of intraocular pressure. A recently completed randomized, single-blind, paralleldesign, phase 2 study evaluated the efficacy and safety of the bimatoprost punctal plug in open-angle glaucoma and ocular hypertensive patients (clinicaltrials.gov ID NCT00824720); the results have not been published.

9.3  Medical Applications for Biodegradable Polymers

Biodegradable polymers have been intensively investigated as surgical biomaterials and in the formulation of drug delivery devices. These materials have included both synthetic polymers, such as polyester derivatives, and natural polymers (biopolymers), such as bovine serum albumin, human serum albumin, collagen, gelatin, chitosan, and hemoglobin (Jain 2000; Gorle and Gattani 2009). Synthetic biodegradable polymers have been successfully used as biomaterials for several decades, being employed initially to produce absorbable sutures, fixation devices for orthopedic surgery (e.g., bone pins and screws), and stents (see Sect. 9.4). Biodegradable polymers can be utilized in a variety of ways to increase the residence time of drugs, slow drug clearance, and enhance drug absorption in the eye. Polymer-based biodegradable ocular drug delivery systems investigated to date include implantable sustained-release drug pellets and inserts, viscosity enhancers, mucoadhesive agents, drug-releasing contact lenses, and injectable formulationssuchashydrogelsandliposomesaswellasmicroemulsions,microsuspensions, microspheres, microcapsules, and their nanoscale counterparts (see Conway 2008; Ghate and Edelhauser 2006; Kearns and Williams 2009; Wadhwa et al. 2009; Gaudana et al. 2009 for review).

Biodegradable polymer-based drug delivery systems show considerable promise for the treatment of ocular diseases and offer a potential solution to many of the limitations of conventional (i.e., systemic, oral, and topical) methods for the administration of ophthalmic drugs (Table 9.3), particularly for treating sight-threatening retinal diseases (Yasukawa et al. 2006; Ghate and Edelhauser 2008; Wadhwa et al. 2009; Gaudana et al. 2009). Biodegradable polymer-based drug delivery can be used to achieve prolonged therapeutic drug concentrations in ocular target tissues, such as the retina, that are not readily accessible by conventional means, and with drugs that may be poorly absorbed by other routes of administration. Drugs formulated with these polymers can be released in a controlled manner, in which the drug concentration in the target tissue is maintained within the therapeutic range (Park et al. 2005), thereby avoiding both insufficient efficacy and high peak drug concentrations associated with pulsed dosing. The tissue specificity of biodegradable drug delivery systems can potentially limit side effects that may otherwise occur with systemic drug exposure. The polymers used in biodegradable drug delivery systems­

are biocompatible, undergoing biodegradation to nontoxic metabolites or polymers that solubilize in vivo and can be eliminated safely by endogenous metabolic ­pathways in the human body without eliciting permanent, chronic foreign-body

reactions (Chu 2008). They also provide sufficient mechanical strength to withstand physical stress in vivo, which enables structural integrity to be maintained for a sufficient duration during therapy. The release rates of drugs from biodegradable systems can be manipulated by choosing polymers with the desired biodegradation kinetics, physicochemical properties, and thermodynamic characteristics, as well as by varying the shape of the delivery system (Park et al. 2005). Biodegradable delivery systems with a long duration of action eliminate the need for frequent retreatment, as occurs with intravitreally injected drugs; unlike nonbiodegradable systems, they do not require removal once the drug supply is exhausted, thereby eliminating the risks associated with such procedures (Conway 2008; Kimura and Ogura 2001). Lastly, such devices remedy the issue of poor compliance with medication treatment regimens, which is problematic with orally and topically administered ocular medications.

Biodegradable devices are most often constructed from synthetic aliphatic polyesters of the poly-a-hydroxy acid family, which include poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and the PGA/PLA copolymer poly(lactic-co-glycolic acid) (PLGA) (Fig. 9.1). Other biodegradable polymers used as biomaterials include poly(e-caprolactone) (PCL), poly(glycolide-co-lactide-co-caprolactone) (PGLC), poly(ortho esters) (POE), polyanhydrides (PAH), polymethylidene malonate (PMM), polypropylene fumarate (PPF), and poly(N-vinyl pyrrolidone) (PVP). These polymers have been explored for ocular drug delivery systems, but the technologies are still in the early stages of development and none have been marketed commercially (see Sect. 9.6).

Fig. 9.1  Chemical structures of biodegradable polymers commonly used in biomedical applications

9.3.1  Polylactic Acid, Polyglycolic Acid,

and Polylactic-Co-Glycolic Acid

The aliphatic poly-a-hydroxy acids – PLA, PGA, and their copolymer PLGA – are the most widely studied of the synthetic biodegradable polymers. This family of polymers shows suitable biocompatibility and biodegradability and is versatile for a range of biomedical applications. Polymers from this family have been approved by the US FDA for drug delivery use (Jain et al. 1998).

The discovery and synthesis of lactideand glycolide polymers was first reported several decades ago, and during the late 1960s and early 1970s their application as suture biomaterials was first described (Jain 2000). The polymers showed ­several useful characteristics such as good mechanical properties, low immunogenicity and toxicity, excellent biocompatibility, and predictable biodegradation kinetics (Jain 2000). The widespread use of lactide/glycolide polymers as suture materials led to interest in their use for other biomedical applications – initially as ­orthopedic

­surgical biomaterials (e.g., bone screws and pins) and later for drug delivery systems (e.g., implants, hydrogels, micro/nanoparticles).

PGA is a semicrystalline polymer with a high melting temperature and low solubility (Bowland et al. 2008). PGA is synthesized using toxic solvents (hexafluoroisopropanol or hexafluoroacetone sesquinhydrate) which can limit its potential as a biomaterial for drug delivery, since drugs could react with residual solvent, or trace amounts of solvent could be incorporated into the end product (Bowland et al. 2008). PLA is a hydrophobic synthetic polyester, which, unlike PGA, can exist as an optically active stereoregular polymer (L-PLA aka PLLA) and an optically inactive racemic polymer (D,L-PLA aka PDLLA) (Jain et al. 1998; Chu 2008). PLLA is semicrystalline because of the high regularity of its polymer chain and has an extremely slow biodegradation rate in vivo, while the amorphous nature of PDLLA results from irregularities in its polymer chain structure. Therefore, PLLA has higher mechanical strength than noncrystalline PDLLA and is used mainly for surgical fixation devices such as pins and screws, while PDLLA is preferable over PLLA for drug delivery because PDLLA enables more homogeneous dispersion of the drug in the polymer matrix (Jain et al. 1998). In comparison with PGA, PLA is more hydrophobic and degrades more slowly due to the presence of methyl side groups.

PLGA, a copolymer of PLA and PGA, is the most widely utilized biodegradable polymer for drug delivery. PLGA is synthesized by a random ring-opening copolymerization of the cyclic dimers of glycolic acid and lactic acid, whereby successive monomeric units of PGA or PLA are linked together by ester linkages. One of the primary advantages of PLGA over other biodegradable synthetic polymers is that the ratio of PLA to PGA used for the polymerization can be adjusted to alter the biodegradation rate of the product (Fig. 9.2). The rate of drug release from PLGA-based drug delivery implants depends on several factors, including the total surface area of the device, the percentage of loaded drug, the water solubility of the drug, and the speed of polymer degradation (Shive and Anderson 1997). The three main factors that determine the degradation rate of PLGA copolymers are the lactide:glycolide ratio (Fig. 9.2), the lactide stereoisomeric composition (i.e., the amount of L-lactic acid vs. D,L-lactic acid), and the molecular weight (Chu 2008; Avgoustakis 2008).

The lactide:glycolide ratio and stereoisomeric composition are most important for PLGA degradation as they determine polymer hydrophilicity and crystallinity. These properties can be manipulated during manufacturing to produce PLGA copolymers with degradation times ranging from weeks to years (Avgoustakis 2008). Lactic acid is more hydrophobic than glycolic acid; therefore, lactide-rich PLGA copolymers are less hydrophilic, absorb less water, and subsequently degrade more slowly (Jain 2000). The most widely used PLGA composition of 50:50 has the fastest biodegradation rate (50–60 days) of the D,L-lactide/glycolide polymers and degrades faster than either PLA or PGA (Yasukawa et al. 2006). PLGA copolymers with lactide:glycolide ratios of 65:35, 75:25, and 80:20 have progressively longer in vivo lifetimes (Mundargi et al. 2008). Lactide-rich PLGA copolymers (up to 70%) are typically amorphous and degrade more rapidly. As the molecular weight of the polymer decreases, the degradation rate increases due to the higher content of carboxylic groups at the end of the

Fig. 9.2  Relationship between lactide/glycolide content and degradation half-life of PLGA. Graph illustrates how PLA and PGA polymer ratios can be adjusted to result in a specific biodegradation half-life and drug release from the implant

polymer chain, which accelerate hydrolysis (Park et al. 2005). PLGAs prepared from PLLA and PGA are crystalline copolymers, and those from PDLLA and PGA are amorphous (Jain et al. 1998). PLGA is degraded in vivo, through hydrolysis and enzymatic action, to lactic acid and glycolic acid which are ultimately converted via the citric acid cycle to water and carbon dioxide (Fig. 9.3).

The versatile properties of PLGA have led to its use in the production of a variety of biomedical devices. PLGA can be used to construct biomaterials of various types and shapes such as resorbable suture materials; rods, screws, plates, and pins for orthopedic surgery; vascular grafts and stents; and surgical meshes and scaffolding for tissue regeneration (see Sect. 9.3). PLGA has also been widely utilized in the development of various types of drug-release systems (e.g., drug implants and micro/nano-particulates) for ocular diseases of the posterior and anterior segments.

PLA, PGA, and PLGA are cleaved predominantly by nonenzymatic hydrolysis of their ester linkages throughout the matrix in the presence of water in the surrounding tissues. This process, referred to as bulk erosion (Fig. 9.4a–f), is distinguished from surface erosion (Fig. 9.4g–l) of the drug/polymer matrix surface (Yasukawa et al. 2006) occurring with polymers such as PAH and POE. Drug release from PLAand PLGA-based matrix drug delivery systems generally follows pseudo first-order or square-root kinetics, and the release rate is influenced by many factors including polymer type, drug load, implant morphology, and porosity. In general, drug release from PLGA-based implants, which is depicted in Figs. 9.4a–f and 9.5, occurs in three phases:

1. Burst release: Drug release from the implant surface occurs, creating a short period of high drug release.

Fig. 9.5  Random scission of polymer chains gradually occurs throughout the biodegradation process, and chain scission occurs more readily in center of the implant. Once chains reach a threshold of reduced molecular weight, the shortened polymer chains become solubilized in the surrounding medium and are liberated from the implant complex. This results in central hollowing of the implant with gradual mass loss of the implant

9.3.2  Poly(e-Caprolactone)

Poly(e-caprolactone) (PCL) is a biodegradable, semicrystalline hydrophobic polymer with a low glass transition temperature (~60°C) and a melting point ranging between 59 and 64°C (Silva-Cunha et al. 2009). Due to its crystallinity and hydrophobicity, biodegradation of PCL occurs very slowly (from months to 1 year), which makes it suitable for a range of biomedical applications such as sutures, orthopedic fixation devices, and a variety of extended-release drug delivery systems (Park et al. 2005; Silva-Cunha et al. 2009). PCL-based drug delivery systems degrade by a twophase process involving an initial phase of bulk hydrolysis followed by a second phase characterized by loss of mass due to chain cleavage and drug diffusion from the polymer matrix (Silva-Cunha et al. 2009). PCL-based implants have been investigated for the ocular delivery of triamcinolone acetonide (Beeley et al. 2005) and dexamethasone (Silva-Cunha et al. 2009; Fialho et al. 2008) and have shown encouraging efficacy and tolerability results.

PCL also can be blended with other polymers to form copolymers with a range of degradation characteristics. Examples of PCL copolymers investigated for biomedical applications include poly(-e-caprolactone)-poly(ethylene glycol) (PCL-PEG), poly(glycolide-co-lactide-co-caprolactone) (PGLC), and poly (glycolide-e-caprolactone-trimethylene carbonate). PCL-PEG copolymers are more hydrophilic than unmodified PCL and therefore have faster degradation times (da Silva et al. 2009). PCL-PEG is being investigated for use in micro/ nanoparticle and hydrogel drug delivery systems (Wei et al. 2009). PGLC, a copolymer of PCL and PLGA, has been studied as a drug delivery vehicle for cyclosporine (Dong et al. 2006a) and tacrolimus (FK-506) (Shi et al. 2005) in animal models of corneal allograft and uveitis (see Sect. 9.6).