mixed results. A paclitaxel dose of up to 2.0 mg/cm3 was well tolerated and the drug remained localized at the injection site. However, pain, injection site bruising, redness, irritation, muscle spasm, and postprocedural discharge were observed as major side effects. In another study, OncoGel demonstrated disappointing results in a phase IIb designed model to determine its impact on presurgical potential in patients with esophageal cancer (Elstad and Fowers, 2009; Jain et al., 2016). Alongside, a phase I/II dose escalation study for local injection of OncoGel in patients having recurring glioma was terminated within 8 weeks owing to dose-limiting toxicities with serious vascular adverse effects mainly subdural hematoma (Clinicaltrials.gov, 2017a,b).

13.2.3 LIMITATIONS OF THERMOSENSITIVE POLYMERS Thermosensitive copolymers are simple and elegant drug delivery systems which are easy to formulate and on administration at body temperature show instantaneous sol gel transition. Nonetheless, it is advised to keep the aqueous copolymer solution at 4 C to maintain good injectability and low viscosity of the system as the viscosity of the system generally tends to increase as the temperature approaches room temperature. This may be a possible hurdle that will need discretion while taking such drug delivery systems to a clinical setting (Gong et al., 2009; Vaishya et al., 2015). Alongside this, drugs with high water solubility and small size readily diffuse from these copolymer matrices because of the highly

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porous microstructure, low degree of crosslinking, and increased hydration and swelling of these copolymers. Studies have also reported that an initial burst release occurs owing to the drug being located close to the surface of these copolymeric delivery systems (Oak, 2012). Numerous studies have also reported irreversible protein aggregation in these delivery systems resulting in incomplete protein release in vitro and less than 100% bioavailability in vivo. Though, this issue can be resolved to some extent by decreasing the drug loading (Vaishya et al., 2015). Furthermore, the addition PEG has been shown to limit the encapsulation efficiency of various drugs and proteins, even when adopting the most pertinent formulation techniques. This effect is suspected to be an effect of steric interference and possible drug/protein polymer interactions (Bret et al., 2011).

13.3 RESORBABLE NANOPARTICLES Nanoparticles have become an extremely popular drug carrier over the years for several reasons. Polymeric nanoparticles offer a lot of potential variations in composition, which plays into their ability to be fine-tuned for a specific therapeutic or delivery target. Furthermore, different methods can be used for incorporating therapeutics to increase entrapment efficiency while maintaining the stability of the system. Generally, polymeric nanoparticles range in size from 10 to 100 nm and can vary structurally as core shell micelles, nanospheres of polymeric matrices, and polymeric shells surrounding aqueous cores as nanocapsules (Roney et al., 2005). Resorbable polymers, such as lactide and glycolide, are useful candidates when designing nanoparticles. Diblock and triblock copolymers based upon PLA and GLA have proven to be biocompatible. In addition, since the formation of nanoparticles relies heavily on hydrophobic forces, the ability to alter the hydrophobic and hydrophilic portions of the amphiphilic block copolymers allows for precise designing of nanoparticles with modifiable physicochemical characteristics, such as size, charge, and entrapment efficiency. This can be done by tailoring their composition either by changing the lactide to glycolide ratio and/or by changing the molecular weight of the hydrophilic or hydrophobic blocks. PEG, polymethacrylate, polyethyleneimine (PEI), and polyethylene oxide (PEO) are commonly used as the hydrophilic blocks of nanoparticles based upon PLA, PGA, and PLGA. (Park et al., 2005)

13.3.1 NANOPARTICLE PREPARATION AND CHARACTERIZATION TECHNIQUES The general polymer synthesis is the same as previously discussed. There are multiple methods that can be employed for nanoparticle formation. Factors such as the hydrophobicity and stability of the therapeutic to be entrapped are considered when deciding which preparation method to use. Each method has potential

13.3 Resorbable Nanoparticles

pitfalls that should be kept in mind, especially when optimizing size and/or entrapment efficiency. Additionally, surface modification is frequently used to target the delivery of nanoparticles to specific tissues or to further stabilize the nanoparticles. It should be noted that some preparation methods may facilitate surface modification better than others. Some of the common preparation methods include emulsification, phase separation, dialysis, and spray drying. In the single emulsion or emulsion/solvent evaporation method the copolymers are dissolved in an organic solvent along with the hydrophobic drug or therapeutic. The organic solvent solution is then emulsified in an aqueous solution often by the aid of a surfactant. Finally, the organic solvent is evaporated and nanoparticles can be isolated (Park et al., 2005). Like with single emulsion, the precipitation solvent diffusion technique produces nanoparticles using an organic solvent for the dissolution of copolymers and hydrophobic therapeutics, but in this method the organic solvent is miscible in water. Upon diffusion of the organic solvent into the aqueous phase, nanoparticles precipitate immediately and are recovered by evaporation of solvent or dialysis against water. A double emulsion technique can be employed by preparing a water in oil in water (W/O/W) system, in which a primary water in oil emulsion is added to an aqueous solution that often contains a surfactant, such as polyvinyl alcohol (PVA) or tween (Park et al., 2005). The primary emulsion may be used to solubilize hydrophilic therapeutics while the organic solvent solubilizes the copolymers. Upon homogenization, a W/O/W emulsion is achieved and subsequent stirring followed by filtration yields nanoparticles. Due to the amphiphilic nature of these block copolymers, self-assembling nanoparticles can be designed. As described previously, the hydrophobic blocks of a diblock or triblock copolymer in aqueous solution will isolate themselves from water to form a hydrophobic core, while the hydrophilic blocks will interact with water to form a hydrophilic shell of the nanoparticle (Fig. 13.2). Additionally, the spray-drying technique requires spraying the drug and copolymer solution or suspension in organic solvents into a stream of heated air. This technique produces particles with advantages, such as small size, reproducibility, and milder preparation conditions. However, product loss is a major disadvantage of this method and has been addressed by developing a double-nozzle to simultaneously spray a mannitol solution, which coats the particles and prevents aggregation (Park et al., 2005). The desired properties of nanoparticles require characterization techniques unique to this type of drug carrier. Some common nanoparticle characterization methods include determining the particle size, entrapment efficiency, and zeta potential. These characteristics are easily detected spectroscopically using light diffraction with a particle sizer, by analyzing the drug concentration of disrupted nanoparticles using HPLC or any other established analysis method for drug/protein and electrostatic interactions within a dispersion respectively (Kaszuba et al., 2010; Park et al., 2005).

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FIGURE 13.2 Schematic representation of diblock and triblock (BAB and ABA type) copolymers showing their auto-arrangement to form micelles or nanoparticles in an aqueous system.

13.3.2 RESORBABLE NANOPARTICLES-BASED DRUG DELIVERY SYSTEMS Nanospheres of PLA and PLGA based diblock and multiblock copolymers have been intently studied for the effect of varying polymer composition and theoretical drug loading on particle size and encapsulation efficiency (Peracchia et al., 1997). Nanospheres were obtained using the emulsification/solvent evaporation technique using methylene chloride or a methylene chloride to chloroform ratio of 1:1 as the organic phase in which the copolymer and drug (lidocaine or prednisolone) were dissolved. The organic phase was subsequently poured into an aqueous phase of double distilled water followed by mixing and sonication to produce the desired emulsion. Nanospheres were formed as a precipitate and collected via centrifugation. Purification was performed by washing twice with water followed by lyophilization to obtain the final nanoparticle product. Regardless of drug, increasing the molecular weight of PEG from 5 to 12 to 20 kDa and the theoretical drug loading from 20% to 33% w/w, when prepared using a PEG-PLGA diblock copolymer, showed a slight change in drug loading and particle size. When investigating multiblock (PEG)3-PLA nanoparticles, the size of the particles increased with increasing chain length of PEG. In addition, the encapsulation efficiency was reduced in the multiblock nanoparticles when compared to the diblock nanoparticles. The release profile for the nanoparticles prepared from diblock copolymers showed a typical fast release initially followed by controlled

13.3 Resorbable Nanoparticles

release, in contrast to the nanoparticles prepared using multiblock copolymers, which showed reduced burst release and a drastically slow drug release rate overall. The authors also highlighted the additional influence of an external PEG coat on the characteristics of drug release (Peracchia et al., 1997). In another study, poly(ester-anhydride) copolymer nano- and microspheres were prepared using a double-emulsion/solvent evaporation technique and characterized with additional attention given to the surface modification potential of these spheres. Surfaces were labeled with cystamine, which was quantified by reducing cystamine with Ellman’s reagent (5,5-dithio-bis-(2-nitrobenzoic acid) (DTNB). Microspheres and nanospheres showed labeling of 0.20 35 and 0.6 100 μmol/g respectively. Furthermore, increased labeling was observed with increasing polyanhydride composition (Pfeifer et al., 2005). A pivotal study was done on targeted self-assembling nanoparticles, evaluating their ability to release drugs in a controlled manner to target prostate cancer cells while evading the immune system by utilizing stealth properties (Gu et al., 2008). The nanoparticles were designed using PLGA, PEG, and the A10 aptamer (Apt). Apt targets the nanoparticles to prostate cancer cells that display prostate-specific membrane antigen (PSMA) on their surface enabling the nanoparticles to bind specifically to these cells and, thus, be endocytosed. This study demonstrates how nanoparticles can be tuned to obtain certain enhanced characteristics. The PLGAb-PEG-b-Apt triblock copolymer was engineered to self-assemble in water by hydrophobic forces driving the PLGA to form a hydrophobic core, while the PEG formed the shell with Apt protruding into the aqueous environment. The incorporation of PLGA-PEG at varying ratios provides a means to alter the concentration of Apt, which optimizes the stealth ability conveyed by the PEG moieties to increase circulation time and decrease accumulation in the liver, while retaining the targeting abilities conveyed by the Apt. Additional optimization was achieved through the tunability of PEG with the size of the nanoparticles being directly proportional to the size of PEG used while being independent of PLGA molecular mass. However, when optimizing the release of docetaxel, the opposite effect was observed and raising the PLGA molecular mass, but not PEG, prolonged the release rate. Cell specific targeting was confirmed in vitro using LNCaP cells in comparison to PC3 cells. In vitro and in vivo optimization of increasing Apt density was found to be 5% for maximum tumor targeting, conversely, a reduction in Apt density results in a lower retention within the liver and should be considered to obtain maximum targeting and minimal retention. In another study of a diblock copolymer comprised of polymethacrylate and PLGA was used to deliver plasmid encoding interleukin-10 to prevent autoimmune disease type 1 diabetes (Basarkar and Singh, 2009). In this study, nanoparticles delivered unique payloads and harnessed the benefits of a polymer other than PEG. Specifically, Eudragit E100 was used, which is a copolymer of n-butyl methacrylate, 2-dimethylaminoethyl methacrylate, and methyl methacrylate (1:2:1) with an average molecular weight of 150 kDa. Double emulsion nanoparticle preparation resulted in PLGA-E100 nanoparticles with a zeta potential of

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58.5 6 2.1 and 42.7 6 1.1 mV before and after plasmid loading respectively, which immensely exceeds that of nanoparticles prepared with PLGA alone with a zeta potential of 4.2 6 0.7 and 1.9 6 0.2 mV before and after loading respectively. The size of the nanoparticles was B800 nm and unaffected by composition. To prepare the nanoparticles, phosphate buffered saline was added to dichloromethane containing copolymers and sonicated to produce a primary emulsion. An aqueous solution of the cationic surfactant cetyltrimethylammonium bromide (0.5% w/v) formed the secondary emulsion upon the addition of the primary emulsion followed by homogenization. Evaporation of dichloromethane, centrifugation, and washing allowed for the isolation and finally the lyophilization of the nanoparticles. Plasmid loading and cellular internalization was facilitated by the cationic polymer and an additional benefit of the E100 was proven to be its buffering ability as demonstrated by acid titration. The ability of the nanoparticle to provide buffering inside the endosome is crucial to the protection and endosomal escape of intact plasmid to provide maximum transfection and, hence, protein expression. The destabilization of endosomes due to the presence of E100 was confirmed via monitoring of the endosomal pH, and higher IL-10 levels of expression were observed. In vivo, mice given low-dose streptozotocin (STZ) treatment developed immune infiltration of the pancreas and while inflammation was reduced with passive treatment with IL-10 plasmid or IL-10 loaded PLGA nanoparticles, only the PLGA-E100 nanoparticles loaded with IL-10 showed the ability to completely protect the pancreas. Therefore, harnessing the cationic nature of E100 allowed for the development of nanoparticles capable of condensing plasmid DNA, cellular uptake, and endosomal escape to accomplish transfection and the expression of protein to protect against autoimmune inflammation of the pancreas. Pagar and Vavia (2013) provide another example of a unique nanoparticle design in which L-lactide is polymerized onto the cyclodepsipeptide, cyclo(GlcLeu) (Pagar and Vavia, 2013). An oil in water single emulsion was used to prepare the nanoparticles. The polymer and the drug, rivastigmine tartrate, were dissolved in methylene chloride and added to the aqueous phase containing 0.5% PVA while stirring. Homogenization and then sonication were employed to reduce the droplet size before evaporating the organic phase. The collected nanoparticles were lyophilized and the formulation variables were analyzed for optimization purposes. It was found that increasing the polymer concentration produces lager sized nanoparticles with increased entrapment efficiency. As can be expected, increasing the concentration of PVA in the aqueous phase reduces particle size, but with reduced entrapment efficiency. The authors hypothesized that an increase in entrapment efficiency should occur with an increase in the amount of drug. However, the results stated that increasing the drug amount increased the nanoparticle size, but decreased the entrapment efficiency. This information was utilized and an optimal drug-to-polymer ratio was achieved, which was found to be 1:5, in order to achieve maximum drug entrapment with the smallest particle size. Another parameter to consider is the organic-to-aqueous phase ratio.

13.3 Resorbable Nanoparticles

Table 13.2 Representative Examples of Resorbable Nanoparticle-Based Drug Delivery Systems of Polylactide and Polyglycolide Diblock and Triblock Copolymers Copolymer

Design Attributes and Key Findings

Reference

PEG5K-PLGA PEG12K-PLGA PEG20K-PLGA (PEG5K)3-PLGA (PEG12K)3-PLGA (PEG20K)3-PLGA PLA:PSA

• Minimal effect of varying polymer composition and theoretical drug loading on particle size and encapsulation efficiency. • Observed additional influence of an external PEG coat on the characteristics of drug release.

Peracchia et al. (1997)

• Increasing polyanhydride content allows for increased surface labeling. • Increasing Apt density to 5% was found to be optimum for maximum tumor targeting. • Reduction in Apt density results in a lower retention within the liver. • E100 produced increased IL-10 transfection and expression to protect the pancreas against autoimmune inflammation. • Optimization of drug to polymer ratio will allow for maximum drug loading and minimal size. • Organic: aqueous phase ratio, cryoprotectant used, and sonication time were also found to affect nanoparticle characteristics.

Pfeifer et al., (2005) Gu et al. (2008)

PLGA-b-PEG-b-Apt

PLGA-E100

L-Lactide-

depsipeptide

Basarkar and Singh (2009) Pagar and Vavia (2013)

The size of nanoparticles can be influenced via this parameter and increased particle size was seen with increased organic phase volume. Once again, the optimal ratio for small particle size and high entrapment was 1:5. Other parameters, such as cryoprotectant used and sonication time, can also have an effect on the characteristics of such nanoparticles. Armed with this knowledge, an optimized nanoparticle was produced with a particle size of 142.2 6 21.3 nm and an entrapment efficiency of 60.72% 6 3.72%. Representative examples of resorbable nanoparticle based drug delivery systems of PLA and PGA diblock and triblock copolymers along with their design attributes and key findings are summarized in Table 13.2.

13.3.3 COMMERCIAL AND INVESTIGATIONAL EXAMPLES The first polymeric nanoparticle, Genexol-PM, has reached phase II trials in the United States. Genexol PM is a self-assembling nanoparticle of mPEG-PLA entrapping paclitaxel for the treatment of nonsmall-cell lung cancer, metastatic breast cancer, and gynecologic cancer. Genexol-PM has a particle size of

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23.91 6 0.41 nm and a zeta potential of 28.1 6 3.1 mV. First-order controlled release kinetics were observed in vitro with B65% of entrapped paclitaxel released within 24 hours and 95% released at 48 hours. When examined in nonsmall-cell lung cancer patients, 46.5% were responsive with favorable antitumor activity, but a high frequency of toxicities was reported. Likewise, phase II trials for metastatic breast cancer showed promising results with an overall response rate of 58.5%, however, associated toxicities were frequent. Nevertheless, gynecologic cancer phase I trials are underway. The consensus surrounding this formulation is that the nanoparticles can prevent some exposure of paclitaxel to normal tissue and potentially increase the dosing schedule and duration in comparison to current treatment with paclitaxel. The lower toxicity in this formulation warrants continued research and clinical trials (Lee et al., 2008; Ahn et al., 2014; Werner et al., 2013; Clinicaltrials.gov, 2017a,b). Meanwhile, docetaxel is also being investigated in an mPEG-PLGA nanoparticle formulation, Docetaxel-PNP. Patients with advanced solid malignancies are participating in the determination of the MTD as well as an evaluation of safety and pharmacokinetic profile (Wang et al., 2014; Clinicaltrials.gov, 2017a,b).

13.3.4 LIMITATIONS OF RESORBABLE POLYMERIC NANOPARTICLES As with any delivery system, nanoparticles have some pitfalls that should be taken into consideration. The limitations given previously for PLA, PGA, and PLGA based copolymers can be apparent when used in nanoparticle formulations, but there are also limitations that are unique to nanoparticles. In general, the variations and optimizations that have been discussed are engineered to overcome the potential limitations of these drug carriers. The major limitations include, but are not limited to, obtaining the desired size, entrapment efficiency, stability, and pharmacokinetics. Nanoparticles of small, uniform sizes (,1000 nm, but preferably 50 300 nm) are desirable due to their ability to reach deep tissues via diffusion while evading the immune system. Stealth properties, like PEG coating, will also allow for prolonged circulation by preventing immune system recognition and removal. Entrapment efficiency is extremely important since the efficacy of many drugs are dependent on the amount of drug that reaches its target. Hydrophilic drugs and small molecules can be particularly troublesome in this aspect and researchers frequently inquest to determine whether nanoparticle formulations would be feasible. Stability can vary depending on the copolymers and surfactant used as well as their concentrations. Drug leaking and premature release are also major hurdles when developing nanoparticle formulations. Finally, pharmacokinetics goes hand in hand with stability, which can ultimately influence important factors of the release profile, such as burst release and the duration of sustained release (Olivier, 2005).

References

13.4 CONCLUSIONS AND FUTURE PERSPECTIVES PLA and PGA based diblock and triblock copolymers have been shown as excellent drug/protein delivery carriers for easy administration and controlled drug delivery. The biodegradability and biocompatibility of these copolymer systems has attracted a lot of attention over the years for their wide application optimized for the delivery of both small and large molecules with good safety profiles. The striking potential of these copolymeric delivery systems is the ability to be tailored in relation to the therapeutic incorporated, by increasing or decreasing the hydrophilic/hydrophobic ratio resulting in accelerated or decelerated degradation for shorter or longer duration of drug release and can be further exploited. For longer drug release periods synthesizing a polymer with high degrees of crystallinity can also be considered. Additionally, chemical alterations in the copolymer backbone can be explored for polyelectrolyte complex formation with charged drug molecules in a way that modifies the release pattern or enhances the stability of the delivery system. The incorporation of additives can be potentially tested with the delivery systems for their effect on drug delivery or for a combination therapy approach (Al-Tahami, 2007; Oak and Singh, 2012a). Overall, these copolymers can be formulated into carriers at multiple scales, such as depots, microspheres, nanoparticles, as well as implants. These systems have the ability to incorporate a wide range of therapeutics of diversified intrinsic characteristics for their controlled delivery in a chemically and structurally stable form over varying time periods with different possible routes of administration. (Bret et al., 2011). Further studies will be effective in making this delivery system an ideal approach to administer a large number of protein and peptide based drugs at a controlled rate for a longer duration.

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