Chapter 12

Hydrogels for Ocular Posterior Segment Drug Delivery

Gauri P. Misra, Thomas W. Gardner, and Tao L. Lowe

Abstract  This chapter discusses emerging hydrogel technology for drug delivery to the back of the eye to treat retinal diseases. The review includes design, characterization and optimization of hydrogels, and advantages and disadvantages of intravitreally and subconjunctivally administrated hydrogels for retinal therapy. Future direction of hydrogel technology for targeted and sustained delivery of drugs to the retina for individualized medicine is also laid out.

12.1  Introduction

The ocular posterior segment, which includes mainly retina, choroid, and optic nerves, plays a vital role in maintaining good vision. Any damage to the back of the eye, especially to the retina, due to disease or disorder leads to vision loss. The major retinal diseases that require better treatment approaches include macular edema, age-related macular degeneration (AMD), diabetic retinopathy, retinitis pigmentosa, cytomegalovirus retinitis, uveitis, retinal detachment, ocular melanoma, and retinoblastoma. Many therapeutic agents including antivascular endothelial growth factor, anti-inflammatory, and neuroprotective drugs/agents have attracted growing interest for the treatments of these retinal diseases (Janoria et al. 2007; Gilhotra and Mishra 2008; Hironaka et al. 2009; Lee and Robinson 2009). These therapeutic agents can be administrated to the retina via topical, intravitreal, subconjunctival, and systemic routes.

T.L. Lowe (*) 

Department of Pharmaceutical Sciences, School of Pharmacy, Thomas Jefferson University, Philadelphia,

PA 19107, USA

Department of Pharmaceutical Sciences, College of Pharmacy, University of Tennessee Health Science Center, Memphis, TN 38163, USA

e-mail: tao.lowe@jefferson.edu

U.B. Kompella and H.F. Edelhauser (eds.), Drug Product Development for the Back of the Eye, 291 AAPS Advances in the Pharmaceutical Sciences Series 2, DOI 10.1007/978-1-4419-9920-7_12,

© American Association of Pharmaceutical Scientists, 2011

However, prolonged efficiency and toxicity of these agents are still major problems to be resolved associated with these routes of administration (Hughes et al. 2005). The reasons are that these therapeutic agents usually have very short half-lives, do not cross the blood–retinal barrier (BRB) and retinal pigmented epithelium (RPE) barrier, and are metabolized at other tissue sites. During the last two decades, various methodologies including intraocular implants and iontophoresis/osmotic pumps have been developed to overcome the BRB and the RPE barrier, and reduce/diminish the toxicology and improve the pharmacoefficacy of the therapeutic agents in the retina (Peyman and Ganiban 1995; Hughes et al. 2005; Myles et al. 2005; Bourges et al. 2006; Gaudana et al. 2010). This chapter will review the current technologies especially with a focus on hydrogel technology for drug delivery to the back of the eye for treating the major retinal diseases. The challenges and future directions for the hydrogel technology are also discussed.

12.2  Hydrogel Technology

Hydrogels are three-dimensional network of polymer chains containing water within the network. The network can be either physically or chemically cross-linked (a covalent bonding) structures between polymer chains. Water content of hydrogels can be adjusted by modulating the composition and conformation of polymers, such as hydrophilic/hydrophobic balance of polymer chains and pendant groups, and degree of cross-linking. Hydrogels can swell and shrink in response to external stimuli, such as changes in temperature, pH, solvent, electric current, magnetic field, and ultrasound, and are called stimuli-responsive hydrogels. Hydrogels can also respond to changes of some biomolecules, such as glucose and single-stranded DNA (Murakami and Maeda 2005). Due to their attractive swelling and responsive properties and resemblance to biological tissues (Serra et al. 2006), hydrogels have been extensively studied for potential drug delivery, tissue engineering, and medical device applications (Hoffman 2002).

Hydrogels can control release of all type of therapeutic agents, hydrophobic and hydrophilic, and small and big molecules. Moreover, the aqueous environment of hydrogels has advantages in protecting peptide, protein, oligonucleotide, and DNA types of therapeutics from degradation and denaturation. Kinetics of drug release from hydrogels can be controlled via three important mechanisms: diffusion of drugs, swelling of hydrogels, and degradation of hydrogels. The chemical composition (hydrophilic/hydrophobic balance), ratios of monomers/macromers and solvents, and cross-linking density (porosity) of hydrogels play important roles in controlling swelling of hydrogels, and subsequent diffusion and release of drugs from hydrogels. Drug release kinetics can also be tuned by controlling the degradation rate of the hydrolytically (carbonate, ester, polyphosphate, and phosphazene) or enzymatically degradable bonds of the hydrogels, and changing external stimuli of stimuli-respon- sive hydrogels. Finally, drug release kinetics from hydrogels is also affected by the nature of drugs. While hydrogels have been used for controlled release of drugs for

treating diseases in heart, lung, kidney, liver, spleen, brain, bone and cartilage, development of hydrogels for treating diseases in the eye, especially posterior segment of the eye, is an emerging field with excitements and challenges. The recent developments of hydrogels for delivering drugs to the back of the eye are reviewed below.

12.3  Design, Characterization, and Optimization

of Hydrogels

Hydrogels for ophthalmic use can be designed by taking into account the chemical and physical properties of the components of hydrogels. Ocular biocompatibility, dosage, and duration of release kinetic are controlled by hydrophilic/hydrophobic balance of polymer chains and pendant groups, degree of cross-linking, and monomer and solvent ratio. Both natural and synthetic polymers can be used to fabricate hydrogels. Degradable hydrogels are made using carbonate, ester, polyphosphate, phosphazene, and enzymatically cleavable linkers in the polymer chains or in the cross-linker. Degradation rate can be controlled by cross-link density and hydrophilic/hydrophobic balance.

Alginate, starch, collagen, chitosan, gelatin, and hyaluronate are most widely used natural polymers for hydrogel fabrication because of their intrinsic biocompatibility (Liu et al. 2006; Gilhotra and Mishra 2008; Al-Kassas and El-Khatib 2009; Wadhwa et al. 2009; Gorle and Gattani 2010; Lai et al. 2010; Tanaka et al. 2010). Enzymes, such as alginase, amylase, collagenase, dextranase, and protease can degrade alginate, starch, collagen, dextran, and amide-bond containing hydrogels (Moriyama and Yui 1996; Dai et al. 2005). The most commonly used synthetic polymerbased hydrogels are based on poly(2-hydroxyethyl methacrylate), poly(ethylene glycol), poly(acrylamide), poly(N-isopropylacrylamide), and poly(acrylic acid) (Eljarrat-Binstock et al. 2008a, b; Kang and Mieler 2008; Swindle-Reilly et al. 2009; Singh et al. 2010). These synthetic hydrogels can be chemically cross-linked hydrogels obtained by radical polymerization of monomers in the presence of crosslinkers, reaction of functional side groups of polymers with themselves or other functional cross-linkers. One example is thermoresponsive hydrogels based on poly(isopropylacrylamide), which become attractive biomaterials for ocular drug delivery in recent years (Kang Derwent and Mieler 2008; Misra et al. 2009). These hydrogels have been made by polymerizing N-isopropylacrylamide with macromer cross-linkers including dextran-lactate 2-hydroxyetheyl methacrylate (dextranlactate HEMA) (Misra et al. 2009) and poly(ethylene glycol) diacrylate (PEGDA) (Kang Derwent and Mieler 2008). These macromers can control the pore size, and/ or the hydrophobicity/hydrophilicity, and/or the degradation of the hydrogel networks, and thus subsequent drug release kinetics. Polymers bearing carboxylic, or hydroxyl, or amino side groups, such as poly(acrylic acid), poly(ethylene glycol), chitosan, can be incorporated to enhance the mucoadhesive property of the hydrogels (Park and Robinson 1987; Huang et al. 2000; Peppas et al. 2000; 2009; Ludwig 2005). Hydrogels can also be made by directly cross-linking polymer chains

without using monomers and macromers containing double bonds. For example, 1-ethyl-3-(3-dimethyl aminopropyl) carbodiimide can cross-link the carboxylic groups of hyaluronic acid to make hyaluronic acid-based hydrogels (Lai et al. 2010). Besides the above chemical composition and synthesis method, the physical geometry of the hydrogels can also strongly affect drug release kinetics, and thus is an important parameter that needs to be considered (Lin and Metters 2006; Misra et al. 2009). The physical geometry of the chemically cross-linked hydrogels also strongly depends on the sites (tissue biology and physiology) of the administration if in situ gelation is performed. In situ chemical gelation can be initiated by ultraviolet and visible light polymerization, and oxidation of thiol groups of polymer chains and then formation of disulfide bonds (Swindle-Reilly et al. 2009). After hydrogels are synthesized and optimized, purification with repeated washing is next important step to remove traces of un-reacted components of chemically cross-linked hydrogels to avoid ocular toxicity.

Physically cross-linked hydrogels are obtained through hydrogen bonding, hydrophobic interaction, electrostatic interaction, or stereo-complex formation. Thermogelation is a physically cross-linking process when liquid polymer solutions turn into hydrogels due to hydrophobic interactions among polymer chains as temperature is elevated (Kumar et al. 1994; Nanjawade et al. 2007; Mundada and Avari 2009; Gao et al. 2010; Shastri et al. 2010; Yin et al. 2010). Commonly used thermogelling polymers are triblock poly-(DL-lactic acid-co-glycolic acid)-polyethylene glycol-poly-(DL-lactic acid-co-glycolic acid) (Gao et al. 2010), Carbopol® (polyacrylic acid cross-linked with allyl penta erythritol) (Kumar et al. 1994), and PoloxamerorPluronics(poly(propyleneoxide)-poly(ethyleneoxide)-poly(propylene oxide)) (Shastri et al. 2010). Thermogelling polymers can be made degradable by introducing a hydrophobic degradable polymer in the main polymer chain. For example, poly(ethylene glycol)-poly(e-caprolactone)-poly-(ethylene glycol) can hydrolytically degrade due to the presence of ester bond in poly(e-caprolactone) (Yin et al. 2010).

After hydrogels are formed, confirmation of the chemical structures, characterization of the physical properties, and evaluation of the biological properties of the hydrogels are the next important and necessary steps before the hydrogels can be used for controlled ocular drug delivery clinically. Attenuated total reflection Fourier transform infrared spectroscope (ATR-FTIR), nuclear magnetic resonance (NMR) spectroscopy, and mass spectroscopy are common instruments to characterize the chemical structures and degradation properties of hydrogels (Huang et al. 2004; Huang and Lowe 2005). X-ray photoelectron spectroscopy is helpful in getting information about the surface chemical composition of the hydrogels (Van Tomme et al. 2008; Sánchez-Vaquero et al. 2010). The swelling properties of hydrogels can be studied by calculating the swelling ratio q:

where Wt and W0 are the weights of the swollen and dry gels, respectively (Huang et al. 2004; Huang and Lowe 2005). The thermoresponsive properties of hydrogels can be characterized by measuring swelling ratios as a function of temperature

(Huang et al. 2004; Huang and Lowe 2005). The degradation kinetics of hydrogels can be determined by measuring the weight loss of the hydrogels as a function of time (Huang et al. 2004; Huang and Lowe 2005). X-ray diffraction (Yokoyama et al. 1986; Pal et al. 2008; Szepes et al. 2008) and differential scanning calorimetry (DSC) (Peppas and Mongia 1997; Alvarez-Lorenzo et al. 2002; Andrade-Vivero et al. 2007) provide crystallinity and melting and glass transition temperatures of hydrogels. Porosity of hydrogels can be measured using mercury porosimetry (Liu et al. 2000; Maia et al. 2005), scanning electron microscopy (Yokoyama et al. 1986; Luo et al. 2000; Zhang et al. 2001; Alvarez-Lorenzo et al. 2002; Zhou et al. 2008), and magnetic resonance microscopy (Mishra et al. 2007). Defect in hydrogels can be detected using scanning laser acoustic microscopy (Luprano et al. 1997). The mechanical properties of hydrogels including storage, loss and compression moduli, tensile strength, and viscoelastic behaviors can be measured by rheometer and dynamic mechanical analyzer (Liu et al. 2006; Schuetz et al. 2008; Zhou et al. 2008; Swindle-Reilly et al. 2009). The common cell viability assays including trypan blue (Chirila et al. 1992) or propidium iodide (Debbasch et al. 2002) exclusion, lactate dehydrogenase (LDH) (Khattak et al. 2006), 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) (Cao et al. 2007; Misra et al. 2009) and 3-(4,5-dimethylthiazol-2-yl)-5-(3- carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS) (Kang Derwent and Mieler 2008) can be used to evaluate the cytotoxicity of hydrogels. The above chemical, physical, and biological characterizations of hydrogels provide important information for further optimization of the hydrogels for desired ocular drug delivery need. Size and hydrophilic/hydrophobic properties of drug molecule should also be taken into account for optimization of hydrogels.

12.4  Intravitreally Administered Hydrogels

for Retinal Therapy

Delivering drugs to the retina via intravitreal route has the least transport barriers. Retinal diseases, such as choroidal neovascularization, diabetic macular edema, ischemic neovascularization, inflammatory and infectious processes, are frequently treated with intravitreal pharmacotherapies. Reduced foveal thickness in diabetic macular edema and improved visual acuity have been observed with intravitreal injections of bevacizumab and triamcinolone acetonide (Shimura et al. 2008). Considerable advances in the delivery of drugs for vitreoretinal disease have been made during recent years (Yasukawa et al. 2004). Both nondegradable and degradable polymeric implants have been used for delivering drugs to the eye. In fact, several commercial intravitreal implants are already in clinical use. For example, sustained release implants, such as Vitrasert™ and Retisert™ (Bourges et al. 2006; Jaffe et al. 2006; Choonara et al. 2010; Kuno and Fujii 2010), are based on nondegradable poly(ethylene-co-vinyl acetate) polymer membrane, which allows transport of mainly hydrophobic/lipophilic drugs. Vitrasert releases ganciclovir to treat cytomegalovirus retinitis for approximately 4–5 months (Sanborn et al. 1992); while a

constant release of fluocinolone actinide, a drug for treating noninfectious uveitis, for 2.5 years can be achieved from Retisert (Jaffe et al. 2006). OZURDEX™ (Allergan 2009), a rod-shaped PLGA intravitreal implant containing dexamethasone, has been approved as first-line therapy for macular edema due to retinal vein occlusion. It can correct visual acuity for 1–3 months.

Intravitreally injectable hydrogels have been recently explored that can be preformulated with several microliter volumes that can be injected via a fine needle or polymer solutions that are liquid at ambient temperature and transform into hydrogels once injected into the vitreous cavity due to their thermogelation property. Intravitreally injectable (3–5 mL) hydrogels composed of poly(N-isopropylacrylam- ide) (PNIPAAm), cross-linked with poly(ethylene glycol) diacrylate have been studied (Kang and Mieler 2008). PNIPAAM is hydrophobic at temperature above 32°C while PEG introduces hydrophilic property into the hydrogels. The hydrogels could encapsulate and release proteins, such as immunoglobulin G (IgG), bevacizumab and ranibiumab in vitro, and be injected into the vitreous via a 30-gauge needle. A dispersion of ganciclovir-loaded PLGA microspheres into poly(D,L- lactide-co-glycolide)-PEG-poly(D,L-lactide-co-glycolide) (PLGA–PEG–PLGA) solution has been used as an injectable hydrogel forming solution (Duvvuri et al. 2007). Upon injection into the vitreal cavity, the solution turned into hydrogels due to the thermogelation of the PLGA–PEG–PLGA polymer. The microsphere-gel delivery system could maintain vitreal concentration of ganciclovir at ~0.8 mg/mL for 14 days, whereas direct injection of ganciclovir could only keep the drug level for slightly over 2 days. The role of PLGA–PEG–PLGA gel was apparently to hold PLGA microspheres within the gel matrix, while ganciclovir release was controlled mainly by the PLGA microspheres.

Although intravitreally implantable/injectable hydrogels can efficiently deliver drugs to the retina and choroid, the implantation/injection process is invasive and can cause complications. Cataract formation, glaucoma, vitreous hemorrhage, choroidal detachment, retinal detachment, hypotony, vitreous loss, exacerbation of intraocular inflammation and wound dehiscence, and endophthalmitis are also associated with the implants (Prasad et al. 2007; Choonara et al. 2010). Moreover, the implants/injections need to be replaced requiring multiple clinic visits and surgical procedures.

12.5  Subconjunctivally Administered Hydrogels

for Retinal Therapy

Because of invasive nature and associated complications of intravitreal treatments, interest in subconjunctival retinal drug delivery is growing (Kompella et al. 2003; Gukasyan et al. 2007; Mac Gabhann et al. 2007). This route is well-suited for retinal delivery since the internal limiting membrane that separates the retina from the vitreous limits diffusion to the vitreous (Mac Gabhann et al. 2007). The subconjunctival space offers the possibility of delivering multiple types of sustained-release formulations

to the posterior eye. Almost 80% of the ocular surface is covered by conjunctival tissue in higher mammalian species. Absorption and transport of drugs via paracellular, transcellular, active, and endocytic routes can play a key role in retinal delivery. The subconjunctival route is safe but less efficient than the intravitreal route. Even though the presence of BRB and RPE barrier can limit drug transport to the retina via the subconjunctival route (Cheruvu et al. 2008), it has been reported in the literature that subconjunctivally administered molecules up to 70 kDa in size can penetrate the sclera to reach the retina and retain biologic activity (Ambati et al. 2000; Ambati and Adamis 2002; Kim et al. 2002; Mac Gabhann et al. 2007), and subconjunctival administration achieved drug level in retina several fold higher than intraperitoneal route did (Ayalasomayajula and Kompella 2004a, b). To achieve long-term release of drugs to the back of the eye via the subconjunctival route, polymer implants, microparticles, and nanoparticles have been developed. Polycaprolactone, a nonwater soluble, hydrophobic and hydrolytically degradable polymer was used to fabricate transscleral implants with 7 mm in diameter and 2.5–3 mm in thickness (Carcaboso et al. 2010). The implants showed sustained release of 102–103 ng/g topotecan, a drug for treating retinoblastoma over 48 h. The Kompella group demonstrated that subconjunctivally injected fluorescent polystyrene nanoparticles (200 nm) and microparticles (2 mm) were retained in the periocular tissue for at least 60 days, suggesting that microand nanoparticles could be used for sustained drug delivery to the back of the eye after subconjunctival administration (Amrite and Kompella 2005). Komplella group further demonstrated that budesonide (Kompella et al. 2003) and celecoxib (Ayalasomayajula and Kompella 2004a, b, 2005), anti-VEGF agents, released from subconjunctivally injected PLAbased nanoand microparticles and celecoxib-loaded PLGA microparticles could maintain therapeutic level in the retina for 2 weeks’ period of time.

Hydrogels have unique cross-linked networks that resemble the nature matrix surrounding all cells within human tissues and control release drugs by changing their swelling and shrinking properties, compared to polymer implant, microparticles and nanoparticles, and thus hydrogels have drawn increasing attentions as biocompatible and subconjunctivally implantable delivery systems for posterior segment of eye diseases (Eljarrat-Binstock et al. 2010). Recently, subconjunctivally implantable, biodegradable hydrogels have been developed for sustained local release of intact insulin to the retina to provide neuroprotective effects without ­risking hypoglycemia (Misra et al. 2009). The hydrogels were synthesized by UV ­photopolymerization of N-isopropylacrylamide (NIPAAm) monomer and a ­dextran macromer containing multiple hydrolytically degradable oligolactate- (2-hydroxyetheyl methacrylate) units (Dex-lactateHEMA) in 25:75 (v:v) ethanol: water mixture solvent. Insulin was loaded into the hydrogels during the synthesis process with loading efficiency up to 98%. Small hydrogels (2 mm diameter, 1.6 mm thickness, 5 mL volume) were fabricated for the purpose of implanting the hydrogels in the subconjunctival space of rat eyes to continuously release insulin to the retina. The hydrogels released biologically active insulin in vitro for at least 1 week and the release kinetics could be modulated by varying the ratio between NIPAAm and Dex-lactateHEMA and altering the physical size of the hydrogels.

Fig. 12.1 Electroretinograms (ERGs) after subconjunctival implantation of insulin-loaded poly (NIPAAm-co-Dex-lactateHEMA) hydrogels in Sprague–Dawley rats. (a) Dark adapted ERGs elicited by increasing intensities of light in implanted eyes, (b) B-wave luminance-response functions from all eyes with insulin-loaded hydrogels (open symbols), (c) ERGs in response to medium-energy light stimuli in four animals implanted with insulin-loaded hydrogels, (d) B-wave amplitude in implanted eyes (insulin-loaded or blank) plotted against contralateral controls (blank or untouched) (Misra et al. 2009)

The poly(NIPAAm-co-Dex-lactateHEMA) hydrogels and their 7-day degradable components were not toxic to R28 retinal cells in vitro. Moreover, subconjunctival implantation of the hydrogels to Sprague–Dawley rats did not cause any morphological change, inflammation, or other adverse effects in the rat eyes over at least a 1-week period of time, as evidenced by the results of H&E staining, immunostaining for microglial cell activation, and electroretinograms (ERGs, Fig. 12.1). After subconjunctival implantation, the ERG data did not show any adverse effect on the retinal function. It was concluded that the developed nontoxic thermoresponsive and degradable hydrogels have high potential to control the release of insulin and other therapeutics to the retina after subconjunctival implantation, which may lead to a new option for treating diabetic retinopathy and other retinal disorders.

To minimize the invasiveness of the subconjunctival implantation process, hydrogels have been formed in situ in the subconjunctival space. In situ matrices containing

fluocinolone acetonide (FA)-loaded biodegradable polymer poly(propylene fumarate) (PPF) have been fabricated by injecting N-methyl-2-pyrrolidone (NMP) solubilized (FA) and PPF in aqueous solution (Ueda et al. 2007). Matrix formation occurs due to the diffusion of NMP into surrounding aqueous fluids. Cross-linked surfaces on the matrix are also formed by using UV-photopolymerization. The PPF matrices released FA for almost 1 year in vitro (Ueda et al. 2007) and achieved more constant release of FA with cross-linked surfaces. This delivery system has potential for prolonged therapy for retinal diseases through the subconjunctival or intravitreal route although more observations will be needed to confirm the in vivo release profile and ocular distribution of drugs. A commercially available ReGel™ solution containing biodegradable and thermosensitive triblock copolymer consisting of poly(lactic-co-glycolic acid) (PLGA) and polyethylene glycol (PEG) was recently injected into the subconjunctival space and formed hydrogels in situ due to the thermogelation property of the copolymer (Rieke et al. 2010). Ovalbumin was mixed with the ReGel™ solution prior to the injection and was held by the ReGel™ hydrogels during the in situ gelation process. The formed ReGel™ hydrogels could continuously release measurable ovalbumin to the retina of rats for 14 days.

Another utility of hydrogels for transscleral drug delivery is to hold drugs in transscleral iontophoresis devices. The iontophoresis devices involve application of a weak electric current to enhance transport of charged drugs across percutaneous tissue. Commercially available iontophoresis probes, Eyegate™ or OcuPhor™, have been used for this purpose (Myles et al. 2005). Iontophoretic ocular devices use hydrogels as reservoir for holding drugs to minimize tissue irritation and current interruptions and control drug release kinetics by modulating the chemical and physical properties of the hydrogels (Eljarrat-Binstock et al. 2005, 2007, 2008a, b). However, iontophoresis requires direct current that may cause a burning sensation. This limits the use of the iontophoretic technique for sustained drug delivery. Nevertheless, short-term delivery of anti-inflammatory drugs such as methylprednisolone, dexamethasone, and methotrexate to the retina of rabbits has been successfully achieved by Eljarrat-Binstock et al. using drug-loaded hydrogels made of hydroxyethyl methacrylate cross-linked with ethylenglycol dimethacrylate that were mounted on a portable iontophoretic device (Eljarrat-Binstock et al. 2005, 2007, 2008a, b). However, the usefulness of the hydrogel-iontophoresis systems for delivering drugs to the retina transsclerally is dependent on the property of the drugs. For example, Eljarrat-Binstock et al. found that the presence of the same hydrogels in the iontophoresis systems for delivering methylprednisolone, dexamethasone, and methotrexate did not improve delivery of carboplatin to treat retinoblastoma due to the high diffusion property of carboplatin (Eljarrat-Binstock et al. 2008a, b).

The newest exciting area in the application of hydrogels for treating retinal diseases is for delivering stem cells. In a recent article, Ballios et al. injected a blend solution of hyaluronan and methylcellulose with retinal stem-progenitor cells (RSPCs) in the subretinal space of adult CD10/Gnat2−/− mice, a mouse model of AMD (Ballios et al. 2010). The polymer blend solution thermally gelated into hydrogels at the physiological temperature. Four weeks postinjection, distribution