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Polymer 52 (2011) 5262e5270

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Polymer

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Evaluation of redox-responsive disulfide cross-linked poly(hydroxyethylmethacrylate) hydrogels

Muhammad Ejaza, Haini Yub, Yong Yana, Diane A. Blakeb, Ramesh S. Ayyalac, Scott M. Graysona,*

aDepartment of Chemistry, Tulane University, New Orleans, LA 70118, United StatesbDepartment of Biochemistry, Tulane University, School of Medicine, New Orleans, LA 70112, United StatescDepartment of Ophthalmology, Tulane University, School of Medicine, New Orleans, LA 70112, United States

a r t i c l e i n f o

Article history:Received 3 June 2011Received in revised form7 September 2011Accepted 9 September 2011Available online 21 September 2011

Keywords:Poly(hydroxyethyl methacrylate) hydrogelStimulus responsiveDisulfide cross-linker

* Corresponding author. Tel.: þ1 504 862 8135; faxE-mail address: sgrayson@tulane.edu (S.M. Grayso

0032-3861/$ e see front matter � 2011 Elsevier Ltd.doi:10.1016/j.polymer.2011.09.018

a b s t r a c t

Poly(2-hydroxyethyl methacrylate) (PHEMA) hydrogels were prepared with a disulfide containing cross-linker bis(2-methacryloxyethyl) disulfide (DSDMA) that exhibited enhanced release in the presence ofglutathione (GSH), a biologically available reducing agent. Varying concentrations of the DSDMA cross-linker were incorporated into the prepolymer before the radical polymerization, enabling the cross-link density to be easily tuned. Dye release studies were performed using rhodamine B and rhoda-mine 6G dyes, and the UV response of the dyes released into the supernatant measured with the additionof GSH. Using ether-based cross-linkers as a control, the disulfide cross-linkers exhibited a substantialincrease in release rates, confirming the responsive nature of the hydrogels to biological reducing agents.The polymers were also tested in a cell culture system for their ability to release the anti-fibroproliferative agent, mitomycin C (MMC). Polymers cross-linked with DSDMA delivered MMC overa slightly longer time period than control polymers prepared with a conventional ether cross-linker.

� 2011 Elsevier Ltd. All rights reserved.

1. Introduction

In the recent years, a number of novel polymer-based deliverysystems have been developed for the entrapment of proteins, drugs,genes or antigens in devices such as micelles, vesicles, or withinhydrogel matrices [1]. Polymer based drug systems have attractedmuch attention in the fields of polymer chemistry, pharmaceuticsand biomedical science because of the potential for both targeteddeliveryand controlled release of the drug payload [2].While solublepolymer-drug conjugates can exhibit mobility in vivo to achieveselective targeting, insoluble hydrogels can offer the complimentaryproperty of immobility once they are implanted or injected into thedesired site of treatment. Hydrogels are three-dimensional cross-linked hydrophilic polymer networks stabilized through physical orchemical cross-links, and can be designed to absorb up to manythousands of times theirdryweight inwater, enabling themtomimicthe physical characteristics of soft tissues. Their chemical composi-tion and three dimensional structure can be easily modified to tunetheir swelling, mechanical properties, biocompatibility, and encap-sulation of drugs [3,4]. The insoluble and often robust character ofthe hydrogels provides them with mechanical stability, yet the

: þ1 504 865 5596.n).

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permeability of hydrogel systems to small molecules and solventsalso maintains a dynamic behavior typical of liquid phases [5].Because they can be designed to exhibit exceptional biocompati-bility, (e.g. reduced inflammatory responses, thrombosis and tissuedamage,) hydrogels have been studied extensively for biomedicalapplications [2,6] including drug delivery [7e13] and tissue engi-neering [14,15]. Hydrogels also offer significant opportunities forinvivo targetedapplicationsdue to theability to control their size andshape, the ability to tune their porosity and their ability to incorpo-rate biorelatedmolecules suchasDNA, proteins anddrugs.Hydrogelsare also easy to functionalize formultivalent bioconjugation [16e18].

Recently stimuli-responsive smart hydrogels have been devel-oped with a variety of triggers including enzymes [19], pH [20],temperature [21] light [22], electric field [23], magnetic field [24]and selective chemical triggers (for example metal ions orglucose) [25] with versatile applications that include controlleddrug and gene delivery systems [6,26] chemical and biologicalseparations [27,28] as well as sensors and actuators [29,30]. Pol-y(hydroxyethyl methacrylate) (PHEMA) is one of the most widelyused hydrogels for various biomedical applications such as drugdelivery systems [31,32], tissue engineering, and contact lenses[33,34] because of its biocompatibility [32]. PHEMA hydrogelsconsist of a framework of PHEMA chains with intermediate water-filled voids that act as pores for the passage of the solute [35]. Thehigh density of hydroxyl groups within PHEMA hydrogels results in

M. Ejaz et al. / Polymer 52 (2011) 5262e5270 5263

a hydrophilic surface that exhibits low interfacial free energy withmost body fluids, and minimal adhesion of proteins and cells totheir surfaces. Because of this attractive set of properties, manyresearch groups have produced PHEMA-based polymeric matriceswith a range of biological function including: adsorption of heparin,proteins and DNA [34], novel scaffolds for bone tissue engineering[36], optical biosensors for the detection of phenol compounds [37]and selective and reversible albumin adsorption [38].

Despite the substantial advantages of PHEMA gels in biologicalsystems, their degradation is relatively slow under physiologicalconditions. However, for several of the most attractive applicationsof these hydrogels, the fast response that might be achieved ina degradable system is desired. Perhaps one of the most chal-lenging, but important requirements of an effective polymer-baseddelivery system is the controlled but relatively rapid release of theirpayload e.g. at very specific times and locations [39]. In the case ofhydrogels, their payload can be released by diffusion, by an inducedchange in the properties (e.g., permeability) of the hydrogel, or bytriggered degradation of the hydrogel into soluble materials. Inaddition to enabling a rapid release of therapeutic cargo, degrad-able hydrogels often yield smaller polymer components that aresoluble and can be removed from the body through renal filtrationor biliary excretion [40].

Degradable “hybrid” PHEMA hydrogels can be prepared by theincorporation of degradable cross-linkers. PHEMA-based degradablehydrogels are advantageous not only because the hydrogel startingmaterials are compatible in vivo, but also because systems can bedeveloped such that the bi-products of the degradation are bothwater soluble and biocompatible. A wide variety of triggers can beutilized to enable the accelerated degradation of, and thereforeaccelerated release from, hydrogels, including enzymes, pH, irradia-tion, etc. An additional stimulus for degradable hydrogels thatwouldbe attractive for some drug delivery applications would involve theredox degradationof cross-linkers. Degradation activated by changesin the redox state isphysiologically relevantbecausea redoxpotentialoccurs across cellular membranes [41]. The cytoplasm is kept ina reducing environment by the presence of reducing agents such asglutathione (GSH), whereas extracellular spaces such as the blood-stream are comparatively oxidizing. Therefore the incorporation ofa redox-active bond, such as a disulfide [42e46], into a deliverysystem can provide stability in oxidizing environments, yet yielda triggered degradation and drug release upon cell entry.

Taking precedence from cysteine’s redox-sensitive role withinprotein structure, hydrogels that contain disulfide cross-links shouldexhibit reversible cross-links which can be degraded by reductivecleavage of thedisulfide cross-linkage (-S-S-) to free thiol groups (-SHHS-). In early studies, Plunkett et al, studied the swelling kinetics ofdisulfide cross-linked HEMA microgels prepared within micro-channels, and “de-cross-linked” by reductionwith dithiothreitol [47]and Armes and coworkers reported biochemically degradable tri-block copolymer gelators by atom transfer radical polymerization(ATRP) [48].More recently Vogt and Sumerlin reported the disulfide-cross-linked hydrogels from ABA triblock copolymers with thiol endgroups prepared by reversible addition fragmentation chain transferpolymerizations [49], Matyjaszewski and coworkers prepareddisulfide-containingnanogels byATRPaspotential drug carriers [50],and Zhang et al. examined the utility of disulfide-cross-linkedhydrogels as a medium for cell proliferation that could be rapidlydissolved by the addition of GSH [51]. Herein, the authors seek toascertain the utility of disulfide cross-linked PHEMA hydrogels fordelivery of anti-fibroproliferative drugs to improve the lifetime ofglaucoma drainage devices. Concurrent with our study, a number ofother groups have also investigated degradable hydrogels based ondisulfide cross-links, confirming their utility for applications such asprotein delivery [52,53], cell encapsulation [52] and dermal wound

healing [54]. Many small molecule biological reducing agents areutilized in vivo, including GSH, and L-cysteine (Cys) [49,55e57]. Theirbioavailability makes them attractive redox triggers to bring aboutthe disintegration of the disulfide cross-linked hydrogels. In the caseof disulfide cross-linked PHEMA-based hydrogels, the resultingsoluble PHEMA fragments should be non-toxic and exhibit solubilityin cell culture medium or in vivo due to the hydrophilicity of PHEMA.Although the disulfide bond should be relatively stable in the cellexterior it should be degraded rapidly in the cytosol because theintracellular concentration of GSH (w10mM) ismore than 500 timeshigher than the extracellular concentration [58]. It has also beenobserved that the concentrationof free thiolgroupsnear tumor tissueismore than 7 times that in normal tissue [59]. Because GSH is one ofthe most ubiquitous reducing agents in vivo [60] the experimentsreported herein will focus on its reductive degradation of disulfidecross-linked hydrogels.

In this study, we report the preparation of new redox-responsivePHEMA hydrogels cross-linked with bis(2-methacryloxyethyl)disulfide (DSDMA) which contains a disulfide linkage between twopolymerizable methacylate functionalities. The hydrogels wereprepared in a one step radical polymerization of a mixture of HEMAand the DSDMA cross-linker. Utilizing glutathione as a reducingagent, the response of the disulfide cross-linked hydrogels toreducing agents was probed. To quantify the enhanced releaseexhibited by the reductive degradation of the disulfide cross-linker,a control was established with equal amounts of a non-degradableether cross-linker di(ethylene glycol) dimethacrylate (DEDMA).The rates of release of a model drug from hydrogels with both thedegradable and the non-degradable cross-linker were quantifiedusing UV fluorescence. The hydrogels were also tested for theirability to release an anti-fibroproliferative drug in a cell culturemodel system.

2. Experimental section

2.1. Materials

2-Hydroxyethylmethacrylate (HEMA, 97%, ALDRICH), di(ethyleneglycol) dimethacrylate (DEDMA, 95%, ALDRICH). N,N,N0,N00-tetra-methyl ethylenediamine (TMEDA, Molecular biology grade,ALDRICH), ammonium persulfate (APS, 98%, Electrophoresis grade,SIGMA-ALDRICH), rhodamine B (RB, 80% dye content, ALDRICH),rhodamine 6G (R6G, ALDRICH), L-glutathione reduced (GSH, SIGMA-ALDRICH), triethyl amine (TEA, 99.5%, ALDRICH), bis(2-hydroxyethyl)disulfide (Technical, SIGMA-ALDRICH), mitomycin C (MMC, SIGMA-ALDRICH), methacryloyl chloride (97%, ALFA AESAR) and all otherreagents were used as obtained from commercial sources withoutadditional purification. Disulfide-based dimethacrylate, bis(2me-thacryloxyethyl) disulfide (DSDMA) was prepared according topublished literature procedure. [61] Tetrahydrofuran (Omnisolv�

THF,without stabilizer, EMDCHEMICALS)was driedwith a systemofdrying columns from the Glass Contour Company. COS-1 cells(AMERICAN TYPE CULTURE COLLECTION) were maintained in Dul-becco’s Modified Eagle’s Medium (INVITROGEN) with 0.4% (w/v)glucose, 4 mM L-glutamine, 1 mM sodium pyruvate, 1.5 g/L sodiumbicarbonate, 10% (v/v) fetal bovine serum (ATLANTA BIOLOGICALS)and antibiotics (100 u/mL penicillin G, 0.25 mg/mL amphotericin Band 100 u/mL streptomycin (SIGMA)). Neutral buffered formalin andtoluidine blue were obtained from SIGMA.

2.2. Instrumentation

UVevis spectral data were recorded using a UV spectropho-tometer (Hewlett-Packard-8452A-Diode array spectrophotometer).The absorbance of solubilized toluidine blue dye from cell culture

M. Ejaz et al. / Polymer 52 (2011) 5262e52705264

experiments was read on a VersaMax Tunable Microplate Readerfrom Molecular Diagnostics.

2.3. Synthesis of hydrogels

Three different types of poly(hydroxyethyl methacrylate)PHEMA-based hydrogels were prepared: hydrogels without anycovalent cross-linker, hydrogels with DEDMA cross-linker andhydrogels with DSDMA cross-linker (Scheme 1). In a typical PHEMAhydrogel synthesis: the desired amount (from 1 to 4% mole withrespect to HEMA monomer) of cross-linking agent (DSDMA orDEDMA) was dissolved in 4 mL of HEMA and mixed with 4 mL ofdouble de-ionized (DI) water and 400 mL of TMEDA. Afterward,a 1 mL aliquot of APS in water (0.5 mol% with respect to HEMAmonomer) was added and solution vigorously mixed. The solutionwas immediately cast between 2 sealed glass slides and allowed tocure for about 15 h at room temperature, following a typicalprocedure for hydrogel preparation using a free radical polymeri-zation pathway initiated by APS and TMEDA [8,9,62]. The resulting1 mm thick hydrogel films were cut into 8 mm diameter round

Scheme 1. Synthetic pathways of the fabrication of cross-linked poly(2-hydroxyethyl met(DSDMA) and (B) nondegradable di(ethylene glycol) dimethacrylate (DEDMA) at room tem

discs using a trephine. Pure PHEMA hydrogels were also preparedusing this procedure (without the addition of any cross-linker,) aswere hydrogels with 1, 2, and 4 mol % of both the DSDMA cross-linker and the DEDMA cross-linker.

The dyes were loaded into hydrogels by two different proce-dures, either before or after the redox polymerization. Tworhodamine dyes, RB and R6G, (Fig. 1) were selected as modelcompounds to quantify hydrogel drug release because of theirsolubility in water and ease of quantification using UVevis spec-trometry. In the pre-polymerization loading procedure, a 1 mg/mLaqueous solution of RB dye was directly mixed with the hydrogelprepolymer in place of pure water. Although this procedure waseffective at incorporating the dye, it prevented post-polymerizationwashing procedures to remove any unreactedmonomer or residualinitiator. After hydrogel formation, the wet dye-loaded hydrogelwas cut into circular discs of 8 mm and used immediately for thecontrol dye release study. In the post-polymerization loadingprocedure, the hydrogels were polymerized without addition ofdye or drug, cut into 8 mm diameter circular discs, and thenallowed to dry completely. Soluble low molecular weight

hacrylate) (PHEMA) hydrogels by a (A) degradable bis(2-methacryloxyethyl) disulfideperature (RT) using redox polymerization method.

OHN NH

O

Cl

O

RB

R6G

GSH

H2N

O

N

O

O

H

H

NH

O

NH2

O

MMC

H

N

N

H

OH

O

O

SH

OO

HO

NH2

Fig. 1. Chemical structure of Rhodamine B (RB), Rhodamine 6G (R6G),Glutathione(GSH) and Mitomycin C (MMC).

Fig. 2. Release profiles of “pre-polymerization-loaded” Rhodamine B from (A) DSDMA-and (B) DEDMA-cross-linked hydrogels at molar percentage of cross-linker of 0 (C),1( ), 2 ( ) and 4 ( ) in water at room temperature.

M. Ejaz et al. / Polymer 52 (2011) 5262e5270 5265

compounds were removed from hydrogels by immersing in 250mLwater for two days with the aqueous medium being replaced every12 h. After the rigorous washing all the discs were immediatelyimmersed in R6G aqueous dye solution (1.5mg/ml or 3.0mM) for atleast 3 days with slow stirring. Excess R6G was removed byimmersing hydrogels in water for two days. The medium wasreplaced initially after every 30 min for 12 h and then every 12 h.The polymer discs were used for the control dye release studiesafter wiping off the excess of dye solution on the surfaces.

2.4. Swelling of hydrogel

The swelling measurement was carried out by immersing thedried hydrogels in DI water at room temperature for two days. Thepercentage of equilibrium water uptake was calculated as follows:

Water uptakeð%Þ ¼ ðWs �Wo=WoÞ � 100

where Ws and Wo are weights of swollen hydrogels and driedhydrogels respectively.

2.5. Control dye release studies

The dye release profiles of PHEMA hydrogels were evaluatedwith respect to the percentage and type of cross-linker, usinga UVevis spectrophotometer. The physically absorbed dye wasreleased from the hydrogels into doubly de-ionized aqueous solu-tions that were slowly stirred in an argon atmosphere to preventany possible oxidative degradation of the dye. Each dye-loadedhydrogel disc was added to 10 mL of water in a 50 mL flask fittedwith three way stopcocks to enable a 10 min cycle of argon purgingto remove oxygen. The flask was then sealed and the aqueoussolutionwas stirred at room temperature. Aliquots (1 ml) of the dyesolution were removed at the desired time intervals and the dyeconcentrations were measured (absorbance at 554 nm for RB and526 nm for R6G) by UVevisible spectrophotometry. The aliquotsremoved for UVevis analysis were returned to the flask after themeasurements to maintain solution volumes and the recombinedsolutionwas purged with argon for 10 min. The dye release amountwas determined from calibration curves based on UVevis absor-bance of RB and R6G control solutions of known concentrations inwater. The amounts of RB and R6G dyes in the aqueous solutionswere determined using the BeereLambert equation with theknown extinction coefficients of RB (554 nm, ε ¼ 86,000 M�1 cm�1

in water) and R6G (526 nm, ε ¼ 82,000 M�1 cm�1 in water),

M. Ejaz et al. / Polymer 52 (2011) 5262e52705266

2.6. Biochemical degradation of disulfide-containing PHEMAhydrogels by reduction with glutathione

The R6G-loaded redox-responsive DSDMA cross-linked 2 mmthick hydrogels were probed to evaluate their redox-acceleratedrelease. Reduced glutathione (GSH) (Fig. 1), a common water-soluble reducing agent within biological systems, was used asa reductive trigger todegrade thedisulfidebondsof theDSDMAcross-linker. The dye release profiles were determined using a protocolidentical to that described for the non-degradable hydrogels, onlywith the addition of GSH at the designated time points. For a typicalprocedure, 3mgof GSHwas added to 10mL aqueousmedium (1mM)containing the R6G-loaded hydrogel. The GSH (as a powder) wasaddedeither simultaneously (t¼ 0h)with the additionof hydrogel, or6 h after the addition of hydrogel to the release media (t ¼ 6 h). Asa control, GSH was also added to the DEDMA-cross-linked hydrogelsand the uncross-linked PHEMA hydrogels. The GHSwas not expectedto enhance dye release in these hydrogel preparations.

2.7. Test of drug release and polymer toxicity in cultured cells

Hydrogel discs were loaded with mitomycin C (MMC) (Fig. 1) byimmersing each 2 mm thick fully swollen hydrogel disc in a 1.5 mLaqueous MMC solution (30 mg/mL) for a desired time interval untilthe required amount of MMC in disc was obtained. The concen-tration of MMC in disc was calculated by difference of MMCconcentration before immersing and after removing the disc fromthe loading solution. The MMC-loaded discs were sterilized by UVradiation for 1 h. All elution/cell culture experiments were per-formed with at least 3 replicate discs. At day 0, each drug-loaded orcontrol disc (with no loaded drug) was immersed in 1 mL completeculture medium and the medium was immediately removed. Anadditional 1 mL of culture medium was added and each disc wasallowed to incubate from0 to 3 days at 37 �C in an atmosphere of 5%CO2/95% air. After the second aliquot of culture medium was

Scheme 2. Schematic representation of (A) reduction of disulfide bonds of bis(2-methacryldi(ethylene glycol) dimethacrylate (DEDMA) cross-linked hydrogels using glutathione (GSH

removed, additional 1 mL aliquots of culture medium were addedand incubation was continued to generate the 3e7 day and 7e14day samples. For some experiments, the culture medium wassupplemented with 2 mM glutathione, the average concentration ofglutathione in human aqueous humor [63].

COS-1 cells were plated in 12-well tissue culture plates (CORN-ING/COSTAR) at 1 � 104 cells/well by adding 1 mL of diluted cellsuspension to eachwell. The plateswere incubated for 4 h at 37 �C inan atmosphere of 5% CO2/95% air, until the cells had attached to theculture well. Then 1 mL of the polymer extract in culture mediumwas added to each well and an additional 1.5 mL of culture mediumwas added so that the final volume of medium in each well was3.5 mL. The cells were cultured for 5 days at 37 �C in an atmosphereof 5% CO2/95% air without a change of culture medium. Preliminaryexperiments demonstrated that uninhibited COS-1 cells grew to fillthe culture well during this 5 day period (data not shown). Thenumber of cells on each well was assessed by a modification of themethod of Leavesley et al. [10]. Briefly, the culture medium wasremoved from each well and the cells were washed twice withphosphate buffered saline and subsequently fixed with 10% neutralbuffered formalin for 1 h. The fixative was removed and 0.5 ml of0.5% toluidine blue in 10% neutral buffered formalinwas used to dyethe cell layer for 1 h. The dyewas removedwith a pipette and the celllayer was rinsed with 4 � 2 mL of water. The plates were air driedovernight. Each culture dish was scanned for a photographic recordof the experiment. The incorporated dye in each well was subse-quently dissolved using 0.5 ml 2% sodium dodecyl sulfate (w/v inwater) for 1h,with rocking.Aliquots (3�100ml) of the dissolveddyewere read with a plate reader at 650 nm.

3. Results and discussion

The PHEMA-based hydrogels used in present study weresynthesized using typical persulfate initiated polymerization che-mistry depicted in the Scheme 1. The free radical polymerizationwas

oxyethyl) disulfide (DSDMA) cross-linked hydrogels and (B) stability of ether bonds of) in water under argon at room temperature.

Table 1Influence of cross-linker concentrations on the hydrogel water uptake.

Comp. PurePHEMA

1%DEDMA

2%DEDMA

4%DEDMA

1%DSDMA

2%DSDMA

4%DSDMA

Wateruptake

85.64 76.63 71.34 69.08 75.62 71.24 67.34

Fig. 4. Release profiles of post-polymerization-loaded Rhodamine 6G from purePHEMA hydrogels with addition of GSH at 0 h ( ), at 6 h ( ) and in absence of GSH inwater at room temperature (C).

M. Ejaz et al. / Polymer 52 (2011) 5262e5270 5267

initiated at room temperature by ammonium persulfate usingN,N,N0,N0-tetramethylethylenediamine to catalyze the decomposi-tion of persulfate ion to a sulfate free radical. With the incorporationof DSDMA or DEDMA, covalent cross-links are incorporated intothe resultant hydrogels, whereas the uncross-linked PHEMA hydro-gels are formed through hydrogen-bonding between the hydroxylside chains on each repeat unit of the polymer. The void spaces inbetween these cross-links can be large enough to enable theencapsulation and controlled release of small molecules. During thepolymerization, increased concentrations of the cross-linker led toa more rapid increase in viscosity and a more rapid gelation, con-firming that the covalent links were critical to increasing themechanical stability of the hydrogels.

A library of hydrogels, containing varying quantities of the twocross-linkers,wasprepared inorder to evaluate the redox-responsiverelease of payload from the loaded hydrogels. Rhodamine dyes, RB(loaded before polymerization) and R6G (loaded after polymeriza-tion), were selected to probe the controlled release because of theirstability, their water solubility, their accurate quantification via UVfluorescence and their similar size to small molecule drugs. Sevenhydrogel formulations were investigated: pure PHEMA hydrogelswithout any cross-linker, PHEMA hydrogels with 1, 2 or 4 mol %DEDMA, or PHEMAhydrogelswith1, 2, or 4mol%DSDMA. Their ratesof dye (RB) release were first measured with the “preloaded”hydrogels in pure water solution at room temperature and all themeasurements were done in triplicate. As expected, the rates ofrelease could be tuned effectively by controlling the degree of cross-linking (Fig. 2). For all of the compositions investigated, the majorityof the dye was released in the first 7 days. (Scheme 2)

The swelling behavior of the above mentioned seven hydrogels(without dye) was studied by immersion in water at room

Fig. 3. Release profiles of post-polymerization-loaded Rhodamine 6G from 4% DSDMAcross-linked hydrogels with addition of GSH at 0 h ( ), at 6 h ( ) and in absence ofGSH in water at room temperature (C).

temperature and water uptake values (%)are given in Table 1. Asexpected, the swelling ratio of hydrogels decreased when the molefraction of cross-linker (either DEDMA or DESDMA) increased dueto an increase in the cross-linking density.

3.1. Hydrogel dye release in the absence of glutathione

Release rates of drugs that are physically entrapped withinhydrogels is known to be governed by a number of factors includingthe molecular weight of the drug, the pore size of the matrix, thedegree of cross-linking of the hydrogel, and the solvent used. In thecase of both the DEDMA and DSDMA cross-linked PHEMA hydro-gels (Fig. 2), as the mole percentage of cross-linker increased, therelease rate slowed, enabling the overall release rates to beempirically tuned by adjusting the cross-linker concentration. Inaddition, despite the different chemical linkages, both cross-linkersexhibited very similar profiles, suggesting that under redox neutralconditions the disulfide cross-linked hydrogels exhibit a similarstability to the ether cross-linked polymers.

3.2. DSDMA cross-linked hydrogel dye release in the presence ofglutathione

In order to confirm the accelerated degradation of the disulfidecross-linked hydrogels under reductive conditions, the releaseprofiles of R6G dye (loaded after polymerization), within DSDMA-

Fig. 5. Release profiles of post-polymerization-loaded Rhodamine 6G from 4% DEDMAcross-linked hydrogels with addition of GSH at 0 h ( ), at 6 h ( ) and in absence ofGSH in water at room temperature (C).

Fig. 6. Photos of representative COS-1 cell cultures after fixing and staining. Cells were cultured in the presence of eluates from triplicate polymer samples as described in Methods.A, DEDMA cross-linked polymers loaded with 0.37 mg MMC; B, DSDMA cross-linked polymers loaded with 0.37 mg MMC; C, DSDMA cross-linked control polymers (no MMC); D,DEDMA cross-linked control polymers (no MMC).

M. Ejaz et al. / Polymer 52 (2011) 5262e52705268

cross-linked hydrogels was studied after the addition of glutathione,and compared to redox-insensitive controls. It is well-known thatthe concentrations of glutathione vary throughout living organismsand as a results, can achieve selective disulfide reduction in vivo,providing a site specific triggering of hydrogel degradation. In orderto probe the utility of GSH as a trigger for accelerate release oftherapeutics from hydrogels, dye release was quantified in 1.0 mMaqueous solutions of GSH. Because RB exhibited signs of reductivedegradation in initial investigations, R6G was selected as an alter-native dye for release studies because it exhibited a stable fluores-cence over the pertinent time period in the presence of GSH.

Because of concern for the toxicity of unreacted monomer orresidual initiator for in vivo applications, the hydrogels were firstprepared without the dye, washed extensively to remove solubleresidues, and then loaded with a solution of the R6G dye as

Fig. 7. Mitomycin released from hydrogel discs is cytotoxic to COS-1 cells. Mitomycin C (0.4 mol% of DEDMA or DSDMA. The drug was eluted from the gels and its toxicity was tested wB, Culture medium with 2 mM glutathione. Data is reported relative to polymer samples w

described above. In the case of the DSDMA cross-linked hydrogels,the effect of GSH on the dye release was probed by comparing 3release studies, onewhere the GSHwas added at 0 h, onewhere theGSHwas added after 6 h and onewhere GSHwas not added (Fig. 3).In the case of the addition of glutathione at the beginning of thestudy, the accelerated release rates yielded approximately 5-foldincreases in observed dye concentrations, relative to drug releasein the absence of GSH. As additional confirmation of the GSH-triggered release, in the case of the 6 h addition of glutathione,the hydrogels first exhibited a nearly identical release profile to thatof the GSH-free study, (Fig. 3 inset) but upon addition of the GSHthe expected rapid release of the dye was observed. These resultsconfirm the utility of the DSDMA cross-linker in providing hydro-gels with redox-responsive drug release. Importantly, the rate ofGSH diffusion into the hydrogel, and the subsequent rate of GSH

37 mg) was loaded into replicate discs of hydrogel (8 mm diameter) cross-linked withith COS-1 cells as described in Methods. A, culture mediumwithout added glutathione;ith no added MMC. Plotted values represent mean þ SD (n ¼ 3).

Fig. 8. Toxicity of polymers without added MMC. Hydrogel samples with DEDMA or DSDMA cross-links were eluted sequentially with culture medium as described in Methods. Thecell density of these cultures was compared to that of cells grown in culture medium with no polymer eluate. Panel A, polymers with 4 mol% cross-linker; Panel B, polymers with8 mol% cross-linker. The plotted values represent mean þ SD (n ¼ 3).

M. Ejaz et al. / Polymer 52 (2011) 5262e5270 5269

reduction of the disulfide cross-links are sufficiently fast to affordthe observed rapid dye release.

3.3. DEDMA cross-linked hydrogel and uncross-linked PHEMAhydrogel dye release in the presence of glutathione

To confirm that the GSH reduction of the disulfide bond was themechanism of enhanced release, two controls studies were alsocarried out. The first involved the addition of GSH at 0 h and 6 h touncross-linked PHEMA hydrogels (Fig. 4). The resulting release pro-files exhibited no change in dye released regardless of GSH addition,verifying that enhanced release was not a result of non-covalentinteractions with the hydrogel. The second control involved theaddition of GSH at 0 h and 6 h to PHEMA hydrogels with 4% DEDMAcross-linker (Fig. 5). The release rates exhibited only a negligibleincreasewith the additionofGSH, suggesting that a limitedamountofester hydrolysis might occur due to the nucleophilic character of thethiol functional group within GSH, however, the observed releaseenhancement was substantially smaller than that of the disulfidecross-linker, suggesting that any contribution of ester hydrolysistoward release wasminor compared to the reduction of the disulfidecross-linker.

3.4. Test of DSDMA and DEDMA cross-linked hydrogels in a cellculture model system

Mitomycin C (MMC) is an anti-fibroproliferative agent that hasbeenusedextensively ineyesurgery to reduceexcessivepostoperativefibrosis and increase surgical success [11,12]. MMC was thereforeincorporated into theDSDMAandDEDMAcross-linkedhydrogels andits ability to inhibit the proliferation of the fibroblastic COS-1 cell lineafter release from the polymer was assessed. The test method, incu-bation of the polymer in culture medium for a given time period andsubsequent assessment of eluate toxicity, approximates in vivo drugelution into thebleb formedafter implantationof aglaucomadrainagedevice [8,9]. Preliminary experiments were carried out to determinethe MMC concentration that would provide a measureable dose-response in this cell culture system. These experiments indicatedthat addition of either 1 or 4 mg of MMC to the hydrogel discs killed>90% of the cells at all time points tested (data not shown). TheMMCconcentration was therefore reduced to w0.4 mg/disc for subsequentexperiments.

Fig. 6 shows a representative 12-well plate of COS-1 cells culturedwith eluates frompolymerdiscs. ColumnsAeDshowresponses fromtriplicate polymer samples. Even with no magnification of the fixedand stained plate, it is evident that cells grown in medium witheluate from polymer discs containing MMC (Columns A and B)

reached a much lower density than cells grown in culture mediumincubatedwith eluate fromcontrol polymers (C andD). The dye fromthese cell layers was subsequently solubilized and the absorbancefromeachwellwas used to quantify the inhibition of proliferation, asshown in Fig.7. Panel A shows the response of cells to MMC releasedimmediatelyupon immersionof thepolymer inculturemedium(day0) and to the MMC released over the indicated periods of elution.Both the DEDMA and the DSDMA cross-linked polymers were eval-uated in this experiment. Most of the MMC was released from thepolymer during the 0e3 day incubation period. These results are inagreementwith the in vitro release kinetics. Eluates from theDSDMApolymer samples were slightly more toxic at later time points (3e7and7e14days of elution). Panel B of Fig. 7 shows the responseof cellsto MMC released in the presence of 2 mM glutathione, the concen-tration of glutathione reported in fluid of the bleb created by a glau-coma drainage device [13]. There were no significant differences inthe release of MMC in the presence of 2 mM glutathione. Therefore,although these hydrogels exhibit excellent release profiles for 1 mMglutathione, they do not appear to be sufficiently sensitive to thew2 mM concentration range typical within the blebs to be of use forthe particular application investigated.

The experiments shown in Fig. 7 did not address the inherenttoxicity of the DEDMA and DSDMA polymers, because eluates frompolymer samples with no loaded MMC were used as controls forthese experiments. Experiments to test the toxicity of polymerswithout added MMC are shown in Fig. 8. In these studies, eluatesfrom control polymers with no added drug were compared to cellsgrown in culture medium without eluate. As shown in Panel A ofFig. 8, when 4 mol% cross-linker was used to prepare the polymers,eluates from the DEDMA and the DSDMA polymers showed a smallbut significant toxicity (p < 0.005) only in the 0e3 day samples.When the concentration of cross-linker was increased to 8%,eluates from the polymers showed small but significant toxicity(p < 0.005) until the 7e14 day elution interval. The DSDMA cross-linked polymers showed slightly more inherent toxicity than theDEDMA polymers in all studies. Preliminary efforts to decrease thistoxicity by increasing washing volumes and times were notsuccessful (data not shown).

4. Conclusions

The disulfide-based dimethacrylate, bis(2-methacryloxyethy)disulfide (DSDMA) cross-linker has been investigated as theresponsive component within degradable hydrogels that cantrigger enhanced release via exposure to in vivo reducing agents.The DSDMA cross-linker exhibited similar stability in non-reducingenvironments to the hardy DEDMA cross-linkers and optimization

M. Ejaz et al. / Polymer 52 (2011) 5262e52705270

of the cross-linker concentration within the hydrogel pre-polymerenabled the release rates of the hydrogels to be tuned. However, theDSDMA cross-linker also showed a rapid enhancement of dye-release upon addition of glutathione in millimolar concentrations.This reductive release mechanism is expected to be useful fora range of controlled release applications in vivo, as the concen-tration of glutathione and other biological reducing agents variessubstantial throughout the body.

Studies with cultured cells demonstrated that an anti-fibroproliferative drug, mitomycin C, could be released from thehydrogels over a 7 day period; but, themicromolar concentration ofglutathione used in these particular in vitro studies was insufficientto stimulate enhanced drug release. However, these studies alsoconfirm that glutathione can be an effective trigger, if addedexternally to these systems. This is particularly attractive, as theblebs are accessible, and glutathione exhibits low toxicity in therelevant concentration range. Additional biological studies will berequired to assure the biocompatibility of the hydrogels and eval-uate their effectiveness for other in vivo delivery applications.

Acknowledgments

This research was supported by Louisiana Board of Regents ITRSgrant to D.A. Blake (Contract Number: LEQSF(2010-12)-RD-B-05),by the Tulane Glaucoma Research Fund (R.S. Ayyala), and by TulaneUniversity’s Research Commercialization and EducationalEnhancement Program. The authors thank Niren Murthy for valu-able discussions.

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