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Advances in Ocular Drug Delivery, 2012: 65-86 ISBN: 978-81-308-0490-3 Editor: Ashim K. Mitra

3. Biodegradable polymers for ocular drug delivery Viral Tamboli, Gyan P. Mishra and Ashim K. Mitra

Division of Pharmaceutical Sciences, School of Pharmacy, University of MissouriKansas City, 2464 Charlotte Street, Kansas City, MO 64108-2718, USA

1. Introduction Ophthalmic drug delivery is challenging due to unique anatomy and physiology of the eye. The natural protective mechanisms of the eye render this organ inaccessible to foreign substances and drug molecules. A successful therapeutic treatment requires maintaining the therapeutic drug concentration at the target site by circumventing the anatomical and physiological barriers. The major ocular drug delivery routes include topical, systemic and local administration. Topical mode of administration is the most preferred route to treat anterior segment diseases because of ease of application. However, the ocular bioavailability of topically applied drugs is less than 5% and it is difficult to achieve therapeutic drug concentration at the target site. Poor bioavailability mainly results from the precorneal factors such as blinking, transient residence time in cul-de-sac, and nasolacrimal drainage. In addition, the lipoidal nature of the corneal epithelium restricts the entry of hydrophilic drug molecules and the water- laden stroma acts as a Correspondence/Reprint request: Dr. Ashim K. Mitra, Division of Pharmaceutical Sciences, School of Pharmacy, University of Missouri-Kansas City, 2464 Charlotte Street, Kansas City, MO 64108-2718, USA E-mail: [email protected]

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rate limiting membrane for lipophilic molecules. Moreover, physicochemical properties of a drug entity itself determines the diffusion resistance and relative impermeability offered by various ocular tissues [1,2]. To overcome these barriers and to increase contact time of the drug on the eye surface, absorption enhancers and/or viscosity enhancers are generally used in the ocular formulations. So far, these approaches have limited success to address the problem of poor bioavailability from topical route to treat anterior segment diseases. On the other hand, for the treatment of posterior segment diseases either systemic or local route is preferred because of the poor corneal drug permeation. Systemic administration requires higher dosage and frequent administration that results in severe adverse effects. Local injections, particularly intravitreal and subconjunctival injections are alternate strategies to achieve therapeutic concentration in the vitreo-retinal disorders. However, to maintain the effective concentration repeated injections are required, which causes clinical complications or patient discomfort [3]. Further, the presence of different efflux pumps such as P-glycoprotein, multidrug resistance associated proteins, and breast cancer resistant protein on various ocular tissues restrict the entry of the drug molecules into the eye [4]. Many approaches have been evaluated to improve ophthalmic drug delivery. Application of controlled drug delivery systems was anticipated as an effective approach to circumvent all these limitations. Controlled drug delivery systems release the drug in a sustained and controlled manner by which the therapeutic concentration is maintained for the prolonged period of time. These systems provide many practical advantages: they avoid frequent administration, which is a major noncompliance with many chronic eye disorders. The delivery of emerging therapeutic macromolecules having very short biological half-lives could be possible as these systems protect the protein drugs in situ and have an ability to deliver them at desire rate by overcoming anatomical and biochemical barriers of drug transport [5,6]. These systems can be based on either erodible or nonerodible matrices. In the early 1960s, first polymeric device was developed for controlled drug delivery. Synthetic biodegradable polymers such as poly (glycolic acid) (PGA) and poly (lactic acid) (PLA) had gained attention for biomedical applications. After five-years, poly (lactide-co-glycolide) (PLGA) sutures emerged on the market. Since then, a wide variety of biodegradable polymers were explored for the drug delivery [7]. In the past two decades the development and application of synthetic biodegradable polymers for ocular drug delivery have gained significant momentum. Polymeric devices such as micro and nanoparticles, microspheres, liposomes, hydrogels and ocular implants have been designed to deliver the therapeutic agents in the controlled manner. The release rate of the drug molecules from these polymeric devices depends on many factors such as, molecular weight and degradation

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mechanisms of the polymer, physicochemical properties of the drug, thermodynamic compatibility between the drug and polymer and the shape and size of the devices [8]. In this chapter, we summarize the properties of natural and synthetic biodegradable polymers and their applications in the ocular drug delivery.

2. Biodegradable polymers Biodegradation is an enzymatic or non-enzymatic hydrolysis of the polymeric backbone into water soluble or insoluble products. Biodegradation involves two complementary processes, degradation and erosion. In the degradation process cleavage of the polymeric backbone into low molecular weight fractions takes place, whereas the erosion mechanism refers to the physical phenomena such as dissolution and diffusion of low molecular weight fractions from the polymer matrix. The degradation products are eventually eliminated from the body via normal metabolic pathway [9].

2.1. Types of biodegradation Heller has described three basic mechanisms of polymer degradation and classified the polymers based on the degradation mechanisms [10]. Schematic representation of polymer degradation mechanisms is shown in Fig. 1 [10].

Figure 1. Schematic representation of polymer degradation mechanisms.

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Type I: Crosslinked water soluble polymers generally follow type-I erosion. Polymers such as gelatin, collagen, polyacrylamides, poly (vinyl alcohol) (PVA) and poly (N-vinyl pyrrolidone) (PVP) upon crosslinking form hydrogel, which is a water insoluble three dimensional structure that undergoes type I hydrolysis. On the basis of hydrolysis product generated, type I erosion mechanism can be further subdivided into type IA and IB. Type IA erosion mechanism produces high molecular weight water soluble polymers, whereas type IB generates low molecular weight polymers. Polymers having type IA erosion kinetics are best suited for topical applications because of faster elimination of high molecular weight water soluble polymers from the ocular surface. Polymers following type IB degradation kinetics are generally utilized for designing implants. Polymeric systems that undergo type I erosion are highly water permeable therefore; they are not suitable for the delivery of low molecular weight compounds with appreciable water solubility. However, the crosslinked polymeric matrix physically entangles the macromolecules and restricts them to diffuse out of the matrix. Therefore, these polymers are well suited for the delivery of macromolecules such as enzymes and antigens or sparingly water soluble molecules, which are released from hydrogels initially via diffusion followed by degradation of the polymer [10]. Type II: Conversion of water insoluble linear polymers into water soluble moiety through hydrolysis, ionization, or protonation of pendant groups is defined as type II erosion. However, since no backbone degradation is involved during erosion process overall molecular weight of polymers does not change significantly. These polymers are generally employed for the topical applications. Copolymers of alkyl vinyl ether and maleic anhydride follow type II erosion mechanism where the degradation rate of the copolymers is affected by the size of the alkyl substitute, pH of the degradation medium and pKa of the carboxylic group [11]. Type III: Type III erosion produces low molecular weight water soluble molecules by the hydrolytic cleavage of water insoluble high molecular weight polymers. The polymers demonstrating type III erosion kinetics produce non-toxic degradation products and thus advantageous for topical and systemic administrations. These polymers are employed for wound healing after surgery and also for chronic ocular diseases. These polymers are available in a wide range of molecular weights having different physico-chemical properties, which can be modulated to formulate drug delivery systems. PLA, PGA and their copolymers PLGA, polyanhydrides, polyurathanes, polycaprolactone (PCL) and its copolymers, poly (ortho esters) and poly (alkyl

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cynoacrylates) (PACA) exhibit type III erosion mechanisms because of the characteristic hydrolytic instability in the polymer backbone [3]. The representative structures of these polymers are shown in Fig. 2. CN O *

C

CH2

COOR

*

*

R

C

OR'

O O

C

*

O

C

O

R

R''

n

Poly(cyanoacrylates)

*

*

n

Poly(anhydrides)

n

Poly(ortho esters)

O * *

O

n

Polyester

Figure 2. Representative structures of different biodegradable polymers.

2.2. Advantages of biodegradable polymers Biodegradable polymers offer several advantages over non-biodegradable polymers for controlled drug delivery. They do not require surgical removal after application, being the most important advantage in ophthalmic drug delivery as it can circumvent surgical complications associated with nonbiodegradable implanted devices. The natural and synthetic biodegradable polymers have many favorable properties such as biocompatibility with ocular tissues, biodegradability and mechanical strength. They provide negligible toxicity and also their degradation products are non-toxic in terms of both local and systemic response. Due to the adequate mechanical properties, they can be tailored to wide range of properties. Natural biodegradable polymers such as gelatin, albumin, chitosan, hyaluronic acid and synthetic biodegradable polymers such as PVP, PACA, PCL, PEO, polyanhydrides and thermoplastic aliphatic polyesters like PLA, PGA and PLGA have been thoroughly explored for ocular delivery systems as summarized in Table 1. These polymers are approved by FDA for human applications.

3. Natural biodegradable polymers used for ocular drug delivery 3.1. Gelatin Gelatin is a natural polymer, derived from the native protein collagen. Due to its biocompatibility and biodegradability gelatin is widely utilized in pharmaceutical and medical applications. Commercially gelatin is available

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Table 1. Polymeric drug delivery systems used in ophthalmic research.

in two types; designated as type A and type B. Gelatin type A is the acid processed collagen, whereas gelatin type B is derived from alkaline treated collagen. These processed gelatins have different isoelectric points, 7-9 for type A and 4-5 for type B. An absorbable cross-linked hydrogel of gelatin can be easily prepared by thermal and chemical modification. This treated hydrogel does not dissolve in water rather it swells upon contact with water [22]. Gelatin based ocular delivery systems can be optimized by changing the electrical and physical properties of gelatin [23]. For example, an aqueous solution of gelatin acts as a thermoreversible hydrogel. This hydrogel has low mechanical strength and starts to break at 30 0C that result in drug loss. Gelatin chains can be crosslinked chemically to provide stability toward thermal degradation. Natu et al. characterized the carbodiimide crosslinked gelatin hydrogel loaded with pilocarpine hydrochloride for the treatment of glaucoma. They observed that the water uptake of hydrogel was reduced by crosslinking, which resulted in slower release of pilocarpine hydrochloride from the gel matrix. In addition, they found that the degree of crosslinking played a major role in both mechanisms. They also prepared the lyophilisates of the gelatine and observed the similar effects of crosslinking [24]. In another study, release of pilocarpine was optimized by embedding the gelatine hydrogel with cetylester wax and polyethylene glycol. This strategy

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had provided zero order release of pilocarpine from the hydrogel due to slower penetration of water inside the matrix. [25]. Vandervoort et al. prepared gelatin nanoparticles encapsulating pilocarpine HCl and hydrocortisone for topical ophthalmic applications. These investigators examined the effect of various parameters (such as gelatin type and pH) on the preparation of nanoparticles. They concluded that gelatin could be an effective polymeric carrier due to prolonged residence time at the ocular surface [26]. Gelatin was extensively employed for fabrication of ocular devices such as microspheres [27] and microcapsules [28].

3.2. Collagen Collagen is an important biodegradable and biocompatible natural polymer widely distributed in ocular tissues such as the sclera, stroma and cornea. Different types of collagens are identified in human body, which are generally 300 nm long having molecular weight of 300,000. Collagen type I is the most common and thoroughly explored protein present in mammals. Collagen derived from the animal sources i.e. porcine and bovine could be utilized for human applications, because of secondary and tertiary structural similarity with the human collagen. It primarily erodes by both type I and type II mechanisms. In vivo degradation primarily occurs with collagenases and metallo proteinases enzymes, which give corresponding amino acids. The rate of degradation can be modulated by various cross-linkers and the enzymatic treatment. It possesses characteristic physico –chemical properties, which could be optimized for different delivery systems such as sponges, nanofibrous matrices, foams, powders, dispersions and injectable viscous solutions [29]. Collagen is widely utilized to prepare collagen corneal shields because of excellent biocompatibility, absorption capacity and low immunogenicity [30,31]. These corneal shields enhance the drug bioavailability by acting as a drug reservoir and providing prolonged contact time. In addition, these systems are applicable for the dry eye syndrome due to its capability to form a gel like structure on the arid eye surface. Shaker et al. evaluated the efficacy of collagen disks made up of porcine scleral tissue for the treatment of dry eye syndrome. In this study, authors observed that collagen shields upon derivatization with succinic acid were well tolerated by dry eye patients. These films have provided enhanced tear film stability and resulted in statistically significant symptomatic relief by reducing the frequency of artificial tear administrations [32]. However, collagen shields in the form of contact lenses interfere with vision and resulted in patient discomfort [33]. Gebhardt et al. evaluated collagen particles to deliver cyclosporine to the cornea and found that particles were equally effective as

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collagen shields in inhibiting corneal allograft rejection [34]. These polymeric systems were also used to deliver antifungal, antiviral, antiinflammation, antibacterial, immunosuppressive and anticoagulant agents [31].

3.3. Chitosan Chitosan (CS) is a natural polymer derived from chitin that upon deacetylation produces chitosan. Chitin constitutes exoskeleton of arthropod and primarily composed of D-glucosamine conjugated through 1-4 linkage [29]. Chemical structure of CS is shown in Fig. 3. In vitro degradation of CS occurs through enzymes such as papain, lysozyme and chitosanase [35], whereas in vivo degradation primarily occurs through lysozyme enzyme. The rate of degradation is inversely proportional to the degree of acetylation and polymer crystallinity [36]. CS possesses various physical properties such as mucoadhesivness and penetration enhancing nature, which render its applicability for ocular applications. Its mucoadhesive property is attributed to the positive charge of amino groups, which interacts with the negatively charged ocular tissues such as the cornea and sclera. Its mucoadhesive nature is advantageous for increasing the ocular bioavailability [37]. Polymers such as PACA and PCL have shorter residence time on the cornea that results in poor absorption across the lipophilic corneal barriers [38]. In addition, these polymers permeate the corneal epithelium mainly via paracellular pathway [39]. In contrast, characteristic pseudoelastic and viscoelastic rheological properties of CS permit prolonged residence time on the cornea. Both CS nanoparticles and CS-coated nanocarriers exhibited high affinity for the corneal and conjunctival epithelial cells [37,40]. CS improves the permeability of molecules by modulating the tight junctions of the corneal epithelial cells in a reversible manner. Reports suggest that CS alters the paracellular and transcellular pathway without disrupting the cellular integrity [14,41]. In a recent review, Alonso et al. summarized the applications of CS nanoparticles as ocular drug delivery vehicle [39]. The effect of CS nanoparticles and CS solution was compared by numerous researchers. It was observed that the ocular bioavailability of CS nanoparticles was 2 to 10 fold higher than suspension of drug in a CS solution [41]. Similar results were observed in the case of fluorescein labeled CS solution and fluorescein loaded nanoparticles [39]. Calvo et al. compared the effect of CS and poly-L-lysine (PLL) coating on the ocular bioavailability of poly- ε -caprolactone nanocapsules and reported that the levels of encapsulated indomethacin is significantly higher in the cornea and aqueous humor as compared to the eye drops. The authors concluded that although

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HO

NH2

O HO

O

OH O

HO

NH2 HO

Figure 3. Chemical structure of Chitosan.

both CS and PLL are cationic polymers, positive charge is not responsible for enhanced penetration of nanocarriers but unique physical properties of CS is responsible for increased intake of CS-coated nanocapsules [42]. Similarly, CS-coated nanocapsules were found more effective than PEG- coated nanocapsules. PEG accelerates the penetration of a colloidal carrier throughout the epithelium, whereas CS enhances the retention of the nanocapsules in the superficial layers of the epithelium [40].

4. Synthetic biodegradable polymers used for ocular drug delivery 4.1. Poly N-vinylpyrrolidone (PVP) PVP is a synthetic and biocompatible polymer widely utilized for vitreoretinal drug delivery. It is mainly employed for the preparation of hydrogels that exhibit viscoelastic properties [43]. The decomposition products of PVPbased hydrogels are easily eliminated from the vitreous through phagocytosis [44-46]. Hydrogel prepared from cross linked PVP was used as a vitreous substitute [47]. Hong et al. evaluated biodegradation of poly (1-vinyl-2pyrrolidinone) cross-linked with 1% 14C-methyl methacrylate. They observed that cross-linked PVP hydrogel did not degrade in vitro in presence of proteolytic enzymes such as trypsin or collagenase. However, in vivo half of the hydrogel disappeared from the rabbit vitreous cavity within 4 weeks by phagocytosis [44]. PVP based hydrogels were transparent materials and remain at the site of injection for several weeks. However, fragmentation of the hydrogels triggers an inflammatory response resulting in the vacuole formation in the retinal pigment epithelium [46]. In addition, clinical studies have shown that PVP-based hydrogels cause intravitreal opacity, hazy corneas and inflammation and might not be suitable as vitreous substitutes [47]. Degradation kinetics of this biomaterial could be easily modulated by varying crosslinking density. Niu et al. investigated injectable hydrogel of acrylamide/N-vinylpyrrolidone copolymer crosslinked with reversible

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disulfide bond for ophthalmic applications. This hydrogel showed characteristic in-situ sol-gel transition that facilitated the designing of complex shapes, which was advantageous as artificial vitreous substance and scaffold for lens regeneration [48]. Hacker et al. explored the matrices composed of photocrosslinked poly (propylenefumarate) (PPF)/ (PVP) for a long term delivery of antiglaucoma drugs, such as acetazolamide (AZ), dichlorphenamide (DP) and timolol maleate (TM). Authors suggested that the use of PVP based implants could be a valuable strategy for controlled release of drugs over a period of 300 days in glaucoma therapy [49].

4.2. Poly (lacticide) (PLA), poly (glycolide), and their copolymers polylactide-co- glycolide (PLGA) PLA and PLGA are the most promising biodegradable polymers [50]. PGA alone is highly prone to hydrolysis and remains insoluble in common organic solvents therefore it is not widely acceptable for the fabrication of controlled drug delivery systems. PLA alone and in combination with PGA with different ratios are mostly utilized in the formulations. These polymers are synthesized by two methods. First involves direct condensation reaction of monomers, which results in low molecular weight polymers and the other method is based on ring opening polymerization of cyclic dimmers, which yields high molecular weight polymers. These polymers upon non-enzymatic or enzymatic hydrolysis produce water soluble metabolic products, which are not harmful to living tissues [51,52]. These polymers belong to polyester class and degrade mainly through bulk erosion. In vitro degradation of polyesters primarily occurs through hydrolytic cleavage. However, in vivo, enzymes play an important role to initiate the degradation process. The degradation products lactic acid and glycolic acid are nontoxic and eliminate in the form of CO 2 and water via Krebs cycle [53]. Chemical structures of different polymers of polyester class are shown in Fig. 4. Polymer degradation rate can be easily modulated by changing the molecular weight, composition, conformation and crystallinity of the polymers [54]. For example, by varying the ratio of lactide and glycolide a wide range of diffusion and degradation profiles can be obtained in PLGAs. PLGA with 50:50 ratio of lactic and glycolic acid degrades faster than either PLA or PGA alone [55,56]. The presence of methyl group provides more hydrophobicty to PLA and it degrades slowly in comparison to PGA. These polymers have glass transition temperature ranging from 45 to 65 °C [57]. Physico-chemical properties of different grades of PLGAs and other polymers are summarized in Table 2.

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Figure 4. Chemical structures of different polymers of polyester class. Table 2. Physico-chemical properties of polymers.

Drug release from the PLGA system depends on the proportion of two monomer used, porosity, surface area of the carrier and physico-chemical properties of the incorporated drug [58,59]. These polymeric materials have been used as surgical sutures due to their good biocompatibility and rapid clearance [60]. PLA and PLGA are widely utilized in ocular drug delivery

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systems such as implants, injectable microspheres and nanoparticles. PLA and PLGA microspheres have been evaluated to reduce the intravitreal administration frequency for various chronic eye diseases such as cytomegalovirus retinitis and endophthalmitis [61,62]. Dillen et al. developed cationic Eudragit® coated PLGA nanoparticles loaded with ciprofloxacin, a most commonly used fluoroquinolone for ocular infections. These authors found that positively charged drug loaded nanoparticles can adhere to the negatively charged bacterial surface. In addition, particulate systems enhanced the therapeutic drug concentration at the target site by providing prolonged diffusion controlled release [63]. Kunou et al. achieved pseudozero-order release kinetics of GCV over a period of one year by employing two monomers of PLA with different molecular weights and ratios for the preparation of biodegradable scleral implant [15]. The PEGcoated PLA nanospheres were more efficient for sustaining the drug release and improving the ocular bioavailability of ACV in the treatment of viral infections [64]. PLGA was also utilized to encapsulate anti-VEGF RNA aptamer (EYE001) in microspheres. In contrast to a characteristic triphasic release pattern of microsphere, the release of EYE001 from PLGA microspheres was a diffusion-controlled process that exhibited drug release in continuous manner over a period of 20 days. Furthermore, the bioactivity of the aptamer was retained in the formulation during the entire release period [65]. Duvvuri et al. discussed the conventional triphasic release pattern from PLGA microspheres and optimized the drug release kinetics by employing various PLGA polymer blends [66]. However, particulate systems such as microspheres and nanospheres may cause vision obstruction or irritation to the retinal tissues after intravitreal injections. In addition, most of the PLGA based drug delivery systems have initial burst release phase. Authors investigated the composite approach to minimize the particulate system related drawbacks. This dual approach involved the use of PLGA-PEGPLGA triblock thermogelling polymer to suspend the particulate system [67, 68]. The thermosensitive polymer exists in the liquid state at room temperature and forms gel upon contact with eye tissue i.e. 34 0C. Thermosensitive hydrogel holds particles at the site of administration and avoids vision interference. In addition, gel matrix protects the microspheres from enzymatic and cellular degradation. This dual system showed release of drug in a more controlled manner for prolonged period of time [61].

4.3. Poly- ε-caprolactone (PCL) PCL is an aliphatic polyester synthesized form monomer ε- caprolactone through ring opening polymerization catalyzed by stannous octoate at 140 0C

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[69]. It is a tough semi crystalline polymer having the melting point in the range of 59 and 64 °C and a glass transition temperature of -60 °C [70]. Permeability and crystallinity of the PCL can be modified by copolymerization with PLA or PGA [69]. Degradation of PCL occurs in two phases. First phase involves molecular weight (Mn) loss up to 5000 due to cleavage of ester linkage in the polymer backbone (chain scission), that produces ε- hydroxyl caproic acid and decreases the intrinsic viscosity of polymer. In the second phase (commonly observed in vivo), chain scission of low molecular weight polymer produces small fragments, which diffuse out of the polymer bulk and break the polymer in small particles that undergo phagocytosis [59]. PCL is utilized for sustained drug delivery due to its higher permeability to various drug molecules and slower degradation in comparison to other polymers [70]. Degradation rate of PCL can be improved by co-polymerizing with other fast degradating polymers. PCL implant loaded with dexamethasone had released the drug within the therapeutic range over the period of more than one year and was well tolerated in the rabbit eye [71]. Rod shaped PCL implant loaded with triamcinolone was also well tolerated in sub-retinal space of rabbit eye and had released the drug over the period of 4 weeks without any clinical complications [72]. Yenice et al. evaluated hyaluronic acid coated PCL nanospheres loaded with cyclosporine. Investigators found that bioavailability of cyclosporine nanospheres was 10-15 fold higher than the drug solution in castor oil. PCL can be utilized to prepare in situ gel-forming sustained drug delivery system. PCL based triblock polymer was recently characterized for ophthalmic applications. Gong et al. evaluated the toxicity of PEG-PCL-PEG triblock copolymer hydrogel after intracameral injections. This hydrogel was biocompatible with ocular tissues and appeared to be a promising controlled release systems for chronic ocular diseases [73].

4.4. Poly (alkyl cynoacrylates) (PACA) PACA is synthesized from monomer alkyl cynoacrylate, which exhibits bioadhesive properties. It can form a strong bond with polar surfaces including skin and living tissues. Polymethylacrylate composed of smaller alkyl chain is not applicable to drug delivery due to tissue toxicity and inflammation. Therefore, larger alkyl chains such as n-butyl, octyl cyanoacrylates are used for clinical applications. In practice anionic or zwitterionic polymerization are commonly used for synthesis of PACA due to rapid initiation at ambient temperature. Polymer degradation occurs by enzymatic hydrolysis of alkyl side chain producing an alkyl alcohol and poly (cyanoacrylic acid). The degradation products are soluble in water and

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eliminate via kidney filtration [74]. This polymer was mostly explored for the preparation of biodegradable nanoparticles. Layre et al. reported the encapsulation of highly crystalline alkylating drug, busulfan, utilizing five different PACA polymers. The highest encapsulation of busulfan was found in poly (isobutyl cyanoacrylate) (PIBCA) and poly (ethyl cyanoacrylate) due to the specific interaction between the drug and the polymers. They suggested that this nanoparticulate formulation when given intravenously have nigligible toxicity and also minimize the variability in bioavailability [75]. Peracchia et al. suggested the potential use of PEG-coated PIBCA nanoparticles as drug delivery carriers, which are rapidly biodegradable. The covalently bound PEG avoids interaction with blood components and prevents recognition by macrophages of the mononuclear phagocyte system after intravenous injection [76]. PEG- coated polyethyl-2-cyanoacrylate nanospheres had increased the ocular bioavailability of ACV by 25 fold when instilled in the conjunctival sac of rabbit eyes. The improved drug bioavailability was attributed to the colloidal nature of nanospheres that can facilitate the transport of drug paracellularly. In addition, presence of PEG provided better mucoadhesion on the corneal surface and improved dug permeation [77].

4.5. Polyanhydrides In 1930s, Hill and Carothers proposed polyanhydrides as substitutes of polyesters for textile applications. However, they were not useful for textile industry due to faster hydrolytic cleavage of anhydride linkage in the polymer backbone. Instead, due to this intrinsic property polyanhydrides are considered as an ideal candidate for formulation of controlled drug delivery systems. Hydrolytically labile linkages of polyanhydrides provide biodegradability and regulate degradation rate. For example, poly [bis (pcarboxyphenoxy) alkane anhydrides] degradation rate can be adjusted from 10-1 to 10-4 mg/hr/cm2 upon changing the methyl group to the hexyl group. Polyanhydrides can be synthesized by three methods: melt condensation, dehydrochlorination and dehydrative coupling. Melt condensation method can produce high molecular weight polymer (up to 50,000) while other two methods are useful for the synthesis of low molecular weight polymers [78]. Polyanhydrides show pH dependent degradation, which can be modulated by additives. Basic additives primarily promote bulk erosion, whereas acidic additives favor surface erosion and produces acetic acid upon degradation [79]. Further, upon changing the polymeric backbone drug release rate can be modulated over a thousand fold [80]. Mostly copolymer of bis(pcarboxyphenoxy propane) and sebacic acid is utilized for drug delivery applications. Chemical structure of this polymer is shown in Fig. 5.

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Figure 5. Chemical structure of Poly [bis (p-carboxyphenoxy) propane-co-sebacic acid.

Release of drug from this polymeric delivery system occurs mainly by surface erosion rather than drug diffusion. It has a potential to provide almost zero-order drug release rate and also undergoes relatively faster in-vivo biodegradation [80]. Rosen et al. demonstrated near zero order degradation and drug release kinetics from poly [bis (p-carboxyphenoxy methane andride] for several months at two different temperatures i.e. 37 and 60˚ C [81]. Polyanhydride microspheres have been employed to avoid repeated intravitreal injections for the treatment of vitreoretinal diseases [80]. Microspheres prepared from poly (adipic anhydride) (PAA) exhibited surface degradation. Release of timolol maleate from these microspheres was sustained for 7 hrs and mainly controlled by polymer degradation. Further, to improve ocular bioavailability of timolol maleate PAA-microspheres were suspended in the Gelrite@ (an in situ polysaccharide gel) [82]. In another study by Lee et al., 5-fluorouracil (5-FU) was incorporated in 3mm bioerodible disc of bis (p-carboxyphenoxy) propane and sebacic acid. They reported that 5-FU was delivered in a sustained manner and maintained intraocular pressure for 3 weeks [83].

4.6. Poly (orthoester) (PEOs) Since early 1970s, four families of poly (ortho esters) have been synthesized. POEs are hydrophobic polymers having hydrolytically labile ortho ester bonds. The amount of water available to react with these bonds is very less under physiological conditions that make the polymer extremely stable. POEs undergo surface erosion and provide zero order release rate for a longer period of time [84]. POEs based formulations have been proven promising in the treatment of ocular diseases such as glaucoma filtration surgery and proliferative vitreoretinopathy (PVR) [85]. They can be injected directly into the eye with a needle of appropriate size. Chemical structures of four types of POEs are shown in Fig. 6 [87]. POE I is synthesized by transesterification reaction between a diol and diethoxy tetrahydrofuran. It is a hydrophobic solid polymer having acid sensitive nature and easily hydrolyzed in an aqueous environment. Basic

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Figure 6. Chemical structures of different Poly (ortho esters).

ingredients such as sodium carbonate are generally utilized to prevent autocatalytic hydrolysis. This polymer has been widely explored for orthopedic applications, treatment of burns and for the delivery of narcotic antagonist and contraceptive steroids. It is not explored much for ocular drug delivery [85-87]. POE II is synthesized by simple addition reaction between diol and diketeneacetal 3,9-di(ethylidene 2,4,8,10-tetraoxaspiro[5.5]undecane) [86]. Monomers are required to dissolve in tetrahydrofuran and trace of acidic catalyst is used to initiate polymer synthesis instantaneously. Polymer hydrolysis occurs in two steps, unlike POE I there is an absence of autocatalytic hydrolysis. This polymer is also synthesized by crosslinking a triol, either alone or as a mixture with diols. It forms a dense polymer upon crosslinking, which biodegrades to small water soluble fragments. The cross linked density can be adjusted by varying the ratio of diol to triol. POE II can be fabricated as a hard glassy material to semi solid material; mechanical and thermal properties can be controlled by using diols having different degrees of chain flexibility [86, 87]. POE II have been extensively explored for the release of 5-FU, which is mainly utilized as an adjunct to glaucoma filtration surgery. The erosion rate of POE II can be controlled by incorporating the acidic excipients such as suberic, adipic and itaconic acids in the polymer matrix. Nearly zero–order release of 5-FU was obtained by incorporating different amount of suberic acid in POE polymeric matrix [84]. According to the United States of Pharmacopoeia, this generation of polymer is nontoxic for cellular, subcutaneous, intramuscular and systemic implant applications.

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POE III is semi-solid at room temperature and synthesized via transesterification of 1, 2, 6- and trimethyl orthoacetate [87]. This polymer has a very flexible backbone and allows incorporation of therapeutic agents at room temperature without using organic solvent. Therefore, it can be used for thermo labile and solvent sensitive drugs. The release rate of incorporated drug can be controlled by modulating the molecular weight of the polymer. Merkli et al. observed that the release of 5-FU occurred within 1 day from 3500 Da and was sustained for 1 week from 33,300 Da POE III [88]. Drug delivery systems fabricated from this polymeric material do not show any burst release and the drug release rate was governed by the polymer degradation rate [89]. Sintzel et al. reported that the drug release rate from POE III can be controlled by modulating the hydrophobicity of the polymer by substituting triol from 1,2,6 hexanetriol to 1,2,10- decanetriol [90]. According to Einmahl et al., this new generation of POE has a potential for application in glaucoma filtering surgery for the patients with higher risk of surgery failure. This injectable polymer can provide sustained release of 5-FU for 2 weeks after subconjunctival injection that can avoid frequent administrations and minimize the adverse effects [89]. These authors have also described that after subconjunctival administration, polymer degradation products follow several pathways. One major pathway involves direct entry into the anterior chamber through the fistula, to the ciliary body, into the vitreous body, and then into the retina [84]. Therefore they evaluated the biocompatibility of this polymer in different parts of the eye including anterior chamber and suprachoroidal space. They found that the anterior chamber of the rabbit eye can tolerate up to 50 µl of polymer solution, which degrades within 1 week [91]. In addition, after suprachoroidal injection the retinal pigmented epithelial (RPE) cells, retinal and choroidal vasculatures were not affected by the polymeric formulation [91, 92]. POE III demonstrated excellent biocompatibility to the different parts of the rabbit eye. Difficulties in synthesis and lack of reproducibility have limited the use of POE III in biomedical applications [87]. POE IV is synthesized by reacting diols with the diketeneacetal 3, 9diethylidene-2, 4, 8, 10-tetraoxaspiro [5.5] undecane. It is a modified form of POE II, which contains latent acid in the polymer backbone that regulates the erosion rate. The latent acid is generally composed of glycolic acid or lactic acid. POE IV does not require external acidic excipients to control the erosion rate, unlike POE II. When the polymer is exposed to an aqueous solution, the latent acid will hydrolyze to give lactic acid or glycolic acid that will further assist in the hydrolysis of polymer. POE IV can be fabricated as solid or gel-like material by changing the nature of diols. POE IV-based

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devices generally undergo surface erosion and produce acidic degradation products which readily diffuse out from the device. Lactic acid based fourth generation POEs are biocompatible and have long residence times following intracameral, subconjunctival, intravitreal and suprachoroidal injections in the rabbit eyes [84]. Polak et al. evaluated the efficacy of 5-chlorouracil (5CU) loaded POE IV formulation in the glaucoma filtration surgery. They found that 5-CU suspended in POE IV has maintained low IOP in the rabbit eye for 5 months [21].

Concluding remarks Polymeric materials contribute a significant role in the controlled drug delivery. In particular, biodegradable polymers have been extensively explored for ocular therapeutics in the recent years. In this chapter we have summarized mainly the properties and applications of biodegradable polymers having natural and synthetic origins. We have exemplified the applications of biodegradable polymers for the delivery small molecules to the different parts of the eye. Two major advantages of polymeric drug delivery devices, enhancing drug bioavailability and minimizing side effects, are significant in ocular drug delivery. The development of new biodegradable block polymers has gained significant momentum in the recent years. These polymers would be advantageous in the delivery of newer therapeutic agents including genes, therapeutic antibodies and bioactive proteins. It is challenging to deliver these macromolecules to targeted tissues of the eye. Therefore, design of novel biodegradable polymeric devices is currently under investigations for targeted delivery of macromolecules.

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