Degradable Poly(2-hydroxyethylmethacrylate)-co-polycaprolactone Hydrogels for Tissue

Engineering Scaffolds

Sarah Atzet,† Scott Curtin,† Phalen Trinh,† Stephanie Bryant,‡ and Buddy Ratner*,†

University of Washington, 1705 Northeast Pacific Street, Box 355061, Seattle, Washington 98195, andUniversity of Colorado, 424 UCB, ECCH 118, Boulder, Colorado 80309

Received June 24, 2008; Revised Manuscript Received August 20, 2008

Biodegradable poly(2-hydroxyethyl methacrylate)(pHEMA) hydrogels for engineered tissue constructs weredeveloped by the use of atom transfer radical polymerization (ATRP), a degradable cross-linker, and a macroinitiator.Hydrogels are appropriate materials for tissue engineering scaffolds because of their tissue-like mechanicalcompliance and mass transfer properties. However, many hydrogels that have seen wide application in medicineare not biodegradable or cannot be easily cleared from the body. pHEMA was selected for the scaffold materialbecause of its reasonable mechanical strength, elasticity, and long history of successful use in medicine as wellas because it can be easily fabricated into numerous configurations. pHEMA was studied at various molecularweights between 2 and 50 kDa. The molecular weight range suitable for renal clearance was an important factorin the experimental design. The fabricated hydrogels contain oligomeric blocks of polycaprolactone (PCL), ahydrolytically and enzymatically degradable polymer, as a cross-linking agent. In addition, a degradablemacroinitiator that also contained oligomeric PCL was used to initiate the ATRP. The chain length, cross-linkdensity, and polymerization solvent were found to affect the mechanical properties of the pHEMA hydrogels.Degradation of the pHEMA hydrogels was characterized by the use of 0.007 M NaOH, lipase solutions, andphosphate-buffered saline. The mass loss, swelling ratio, and tensile modulus were evaluated. Degradation productsafter sodium hydroxide treatment were measured by the use of gel permeation chromatography (GPC) to verifythe polymer lengths and polydispersity. Erosion was observed in only the sodium hydroxide and lipase solutions.However, the swelling ratio and tensile modulus indicate bulk degradation in all PCL-containing samples.Degradable hydrogels in enzymatic solutions showed 30% mass loss in 16 weeks. Initial cell toxicity studiesindicate no adverse cellular response to the hydrogels or their degradation products. These hydrogels haveappropriate mechanical properties and a tunable degradation rate, and they are composed of materials that arecurrently in FDA-approved devices. Therefore, the degradable pHEMA developed in this study has considerablepotential as a scaffold for tissue engineering applications, in cardiac and other applications.

Introduction

This study addresses scaffold polymers that will be used inheart muscle tissue engineering. Heart failure and relatedcardiovascular disease continues to be the leading cause of deathin the United States.1 Myocardial infarctions result fromcoronary artery blockages and often lead to heart failure.Because cardiomyocytes have limited regenerative capability,tissue damaged by an acute myocardial infarction is replacedwith nonfunctional scar tissue. Despite advances in pharmaco-logical, interventional, and surgical therapies, the prognosis forpatients with heart failure is unfavorable.2 Consequently,research has focused on regenerating functional myocardium,often by the use of tissue engineered scaffolds. It has beenproposed that scaffolds seeded with functional cardiomyocytescan be placed in the infracted region to restore viablemyocardium.3-6 The role of the scaffold in this process is tosupport and direct 3D cellular growth, leading to restored tissue.Compared with the direct injection of cells, a seeded scaffoldallows for a higher cell density, and more importantly, itlocalizes injected cells.7 As cells develop and lay down the

extracellular matrix, the scaffold should begin to degrade. Therate of degradation can affect the macroscopic shape and thetimely development of new tissue.8,9 Therefore, it is importantto have a scaffold material with a tunable degradation rate. Asuitable scaffold for cardiac tissue engineering should also bebiocompatible, integrate with the host tissue, and exhibit tissue-like mechanical properties.10 Additionally, to ensure that scaffolddegradation products can easily egress by renal clearance themolecular weight should be <50 kDa and products should besoluble in the bloodstream.11

The majority of the scaffolds used in tissue engineeringstudies published to date have been based on poly(glycolic acid)(PGA), poly(lactic acid) (PLA), and their copolymers. Althoughthey have a long history in medicine and a wide acceptance,these polymers exhibit a number of disadvantages for theirapplication as tissue engineering scaffolds, especially for heartmuscle tissue engineering: (1) They are not elastomeric and thusdo not match the modulus of soft tissue. (2) They break downinto acidic products that are incompatible with cell growth. (3)The acidic breakdown products autocatalyze further polymerbreakdown, often leading to catastrophic disintegration of largermasses of polymer. (4) They are hydrophobic. (5) They aredifficult to derivatize chemically, which makes surface im-mobilization problematical. These issues are addressed in thenew polymer that was developed in this work.

* To whom correspondence should be addressed. E-mail: ratner@uweb.engr.washington.edu. Fax: 206-616-9763.

† University of Washington.‡ University of Colorado.

Biomacromolecules 2008, 9, 3370–33773370

10.1021/bm800686h CCC: $40.75 2008 American Chemical SocietyPublished on Web 12/08/2008

Methacrylate and acrylate polymers are widely used inmedicine and biology because they are well tolerated in vivo.Several methacrylates and acrylates with hydrophilic substituentgroups such as poly(2-hydroxyethyl methacrylate) (pHEMA)can form hydrogels that, because of the water content andfavorable mechanical properties, have found use in applicationssuch as contact lenses, drug delivery vehicles, and tissueengineering scaffolds.12 Hydrogels of pHEMA meet several ofthe scaffold requirements. They elicit an in vivo response thatis considered biocompatible, they can be fabricated in variousarchitectures, and they have mechanical properties that aresimilar to those of natural tissue.13-16 However, pHEMA hasnot been previously used as a bioresorbable scaffold becauseof its biostablility. Polycaprolactone (PCL) is a semicrystallinepolyester that contains aliphatic ester linkages that are suscep-tible to both hydrolytic and enzymatic degradation.17-19 Ex-tensive research has been performed with PCL that has led toFDA approval of several PCL-containing medical and drugdelivery devices.20,21 The degradation kinetics of PCL areconsiderably slower than other aliphatic polyesters because ofits hydrophobicity and crystallinity. This slow degradation canbe desirable for scaffolding material because it may reflect therate of tissue development.

Atom transfer radical polymerization (ATRP) can be usedto prepare controlled molecular weight pHEMA with lowpolydispersity, and the water solubility of that pHEMA has beenevaluated.23 ATRP is a versatile technique that can be used tocontrol the polymerization of several monomer classes includingmethacrylates.22,23 ATRP involves polymer radicals, as opposedto ionic species, and therefore is well suited for the polymer-ization of functional monomers such as pHEMA.24 Thistechnique is used here to ensure that the degradation productsare appropriately small for excretion from the body.

In this study, we have combined pHEMA and PCL by usingthe ATRP technique and advanced macromolecular design ideasto develop a scaffold that can meet all of the aforementionedrequirements. The specific rationale for the polymers that aredescribed in this paper is as follows: (1) For pHEMA to becleared after biodegradation, fragments must be smaller than10 kDa (5 kDa is preferable). (2) To provide hydrolytically labilesites for the breakdown of the stable pHEMA, dimethacrylatedoligo-PCL segments were incorporated. (3) When pHEMAsegments were 5 kDa, poor mechanical properties were observedin the resulting polymer. Therefore, a difunctional PCL oligomerthat was terminated with ATRP initiator was used to createpHEMA units with molecular weights of 10 kDa plus themolecular weight of the backbone PCL unit. Here, we describethe synthesis, characterization, and degradation of low-molec-ular-weight pHEMA that is initiated and cross-linked with PCLmoieties.

Materials and Methods

Materials. 2-Hydroxyethyl methacrylate monomer (ophthalmologicgrade, 99% minimum purity, acid content maximum ) 0.05%, cross-linker content maximum ) 0.15%) (HEMA) and tetraethylene glycoldimethacrylate (TEGDMA) were purchased from Polysciences (War-rington, PA) and were used without further purification. PCL diols (MW) 530, 1250, 2000), methacryloyl chloride (97%), copper(I) chloride,2,2′-bipyridyl, ethyl 2-bromoisobutyrate (EBiB), and R-bromoisobutyrylbromide were purchased from Sigma-Aldrich (Milwaukee, WI) andwere used without further purification. Amano lipase PS and MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) were alsopurchased from Sigma, along with all other reagent-grade solvents.

Instrumentation. 1H NMR spectra were recorded on a BrukerAvance series instrument (300 MHz) using a Bruker BBI probe. ATosoh Corp gel permeation chromatograph was used with TSRGelR-3000 and R-4000 columns eluted at 1 mL/min. Dimethylformamidecontaining 1 wt % lithium bromide was used as the solvent, andpoly(methyl methacrylate) standards were used for calibrating themolecular weights. Mechanical properties of the hydrogels such as thetensile modulus, ultimate tensile strength, and ultimate tensile strainwere evaluated by the use of an Instron universal tester 3340 seriessingle column system that was equipped with a 10 N load cell.

Polycaprolactone Cross-Linker Synthesis. Synthesis of the PCLcross-linker was performed as previously reported by Sawhney et al.25

We reacted an oligomeric caprolactone diol (MW ) 530) withmethacryloyl chloride to produce a degradable PCL cross-linker, PCLX.Figure 1 shows the reaction scheme for the synthesis of PCLX. Briefly,PCL diol (25 g, 0.047 mol, MW ) 530) was dissolved in dichlo-romethane (100 mL) in a round-bottomed flask. Triethylamine (19.7mL, 0.14 mol) was added directly to the flask, and the solution wascooled to 0 °C. Methacryloyl chloride (13.8 mL, 0.14 mol) anddichloromethane (50 mL) were placed in an addition funnel over theround-bottomed flask. The air was removed and replaced with nitrogen.The methacryloyl chloride solution was added dropwise to the stirredsolution over a 2 h period. The solution was allowed to react for 12 hat 0 °C and then for 24 h at room temperature. The triethylaminehydrogen chloride salts were removed by filtration, and then the solventwas removed under vacuum. The remaining oil was precipitated intocold hexanes to yield a slightly yellow wax. The purchased PCL diol(MW ) 530) has an average n of 2.3. The product was evaluated byNMR (Figure 1), and the chemical shifts were obtained relative totetramethylsilane. 1H NMR (300 MHz, CDCl3, δ): 6.13 (b), 5.58 (a),4.28 (1), 4.08 (e), 3.68 (2), 2.35 (j), 1.95 (c), 1.68 (h), 1.39 (g).

Polycaprolactone Initiator Synthesis. An ATRP macroinitiator wasalso synthesized to improve mechanical properties of the hydrogel whilemaintaining low-molecular-weight degradation products. We synthe-sized the PCL macroinitiator (PCLI) by reacting an oligomericcaprolactone diol with R-bromoisobutyryl bromide, which resulted ina difunctional initiator.26 PCL diol (MW ) 530, 10 g, 0.02 mol) wasdissolved in anhydrous tetrahydrofuran (300 mL). Triethylamine (13.1mL, 0.1 mol) was added to the stirred solution, followed by the addition

Figure 1. PCLX reaction scheme. PCL diols of various molecular weights (530, 1250, and 2000 g/mol) were reacted with methacryloyl chlorideto produce a dimethacrylated PCL moiety that was used as a cross-linker (PCLX).

pHEMA/PCL Hydrogels Biomacromolecules, Vol. 9, No. 12, 2008 3371

of R-bromoisobutyryl bromide (5.8 mL, 0.05 mol). The reaction wasallowed to proceed overnight at ambient temperature. The resultinghydrogen chloride salt was removed by filtration, and the THF wasremoved under vacuum. The residual yellow oil was redissolved indichloromethane (300 mL) and was then washed three times with asaturated solution of bicarbonate. The organic layer was separated, driedover magnesium sulfate, and filtered, and the dichloromethane wasremoved under vacuum. The final product, a yellow oil, was character-ized by NMR.26 Figure 2 shows the reaction scheme for the synthesisof PCLI. 1H NMR (300 MHz, CDCl3, δ): 4.31 (1), 4.11 (e), 3.72 (2),2.35 (j), 1.95 (d), 1.65 (h), 1.53 (f), 1.25 (g).

Preparation and Characterization of Scaffold and Linear Poly-mer. The synthesis of linear pHEMA for solubility and macroinitiatorstudies was done according to Weaver et al.23 In a typical experiment,HEMA monomer (7.16 mmol, 0.9 mL), the initiator EBiB (0.094 mmol,14 µL), and methanol (1 mL) were degassed with nitrogen for 30 min.A catalyst solution containing CuCl (0.096 mmol, 9.5 mg), 2,2′-bipyridyl (0.24 mmol, 37.5 mg), and methanol (100 µL) was alsodegassed and was then added under nitrogen to the previous monomersolution. The methanolic solution was stirred for 24 h, and upon expo-sure to air, the solution changed from dark brown to blue, indicatingthe oxidation of the copper. Additional methanol (5 mL) was addedto the polymer solution before it passed through a silica column toremove the catalyst. Excess methanol was removed by vacuum, andthen the reduced volume solution was precipitated into cold diethylether. The white precipitate was then collected and dried under vacuumfor 12 h. pHEMA hydrogels with various molecular weight blocksegments and cross-link densities were prepared by the use of ATRPtechniques. The initiator (PCLI) and cross-linker (PCLX) that wereused for the degradable gels were synthesized as described above. Thefollowing experiments were all conducted on hydrogels that wereprepared by the use of PCLX that was synthesized from PCL diol (MW)530 g/mol). EBiB and TEGDMA were used in the nondegradablecontrols as the initiator and the cross-linker, respectively. The ATRPtransition metal and ligand that were used for both polymerizationswere copper(I) chloride and 2,2′-bipyridyl.

For a typical polymerization to form the degradable pHEMAhydrogels (MW ) 10 kDa), PCLI (0.46 mmol, 0.4 g) and PCLX (1.16mmol, 1.0 g) were dissolved in dimethylformamide (1.5 mL). HEMA(35.8 mmol, 4.5 mL) and ethylene glycol (3.0 mL) were then added tothe solution, and the solution was purged with nitrogen for 30 min.We separately prepared the catalyst solution by dissolving CuCl (0.93mmol, 92.1 mg) and 2,2′-bipyridyl (2.33 mmol, 363 mg) in dimeth-ylformamide (0.3 mL) and ethylene glycol (0.7 mL). The catalystsolution was also purged with nitrogen for 30 min. In a nitrogenenvironment, the two solutions were mixed, cast into molds, andallowed to polymerize for 24 h. The hydrogels were then removed fromthe molds and rinsed several times with acetone/water (90:10) to removecopper and unreacted monomer. We note that the term cross-linkingdensity will be used in this article to refer to the mol percent of cross-linking agent in the initial polymerization solution.

The uncrosslinked pHEMA and hydrogel degradation products werecharacterized by gel permeation chromatography (GPC). Mechanicalproperties were evaluated from dog-bone-shaped samples (4.5 × 0.65× 20 mm3) that were punched from a film. The crosshead speed wasset to stretch at 10% strain/min until failure. At least three sampleswere used for each measurement.

Scaffold Degradation and Cytotoxicity. We evaluated in vitrodegradation profiles and rates by measuring the percent mass loss andthe mass swelling ratio. We determined the percent mass loss andswelling ratio by the using the following equations

% mass loss)mo -mD

mo× 100

swelling ratio)mw

mD

where mo is the original dry mass of the sample, mD is the residual drymass of the sample after a degradation period, and mw is the mass ofthe hydrated sample after a degradation period.6

Degradation was studied in several environments including 0.007M NaOH, 1.0 mg/mL lipase, 0.5 mg/mL lipase, and phosphate-bufferedsaline (PBS). Before degradation studies, all hydrogels were lyophilizedand weighed to determine the original dry weight. In the accelerateddegradation study, samples were place in 0.007 M NaOH at roomtemperature and were placed on an orbital shaker that rotated atapproximately 60 rpm. Samples were taken twice a day for 5 days,and the hydrated weight was measured before the samples were frozenand lyophilized. The dry weight was again recorded for use incalculations. The degradation studies in lipase and PBS solutions wereconducted in a similar fashion, except the temperature was controlledat 37 °C, and samples were drawn over a 20-week time period.Degradation media were changed every 4 days to prevent contaminationand to ensure enzyme activity. The Young’s modulus of the hydrogelswas measured over the degradation time. We determined the modulusby examining the slope of the linear region of the stress versus straincurve (<10% strain).

The solubility of the degradation products was evaluated in a 0.15M NaCl solution. In each case, polymer (20 mg) was dissolved in 2mL of solution and was stirred at 0-5 °C for 24 h. The solution wasthen filtered, and the remaining insoluble material was vacuum driedand weighed.23

The cytotoxicity of the hydrogels and their degradation products wasmeasured by an MTT assay. Hydrogels were soaked in media for 24 hbefore the eluent was removed and placed on 3T3 mouse fibroblaststhat were plated 24 h earlier at a concentration below confluency. Thecells were then evaluated at 12, 24, and 48 h time points. Thedegradation products were dissolved in media at concentrations rangingfrom 5 to 15 mg/mL and were evaluated in a procedure that was similarto that of the hydrogel eluent. Tissue culture polystyrene (TCPS) andlatex were used as the negative and positive controls for cytotoxicity,respectively.

Figure 2. PCLI reaction scheme. A PCL diol (MW ) 530) was reacted with R-bromoisobutyryl bromide to produce a difunctional macroinitiator(PCLI).

3372 Biomacromolecules, Vol. 9, No. 12, 2008 Atzet et al.

Results and Discussion

Figures 1 and 2 show the reaction scheme for both PCLXand PCLI production. We calculated the extent of methacrylationof the PCL diol by comparing the integral area under the NMRpeaks at δ 6.13 (methacrylate functional group) to 3.68 (PCLdiol backbone). We determined the functionality of the initiatorby comparing the peak areas of δ 1.95 (bromine functionalgroup) to 3.72 (PCL diol backbone). The functionality of thePCLX and PCLI were determined to range from 80-95 and70-85%, respectively. These percentages closely correspondto those reported by Sawhney et al. and Rice et al. for similarreaction schemes.18,25,27 Additionally, the comparison of HNMR shifts between the initial PCL diol and either PCLX orPCLI shows the disappearance of the peak at 3.53, whichcorresponds to OH terminals. This again confirms that the endgroup functionality of the PCL has been converted from a diolto either a dimethacrylate or diorganohalide.

To verify the living nature of the ATRP/HEMA system andthe ability to synthesize pHEMA with low molecular weights,we examined various monomer-to-initiator ratios by utilizingethyl 2-bromoisobutyate as the initiator. Table 1 lists themonomer-to-initiator ratios, theoretical molecular weights, andmeasured molecular weights (GPC) for several polymerizationsof linear pHEMA.

As Table 1 indicates, the molecular weight measured by GPCis significantly higher than the theoretical molecular weight;however, discrepancies in the low-molecular-weight range havebeen reported as being systematic errors in the GPC analysis.23

Increased polydispersities and standard deviations were observedat higher molecular weights. This is thought to be partially dueto impurities in the HEMA monomer. Additionally, measure-ment errors for low molecular weights (<5 kDa) may bepartially due to the fact that this approaches the lower limits ofthe GPC column. Polydispersity (PDI) was measured to be aslow as 1.17, which indicates low variance in the molecularweight of the pHEMA chains.

The molecular weight of the polymer is an important factorin the ultimate excretion. There are literature reports thatmolecular weights that are suitable for renal clearance are inthe range of 10-50 kDa.11,28 This corresponds to monomer-to-initiator ratios of 75:1 to 375:1. The linear correlation betweenmeasured and theoretical molecular weight, as shown in Figure3, indicates that the polymerization is controlled.

Water solubility is another significant factor for renal clear-ance. As shown by Figure 4, there is a significant decline inwater solubility as molecular weight increases. A molecularweight of >5 kDa (monomer-to-initiator ratio 38:1) results inlimited water solubility. Taking into consideration the GPCsystematic errors, this solubility limit is in agreement with thatreported by Weaver et al.23 Unfortunately, the low molecularweights required for renal clearance and water solubility producehydrogels that do not have mechanical properties that are inthe natural tissue range. Therefore, a macroinitiator was syn-

thesized to allow the backbone chain length to be double thelength of the degradation products.

Polymerizations of various molecular weights were performedto evaluate the effectiveness of the macroinitiator PCLI. GPCwas used to evaluate the chain length before and after thepolymer was placed in 1 M NaOH to accelerate degradation(incubation time 24 h). As shown by Figure 5, the molecularweight of the degraded polymer chain is half the original length.This indicates that the macroinitiator is indeed difunctional anddoes not affect the ability to control the polymerization. Thecontribution of the PCL in the polymer chain is small and thusfalls within the standard error of the measurements. Therefore,the hydrogels used for degradation experiments were polym-erized in a ratio of 75:1, which has a predegradation theoreticalbackbone length of 9.76 kDa and a degraded backbone lengthof 4.88 kDa.

Hydrogels were polymerized with cross-linker concentrationsthat ranged from 0.8 to 13.5 mol % and backbone lengths thatranged from 10 to 50 kDa. Mechanical properties were foundto be greatly influenced by the polymerization solvent, cross-linking density, and monomer-to-initiator ratio. The minimumcross-linking density is dependent on the monomer-to-initiatorratio used. A minimum of two cross-links per chain was usedto ensure the formation of an elastic hydrogel. Therefore, for a

Table 1. Summary of Monomer-to-Initiator Ratios Used and theCorrelating Theoretical and Measured Molecular Weight for LinearpHEMA Initiated with Ethyl 2-Bromoisobutyrate

Mo:I ratio Mn (theory) Mn (GPC) PDI

15:1 2000 6000 ( 120 1.17 ( 0.00930:1 4000 8200 ( 270 1.22 ( 0.00560:1 8000 11600 ( 500 1.26 ( 0.008120:1 16000 17600 ( 730 1.35 ( 0.05240:1 32000 33500 ( 3800 1.71 ( 0.05360:1 48000 46500 ( 4800 2.17 ( 0.2

Figure 3. The molecular weight of linear pHEMA initiated with ethyl2-bromoisobutyrate as measured by GPC versus the theoreticalmolecular weight calculated from the monomer-to-initiator ratio. Thelinear correlation indicates a controlled polymerization.

Figure 4. Single pass solubility of linear pHEMA initiated with ethyl2-bromoisobutyrate in 0.15 M NaCl versus the theoretical molecularweight calculated from monomer-to-initiator ratios.

pHEMA/PCL Hydrogels Biomacromolecules, Vol. 9, No. 12, 2008 3373

backbone chain length of 10 kDa, the minimum cross-linkingdensity was calculated to be approximately 4 mol %. Figure 6shows a schematic of the hydrogel polymerized with PCLI andPCLX.

The Young’s modulus of the hydrogel was measured forseveral different polymerizations. The polymerization detailsand the resulting Young’s moduli are shown in Table 2. Asexpected, significant increases in the Young’s modulus resultfrom increasing the cross-linking density, cross-linking withPCLX instead of TEGDMA, and by decreasing the molecularweight (Table 2). The modulus of all of the measured sampleswas comparable to the modulus of rat heart tissue and pHEMApolymerized with a TEGDMA cross-linker. The incorporationof the PCLX nearly doubles the tensile modulus of the material,increasing it from 0.29 to 0.43 MPa (experiments 1 and 2). Thiseffect could be due to the fact that PCL can crystallize, andthis increase in crystallinity may be contributing to the increasedtensile modulus. Further studies will be performed that usedifferential scanning calorimetry to elucidate this effect. Dou-bling the cross-linking density also doubles the tensile modulus(experiments 2 and 3). It should be noted that the modulus isalso significantly impacted by the type and amount of solventthat is used in the polymerization. Increasing the amount ofethylene glycol in the polymerization by 50% resulted in adecrease in the modulus from 0.76 ( 0.05 to 0.43 ( 0.02 MPa(experiments 2 and 5). All hydrogels were subjected to a pretestcycle that extended the sample to 10% strain followed byrelaxation three times. Hysteresis was observed on only the firstcycle, indicating that after an initial conditioning, the hydrogels

exhibit elastic behavior. Conventional pHEMA was polymerizedby a free-radical mechanism that used TEGDMA cross-linkerand ammonium persulfate and sodium bisulfate as the initiatorsystem. Unfortunately, reliable measurements could not beobtained from hydrogels with cross-linking densities of 13.5mol % (approximate mass equivalent to the PCLX) because ofcracking and therefore were not included in the subsequent table.

We investigated the degradation of the hydrogels in sodiumhydroxide solutions of varying concentrations, enzymatic solu-tions, and PBS by following the changes in the weight loss,swelling ratio, and tensile modulus. The hydrogels undergo bulkdegradation uniformly throughout the gels. As PCL cross-linksor chain segments are cleaved, the effective cross-link densitywill be reduced, which increases the polymer mesh size, lowersthe tensile modulus, and increases the swelling ratio or watercontent of the hydrogel. When enough PCL bonds have beenbroken to free a pHEMA backbone segment, the chain will gointo solution, and the mass of the hydrogel will decrease.

Accelerated degradation studies were performed primarily toensure that the hydrogel’s hydrolytically labile bonds areaccessible and that the degradation products will be soluble.This was confirmed by degradation in 0.007 M NaOH at roomtemperature. The monomer-to-initiator ratio for these hydrogelswas 75:1 (MW ) 10 kDa), the cross-link density varied from4.5 to 13.5 mol %, and the solvent-to-monomer ratio was keptconstant at 1:3. The degradation profile is shown in Figures 7and 8 as assessed by the swelling ratio and the percent massloss. Degradation profiles followed the expected trend. Thelowest cross-link density hydrogel completely degraded in under24 h, whereas the gel with the highest cross-link density didnot completely degrade within the experimental time frame. Thisverifies that the cross-link density can be used to tailor thedegradation rate of the hydrogel. The nondegradable control,which was initiated with EBiB and cross-linked with TEGDMA,showed no statistically significant mass loss or change inswelling ratio.

Experiments in NaOH confirm degradation, but to match thein vivo environment more closely, we examined degradationin lipase solutions of 1.0 and 0.5 mg/mL and in PBS. Themonomer-to-initiator ratio for these hydrogels was 75:1 (MW) 10 kDa), the cross-link density varied from 4.5 to 13.5 mol%, and the solvent-to-monomer ratio was kept constant at 1:3.Figures 9 and 10 illustrate how the swelling ratio and tensilemodulus change over degradation time.

Degradation, as measured by swelling and weight loss, isobserved in all media for the PCLX hydrogels, whereas thereis no statistical difference for control samples (cross-linked withTEGDMA). Both the swelling ratio and the tensile modulusfor the degradable hydrogels vary linearly with the degradationtime with correlation values above 0.95. For PCLX/PCLIhydrogels, the swelling ratio increases from 1.4 to 2.0 over 16weeks, whereas nondegradable hydrogels start at 1.8 and remainconstant over that time period. The nondegradable hydrogelswelling ratio is initially higher because the PCLX and PCLIcontribute hydrophobic moieties that lower the water uptakeby the hydrogel. As expected, the degradation that is observedby monitoring the swelling ratio is more pronounced in theenzymatic solutions, and the rate correlates with the enzymeconcentration. The tensile modulus for degradable hydrogelsdecreases from 0.4 to 0.05 MPa over 16 weeks. Again, nostatistical difference is observed in the nondegradable TEGDMA/EBiB hydrogels. From these data, it is clear that the hydrogelis subjected to bulk degradation because there are measurableinitial changes in properties for all degradable samples. Figures

Figure 5. Molecular weight as measured by GPC for linear pHEMAinitiated with PCLI at various monomer-to-initiator ratios both beforeand after exposure to 1 M NaOH.

Figure 6. Schematic of pHEMA hydrogel cross-linked and initiatedwith PCLX (- - -) and PCLI (---). pHEMA backbone chain lengthcan be varied from 5 to 20, and a minimum of two cross-links perchain is necessary for hydrogel formation.

3374 Biomacromolecules, Vol. 9, No. 12, 2008 Atzet et al.

11 and 12 show the mass loss for control and degradablehydrogels, respectively.

There is a significant lag time in the mass loss where thehydrogel properties are changing but no erosion has occurred.Day 42 is the first time point in which a statistically significantmass loss can be measured. After that time, the mass lossproceeds quickly until the gels completely dissolve in thedegradation media. Again, there is no statistically measurablemass loss for the control samples. The mass loss is observed

for only the degradable samples in the enzymatic solutions. At16 weeks the PCLX/PCLI hydrogels that were degraded in 1.0mg/mL lipase have lost 30% of their original weight. As shownby the accelerated degradation profiles, the hydrogel rapidlyerodes after this point, and reliable measurements are difficult.Whereas it is clear from swelling studies and tensile measure-ments that the hydrogel is subjected to bulk degradation, thesamples that were exposed to enzymatic solutions may also beundergoing surface degradation, leading to measurable massloss. The enzymatically accelerated degradation is likely a

Table 2. Summary of Experimental Factors Affecting the Tensile Modulus of the Hydrogelsa

expt no. Mo:I ratio theoretical Mn (Da) cross-linker cross-link density (mol %) ethylene glycol (mL) tensile modulus (MPa)

1 75:1 10 000 TEGDMA 4.5 3 0.29 ( 0.032 75:1 10 000 PCLX 4.5 3 0.43 ( 0.023 75:1 10 000 PCLX 9 3 0.83 ( 0.074 75:1 10 000 TEGDMA 4.5 2 0.48 ( 0.035 75:1 10 000 PCLX 4.5 2 0.76 ( 0.056 150:1 20 000 TEGDMA 2.5 2 0.31 ( 0.027 150:1 20 000 PCLX 2.5 2 0.49 ( 0.03

conventional pHEMAb TEGDMA 1.8 0.63 ( 0.04heart tissue (rat)c 0.59 ( 0.22scar tissuec 4.7 ( 1.4

a Cross-linker types TEGDMA and PCL dimethacrylate (PCLX) were used. b Conventional pHEMA is polymerized with ammonium persulfate andsodium bisulfate. c Heart and scar tissue provided by Steve Korte, Alicia Gonzalez, and Michael Reigner.

Figure 7. Swelling ratio as a function of time for pHEMA hydrogelscross-linked with 4.5, 9.0, and 13.5 mol % PCLX and 4.5 mol %TEGDMA in a 7 mmol solution of NaOH.

Figure 8. Percent mass loss as a function of time for pHEMAhydrogels cross-linked with 4.5, 9.0, and 13.5 mol % PCLX and 4.5mol % TEGDMA in a 7 mmol solution of NaOH.

Figure 9. Swelling ratio as a function of time for pHEMA hydrogelscross-linked and initiated with PCLX/PCLI or TEGDMA/EBiB insolutions of 1.0 and 0.5 mg/mL lipase and PBS.

Figure 10. Tensile modulus of pHEMA hydrogels cross-linked andinitiated with either PCLX/PCLI or TEGDMA/EBiB in 1.0 mg/mL lipaseand PBS.

pHEMA/PCL Hydrogels Biomacromolecules, Vol. 9, No. 12, 2008 3375

combination of surface and bulk erosion because the diffusionof lipase into the bulk of the hydrogel is limited by the size ofthe enzyme. The mass loss varies exponentially with time, andcorrelation coefficients for the two enzymatic experiments werefound to be 0.85 and 0.95 for 1.0 and 0.5 mg/mL, respectively.This slow rate of mass loss is not unexpected. Kwoen et al.reported a weight loss of 10% at 40 days for PCL and PCLnetworks in PBS.17 However, Rice et al. reported the completedegradation of PEG-PCL hydrogels in 1.0 mg/mL lipasesolutions in <9 days.18 The swelling ratios of the PEG-PCLhydrogels in that study were 10-fold higher than those of thesepHEMA gels, and the PCL segments were presumably moreaccessible to hydrolysis and the hydrolytic enzyme. The enzymeconcentrations used were not selected on the basis of knownesterase concentrations in vivo and thus may have differentactivities than what will be encountered intramuscularly. Ad-ditionally, to address the slow rate of degradation, current studiesare investigating the use of PLA and PGA as a replacement forthe PCL segments in the degradable cross-linker. Preliminaryresults for PLA cross-linked hydrogels indicate a much fasterdegradation compared with PCL cross-linked hydrogels. Furtherinvestigations into tuning the degradation rate are in progress.

Prior to in vivo experiments, the cytoxicity of the hydrogelsand their degradation products must be evaluated. The MTTassay is a colorimetric method that is used to measure cellproliferation and cytotoxicity. Mitochondria of live cells reducethe yellow tetrazole to a purple formazan that can be solubilizedby acidic isopropanol. The absorbance of this colored solutioncan then be used to quantify the number of live cells. For thisstudy, the absorbance values of all samples were normalized tothe negative control (TCPS). The hydrogels tested were notcytotoxic to fibroblasts, and the degradation products showedno cytotoxicity at any of the concentrations examined. Cyto-toxicity was evaluated at time points of 12, 24, and 48 h. Nostatistical difference was observed in this time period. Figure13 shows the normalized absorbance for the three testedconcentrations as well as the positive (latex) and negativecontrols at 24 h. These results justify the examination of thesepolymers in in vivo models for degradation, and this will bethe subject of a future study.

Conclusions

A novel degradable hydrogel has been synthesized andcharacterized. The addition of degradable PCL segments in boththe initiator and cross-linker as well as control over the pHEMAmolecular weight resulted in a bioresorbable pHEMA hydrogel.The degradation rate of this material can be controlled by thecross-linking density, the PCL chain length, and the pHEMAbackbone chain length. The hydrogel exhibited mechanicalproperties that are suitable for cardiac tissue engineeringscaffolds. Degradation products are soluble in solutions withionic strength similar to that of blood, and we expect that renalclearance take place. Bulk degradation was observed over aperiod of 16 weeks and was seen in only enzymatic solutions.Because of the long history of the successful application ofmethacrylate polymers in medical devices, we envision this newmaterial to be readily accepted and to have applications in tissueengineering and drug delivery systems.

Acknowledgment. This work was supported by the NationalInstitutes of Health (grant no. HL64387). We thank Steve Korte,Alicia Gonzalez, and Michael Reigner for heart and scar tissuemechanical data.

References and Notes(1) Heart Disease and Stoke Statistics-2006 Update. In Circulation;

American Heart Association: Dallas, 2006; pp. e85-e101.

Figure 11. Percent mass loss versus time for pHEMA hydrogelscross-linked and initiated with TEGDMA/EBiB in 1.0 and 0.5 mg/mLlipase and PBS.

Figure 12. Percent mass loss versus time for degradable pHEMAhydrogels cross-linked and initiated with PCLX/PCLI in 1.0 and 0.5mg/mL lipase and PBS.

Figure 13. Cytotoxicity, as measured by MTT, of degradation productsat 5, 10, and 15 mg/mL. TCPS and latex were used as the negativeand positive controls, respectively, for cytotoxicity. Absorbance hasbeen normalized to the value of TCPS.

3376 Biomacromolecules, Vol. 9, No. 12, 2008 Atzet et al.

(2) Caspi, O.; Lesman, A.; Basevitch, Y.; Gepstein, A.; Arbel, G.; Habib,I. H.; Gepstein, L.; Levenberg, S. Circ. Res. 2007, 100, 263–272.

(3) Leor, J.; Gerecht, S.; Cohen, S.; Miller, L.; Holbova, R.; Ziskind, A.;Shachar, M.; Feinberg, M. S.; Guetta, E.; Itskovitz-Eldor, J. Heart2007, 93, 1278–1284.

(4) Leor, J.; Aboulafia-Etzion, S.; Dar, A.; Shapiro, L.; Barbash, I. M.;Battler, A.; Granot, Y.; Choen, S. Circulation 2000, 102, III56–III61.

(5) Hidalgo-Bastida, L. A.; Barry, J. J.; Everitt, N. M.; Rose, F. R.; Buttery,L. D.; Hall, I. P.; Claycomb, W. C.; Shaksheff, K. M. Acta Biomater.2007, 3, 457–462.

(6) Eschenhagen, T.; Zimmermann, W. H. Circ. Res. 2005, 97, 1220–1231.

(7) Park, H.; Cannizzarro, C.; Vunjak-Novakovic, G.; Langer, R.; Vacanti,C. A.; Farokhzad, O. C. Tissue Eng. 2007, 13, 1867–1877.

(8) Bryant, S. J.; Anseth, K. S. J. Biomed. Mater. Res. 2002, 59, 63–72.(9) Rice, M. A.; Anseth, K. S. J. Biomed. Mater. Res., Part A 2004, 70,

560–568.(10) Nair, L. S.; Laurencin, C. T. Prog. Polym. Sci. 2007, 32, 762–798.

(11) Healy, J. M.; Lewis, S. D.; Kurz, M.; Boomer, R. M.; Thompson,K. M.; Wilson, C.; McCauley, T. G. Pharm. Res. 2004, 21, 2234–46.

(12) Metters, A. T.; Anseth, K. S.; Bowman, C. N. J. Phys. Chem. B 2001,105, 8069–8076.

(13) Homsy, C. A. J. Biomed. Mater. Res. 1970, 4, 341–356.

(14) Peppas, N. A.; Moynihan, H. J.; Lucht, L. M. J. Biomed. Mater. Res.1985, 19, 397–411.

(15) Ratner, B. D.; Miller, I. F. J. Biomed. Mater. Res. 1973, 7, 353–367.

(16) Bakshi, A.; Fisher, O.; Dagci, T.; Himes, B. T.; Fischer, I.; Lowman,A. J. Neurosurg. Spine 2004, 1, 322–329.

(17) Kweon, H.; Yoo, M. K.; Park, I. K.; Kim, T. H.; Lee, H. C.; Lee,H. S.; Oh, J. S.; Akaike, T.; Cho, C. S. Biomaterials 2003, 24, 801–808.

(18) Rice, M. A.; Sanchez-Adams, J.; Anseth, K. S. Biomacromolecules2006, 7, 1968–75.

(19) Tay, B. Y.; Zhang, S. X.; Myint, M. H.; Ng, F. L.; Chandrasekaran,M.; Tan, L. K. A. J. Mater. Process. Technol. 2007, 182, 117–121.

(20) Woodward, S. C.; Brewer, P. S.; Moatamed, F.; Schindler, A.; Pitt,C. G. J. Biomed. Mater. Res. 1985, 19, 437–444.

(21) Darney, P. D.; Monroe, S. E.; Klaisle, C. M.; Alvarado, A. Am. J.Obstet. Gynecol. 1989, 160, 1292–1295.

(22) Beers, K. L.; Boo, S.; Gaynor, S. G.; Matyjaszewski, K. Macromol-ecules 1999, 32, 5772–5776.

(23) Weaver, J. V. M.; Bannister, I.; Robinson, K. L.; Bories-Azeau, X.;Armes, S. P. Macromolecules 2004, 37, 2395–2403.

(24) Matyjaszewski, K.; Xia, J. Chem. ReV. 2001, 101, 2921–2990.(25) Sawhney, A. S.; Pathak, C. P.; Hubbell, J. A. Macromolecules 1993,

26, 581–587.(26) Huan, K.; Bes, L.; Haddleton, D. M.; Khoshdel, E. J. Polym. Sci.

2001, 39, 1833–1842.(27) Baez, J. E.; Marcos-Fernandez, A.; Lebron-Aguilar, R.; Martinez-

Richa, A. Polymer 2006, 47, 8420–8429.(28) Jiang, X.; Lok, M. C.; Hennink, W. E. Bioconjugate Chem. 2007, 18,

2077–2084.

BM800686H

pHEMA/PCL Hydrogels Biomacromolecules, Vol. 9, No. 12, 2008 3377