Enzymatically cross-linked gelatin-phenol hydrogels with a broader stiffness range for osteogenic differentiation of human mesenchymal stem cells

Acta Biomaterialia 8 (2012) 1826–1837

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Enzymatically cross-linked gelatin-phenol hydrogels with a broader stiffness range for osteogenic differentiation of human mesenchymal stem cells Li-Shan Wang, Chan Du, Joo Eun Chung, Motoichi Kurisawa ⇑ Institute of Bioengineering and Nanotechnology, 31 Biopolis Way, The Nanos, Singapore 138669, Singapore

a r t i c l e

i n f o

Article history: Received 2 September 2011 Received in revised form 23 January 2012 Accepted 1 February 2012 Available online 8 February 2012 Keywords: Hydrogel Human mesenchymal stem cells Osteogenesis Differentiation Stiffness

a b s t r a c t An injectable hydrogel system, composed of gelatin–hydroxyphenylpropionic acid (Gtn–HPA) conjugates chemically cross-linked by an enzyme-mediated oxidation reaction, has been designed as a biodegradable scaffold for tissue engineering. In light of the role of substrate stiffness on cell differentiation, we herein report a newly improved Gtn hydrogel system with a broader range of stiffness control that uses Gtn–HPA–tyramine (Gtn–HPA–Tyr) conjugates to stimulate the osteogenic differentiation of human mesenchymal stem cells (hMSCs). The Gtn–HPA–Tyr conjugate was successfully synthesized through a further conjugation of Tyr to Gtn–HPA conjugate by means of a general carbodiimide/active estermediated coupling reaction. Proton nuclear magnetic resonance and UV–visible measurements showed a higher total phenol content in the Gtn–HPA–Tyr conjugate than that content in the Gtn–HPA conjugate. The Gtn–HPA–Tyr hydrogels were formed by the oxidative coupling of phenol moieties catalyzed by hydrogen peroxide (H2O2) and horseradish peroxidase (HRP). Rheological studies revealed that a broader range of storage modulus (G0 ) of Gtn–HPA–Tyr hydrogel (600–26,800 Pa) was achieved using different concentrations of H2O2, while the G0 of the predecessor Gtn–HPA hydrogels was limited to the range of 1000 to 13,500 Pa. The hMSCs on Gtn–HPA–Tyr hydrogel with G0 greater than 20,000 showed significantly up-regulated expressions of osteocalcin and runt-related transcription factor 2 (RUNX2) on both the gene and protein level, with the presence of alkaline phosphatase, and the evidence of calcium accumulation. These studies with the Gtn–HPA–Tyr hydrogel with G0 greater than 20,000 collectively suggest the stimulation of the hMSCs into osteogenic differentiation, while these same observations were not found with the Gtn–HPA hydrogel with a G0 of 13,500. Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Hydrogels are widely used as biomaterial formulations for drug delivery or as scaffolds in tissue engineering because their highly hydrophilic characteristics provide an excellent environment for bioactive agents that include therapeutic proteins, growth factors and cells [1–3]. Injectable hydrogel systems are of particular interest for biomedical applications because bioactive agents and cells can be easily encapsulated in hydrogels with a simple injection of a mixture of gel precursors and bioactive agents in solution [4,5]. Thus, no surgical procedures are required for the implantation of hydrogels, or their removal in the case of the biodegradable ones. Hydrogels can be formed in situ either via chemical and/or physical cross-linking reaction mechanisms. It is generally accepted that chemically cross-linked hydrogels are superior to physically cross-linked hydrogels in terms of stability and control in mechanical strength, although the latter has the advantage of

⇑ Corresponding author. Tel.: +65 6824 7139; fax: +65 6478 9083. E-mail address: [email protected] (M. Kurisawa).

being free of cross-linkers. Several strategies have been adopted to prepare chemically cross-linked hydrogels using either natural or synthetic polymers. Chemically cross-linked hydrogels are formed by radical polymerization that adopts redox- or photo-initiators, Michael-type addition reactions, disulfide bond formations, and aldehyde-mediated cross-linking [6–11]. Recently, an enzymatic cross-linking strategy has attracted intensive attention in the preparation of chemically cross-linked hydrogels [12–21]. Hydrogels composed of biopolymer-phenol conjugates were formed using the oxidative coupling of phenol moieties catalyzed by hydrogen peroxide (H2O2) and horseradish peroxidase (HRP). In our previous reports, we have developed an injectable hydrogel scaffold system composed of gelatin–hydroxyphenylpropionic acid (Gtn–HPA) conjugates. This injectable hydrogel scaffold had tunable stiffness for controlling the proliferation and differentiation of human mesenchymal stem cells (hMSCs) in a two-dimensional (2-D) and three-dimensional (3-D) cell culture environment [16,17]. The stiffness of the hydrogels was readily tuned by varying the H2O2 concentration without changing the concentration of its polymer precursor. In 2-D cell culture systems, the hMSCs on a softer hydrogel (storage modulus, G0 = 600 Pa)

1742-7061/$ - see front matter Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2012.02.002

L.-S. Wang et al. / Acta Biomaterialia 8 (2012) 1826–1837

expressed more neurogenic protein markers, while cells on a stiffer hydrogel (G0 = 12,800 Pa) showed a higher up-regulation of myogenic protein. It is suggested that the observation of this cell-lineage specific differentiation of hMSCs in relation to hydrogel stiffness could correlate with previous results that showed the differentiation of hMSCs cultured on collagen-coated polyacrylamide gels – where the cultured hMSC differentiation was directed by the elasticity of the gels that corresponded to the elasticity of the respective tissue [22]. In contrast, in the 3-D cell culture setting, it was found that neurogenic differentiation was enhanced when the cells were cultured in hydrogels with lower stiffness. However, due to the inherent difficulties of 3-D cell culture systems such as poor transportation of nutrients and low degradability of hydrogels as a result of increasing stiffness, the storage modulus (G0 ) of hydrogels studied in three dimensions was limited to no higher than 1000 Pa and the differentiation of hMSCs in three dimensions was only limited to neurogenic differentiation. To address this issue, the design of injectable hydrogels with porous structures would be beneficial for the 3-D culture and differentiation of progenitor cells in hydrogels with high stiffness. It has been reported that hydrogels with porous structures are formed by using a simple stirring process [23]. It is envisioned that the application of such an injectable hydrogel system can be further extended to stimulate cell differentiation to other lineages from progenitor cells, particularly when high stiffness is often required for such differentiation. In our previous study, intensive efforts have been made to maximize the stiffness of the Gtn–HPA hydrogel by optimizing the concentrations of Gtn–HPA and H2O2; however, the stiffness achieved using the Gtn–HPA hydrogel system was still not high enough to induce osteogenic differentiation of hMSCs [17]. Thus, this report has focused on broadening the range of stiffness in the Gtn–phenol hydrogen system further for the stimulation of hMSCs to osteogenic differentiation in a 2-D context. However, as shown in our earlier studies, the amine groups in Gtn were already highly conjugated with HPA [16]. In a similar manner, tyramine (Tyr) was employed additionally in this study to modify the Gtn by a carbodiimide/active ester-mediated coupling reaction, to increase the phenol content in this Gtn–phenol conjugate further in order to achieve a greater degree of stiffness control. The total phenol content was carefully monitored so as not to compromise the solubility of the resultant Gtn–HPA–Tyr conjugate in water. This study defined the range of stiffness of Gtn–HPA–Tyr hydrogel system and its specific range for the stimulation of osteogenic differentiation of hMSCs. We believe that a Gtn–HPA–Tyr hydrogel system that offers a broader range of stiffness extends the application of the enzymatically cross-linked hydrogel system in tissue engineering and regenerative medicine, because substrate stiffness is becoming an essential design variable for tissue engineering [24,25].

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assay kit were obtained from Millipore (USA). Human fibroblasts (HFF-1) were acquired from ATCC (USA). Human mesenchymal stem cells (hMSCs) were provided by Cambrex Bio Science Walkersville, Inc. (USA). Mesencult human basal medium supplemented with mesencult human supplement was purchased from Stem Cell Technologies (Canada). Dulbecco’s modified Eagle medium (DMEM), fetal bovine serum (FBS), penicillin-streptomycin, fluorophore-conjugated secondary antibody and TaqmanÒ gene expression assay kit were provided by Invitrogen (Singapore). Phosphate buffered saline without Ca2+ and Mg2+ (PBS, 150 mM, pH 7.3) solution was supplied by the media preparation facility in Biopolis (Singapore).

2.2. Synthesis of Gtn–HPA–Tyr conjugate The synthesis of the Gtn–HPA–Tyr conjugate was achieved in a two-step reaction process (Fig. 1). First, the Gtn–HPA conjugate was synthesized as described previously [16]. After purification of Gtn–HPA conjugate by dialysis, instead of lyophilizing to obtain Gtn–HPA conjugates, Tyr.HCl (0.50 g, 2.87 mmol), NHS (0.12 g, 1 mmol) and EDC (0.14 g, 0.75 mmol) were added to the purified Gtn–HPA conjugate solution to synthesize Gtn–HPA–Tyr conjugate. The solution was stirred again overnight at room temperature at pH of 4.7. Then, the solution was dialyzed in the same manner as previously described [16]. The purified solution was finally lyophilized to obtain the Gtn–HPA–Tyr conjugate. However, precipitation occurred when Tyr.HCl (1.0 g, 5.74 mmol), NHS (0.24 g, 2.0 mmol) and EDC (0.28 g, 1.5 mmol) were added into the purified Gtn–HPA conjugate solution under the same reaction conditions. This insoluble conjugate, abbreviated as Gtn-insoluble, was harvested and lyophilized for characterization.

2.3. Characterization of Gtn–HPA–Tyr conjugates Proton nuclear magnetic resonance (1H NMR) spectra were recorded on a Bruker AV-400 (400 MHz) spectrometer at room temperature to characterize the conjugation of phenol compounds of Gtn–HPA and Gtn–HPA–Tyr conjugates (10 mg ml1 in D2O). The absorbance of Gtn–HPA–Tyr and Gtn–HPA conjugates (1 mg ml1) was measured at 276 nm using a UV–visible spectrophotometer (U-2810, Hitachi, Japan) to determine the phenol content of Gtn– phenol conjugates. The phenol content of each sample was estimated by comparison with the HPA standards. No significant difference in the absorbance between the HPA and Tyr standards was observed. Differential scanning calorimetry (DSC) was performed with a DSC-Q100 (TA Instruments, USA). The Gtn, Gtn– HPA, Gtn–HPA–Tyr conjugates and Gtn-insoluble (6–8 mg) were heated from 35 to 230 °C at 3 °C min1 in crimped standard aluminum pans to obtain the DSC thermograms. The Gtn-insoluble was also characterized by DSC for comparison.

2. Materials and methods 2.4. Rheological measurement 2.1. Materials Gelatin (Gtn) (MW = 80–140 kDa, pI = 5) and horseradish peroxidase (HRP) (100 units mg1) were obtained from Wako Pure Chemical Industries (Japan). 3,4-hydroxyphenylpropionic acid (HPA), tyramine hydrochloride (Tyr.HCl), N-hydroxysuccinimide (NHS), 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDCHCl), type I collagenase (246 units mg1) Triton X100 and hydrogen peroxide (H2O2, 30%) were purchased from Sigma-Aldrich (Singapore). Anti-osteocalcin (anti-OC) was purchased from R&D System (USA). Anti-runt-related transcription factor 2 (anti-RUNX2), alkaline phosphatase detection kit and osteogenesis

Rheological measurements of the hydrogel formation were performed as described previously [16]. Briefly, the measurements were taken at 37 °C in the dynamic oscillatory mode with a constant deformation of 1% and frequency of 1 Hz. The solutions of HRP and H2O2 with different concentrations were added sequentially to an aqueous solution of Gtn–HPA–Tyr (10 wt.%, 250 ll in PBS). The solution was vortexed and then immediately applied to the bottom plate. The measurement parameters were determined to be within the linear viscoelastic region in preliminary experiments. Rheological measurement was allowed to proceed until the storage modulus (G0 ) reached a plateau.

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1st step OH

OH EDC/NHS

NH CO

O COO N

COOH

Gly X Y n

O

3-(4-Hydroxyphenyl)propionic acid(HPA)

Gelatin

Gelatin-HPAconjugate 2nd step

CO NH

OH

NH CO

NH CO

EDC/NHS

+ NH2 Tyramine

Gelatin-HPA-Tyrconjugate

Gelatin-HPAconjugate X, Y COOH NH proline

COOH H2N CH2 CH2 CH2 CH2 CH COOH NH NH2

HO hydroxyproline

lysine

Fig. 1. Reaction scheme for the synthesis of Gtn–HPA–Tyr conjugates.

2.5. Time course assay on cell attachment This experiment was designed to assess whether the further conjugation of Tyr to Gtn–HPA and the resultant increase of its hydrogel stiffness interfere with the attachment of hMSCs. A filter sterilization process was performed on all the solutions including H2O2, HRP and Gtn–phenol conjugates using a Minisart syringe filter (0.2 lm, Satorius Stedim Biotech GmbH, Germany), prior to hydrogel preparation for both in vitro and in vivo studies. For the hydrogel preparation, 6 ll of HRP was added to 1 ml of the Gtn– HPA and Gtn–HPA–Tyr conjugate solutions (10 wt.%) to obtain a final concentration of 0.15 units ml1. Cross-linking was initiated by adding 6 ll of H2O2 solution to give a final concentration of 13 mM. The mixture was vortexed vigorously before it was transferred to a six-well plate. The hydrogels were allowed to set for 4 h before they were briefly washed with PBS. 250 ll of hMSCs in mesencult human basal medium supplemented with mesencult human supplement (passage number <6) at a cell density of 3  105 was seeded onto these hydrogels. The plates were returned to the incubator at 37 °C under a humidified atmosphere of 5% CO2 for an appropriate period of time ranging from 1 to 6 h. At selected time intervals, the media with unattached cells were aspirated and the wells were washed with PBS. A cell culture well plate without the hydrogel served as a comparison. The cells attached to the hydrogels were harvested by incubating the hydrogels with a collagenase solution (6.7 units ml1) to digest the hydrogels, whereas the cells attached to the culture plate were harvested by trypsin-

ization. The DNA was quantified to determine the number of attached cells at specific times on the hydrogels or culture well plate. The cell pellets were lysed with a freeze–thaw cycle in 200 ll of DNA-free lysis buffer. Samples were then incubated with 200 ll of PicoGreen working solution. The number of cells attached to the surface was then determined by using the fluorescence measurement of the sample solution along with the known concentration of cell suspension for the standard curve. The fluorescence measurement was performed in four replicates by a microplate reader with excitations and emissions at 480 and 520 nm, respectively. 2.6. Immunocytochemistry study In the study of cell focal adhesion, hMSCs were seeded onto the hydrogels and maintained for 1 week before being immunostained using an actin/focal adhesion stain kit. For the immunofluorescence staining of osteocalcin (OC) and runt-related transcription factor 2 (RUNX2), hMSCs were cultured onto surfaces of the hydrogels for 3 weeks before being immunostained. Prior to the immunostaining, the hydrogels together with the encapsulated cells were fixed with 4% formaldehyde solution at room temperature for 20 min. After washing, the cells were permeabilized using 0.5% Triton X-100 in PBS solution at room temperature for 5 min. The cells were then blocked in 0.05% Triton X-100 containing 1% bovine serum albumin at room temperature for 1 h. The samples were then incubated with the respective antibody in blocking

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buffer solution at 4 °C overnight. The cells were washed and then incubated with a FITC-conjugated secondary antibody in the dark for 30 min. In the study of cell focal adhesion, TRITC-conjugated phalloidin was incubated simultaneously with the secondary antibody. Confocal images were acquired using confocal laser scanning microscopy (Olympus FV300, Japan). 2.7. Real-time polymerase chain reaction (PCR) analysis The hydrogels were digested with collagenase solution (0.5 wt.%) and the cells were then harvested for RNA extraction to measure the relative expression of genes of interest in hMSCs that were cultured on Gtn–HPA and Gtn–HPA–Tyr hydrogels for 3 weeks. This extraction was performed according to the protocols specified in the RNeasy Mini Kit (Qiagen, USA). The RNA samples were reverse-transcribed to cDNA using a First strand cDNA synthesis kit (Fermentas, Canada). The relative expression of enolase 2 (ENO2), myogenic differentiation factor 1 (MYOD1), RUNX2 and OC was then determined via real-time PCR using a Bio-Rad iQ5 multicolor real-time PCR detection system (Bio-Rad, USA). Specific primers for these genes were inventoried by Invitrogen. Each PCR reaction was performed in 20 ll of a reaction mixture containing 2 ll cDNA, 1 ll of each primer, 10 ll of TaqmanÒ gene expression master mix (Invitrogen) and 7 ll of diethylpyrocarbonate (DEPC)-treated water (Invitrogen). The final concentration of the primers used in this study was 100 lM. The samples were then subjected to cycling conditions as specified in the TaqManÒ Gene expression assay kit protocol. The experiment was performed in triplicate. The results were normalized to b-actin gene expression and were expressed as fold change values relative to undifferentiated hMSCs. 2.8. Alkaline phosphatase detection The alkaline phosphatase detection was performed using the alkaline phosphatase detection kit, adhering to the protocol provided by the manufacturer. Briefly, the cells cultured on the surfaces of the hydrogels were fixed with 4% paraformaldehyde for 2 min, and followed by incubation with a stain solution composed of Fast Red violet and Naphthol AS-BI phosphate solutions. The samples were then washed and photographed. 2.9. Calcium deposit Osteogenesis quantitation kit (Millipore) was employed to study the mineralized matrix of calcium and phosphorous as a result of the osteogenenic differentiation of hMSCs stimulated by the hydrogel stiffness in accordance with the protocol provided by the manufacturer. In brief, the cells were fixed with 4% paraformaldehyde for 15 min and followed by three times washing with 10 min intervals. They were then incubated with Alizarin Red Stain solution for 20 min. Image acquisition was performed after washing four times with deionized water. Quantitative analysis of the Alizarin Red Staining was subsequently performed through incubation with 10% acetic acid for 30 min. The cells were scraped off, transferred, heated at 85 °C for 10 min and centrifuged at 20,000g for 15 min. The supernatant was then neutralized to a pH of 4.2 and its absorbance at 405 nm was recorded using a microplate reader (Infinite M200, Tecan, Switzerland). The concentration of Alizarin Red was calculated against a standard curve obtained by a set of known concentrations of Alizarin Red solution and normalized to the number of cells on the hydrogels. The number of cells was determined by a separate experiment using a Picogreen assay as described above. Detection of such a mineralized matrix was also performed using a field emission scanning electronic microscope (FESEM, JEOL JSM-7400F) equipped with energy dispersive X-ray

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spectrometry (EDX). The hydrogels with the cells on the surfaces were washed three times with deionized water after fixing with 4% paraformaldehyde solution. They were then freeze-dried before being coated with a thin layer of platinum (Pt) using a JEOL auto fine coater (JFC-1600). 2.10. Enzymatic degradation of Gtn–HPA–Tyr hydrogels Slab-shaped Gtn–HPA–Tyr hydrogels of different stiffness were prepared as described previously [16] with variation in the concentration of the H2O2 solutions. The G0 of resultant hydrogels were 600, 3200, 13,500, 14,800 and 26,800 Pa, when the H2O2 solution was given with final concentrations of 1.7, 3.4, 8.5, 17 and 13 mM, respectively. The hydrogel disks (1 mm  16 mm) were immersed in 20 ml of PBS containing 6.7 units ml1 of type I collagenase and incubated at 37 °C and 100 rpm in an orbital shaker. The degree of degradation of the hydrogels was estimated by measuring the residual hydrogel weight. The hydrogels were removed from the solution, blotted dry and weighed at specific time points to measure their residual weight. 2.11. Cytotoxicity study of fragmented Gtn–phenol hydrogels Degraded products of Gtn–phenol hydrogels by collagenase were tested for their cytotoxicity. The Gtn–phenol hydrogels (10 wt.%) were immersed in PBS containing 6.7 units ml1 of type I collagenase and incubated at 37 °C in an orbital shaker at 100 rpm until they were fully degraded. The solution, containing the degraded hydrogels, was aseptically diluted using the culture medium to give final concentrations of 0.25 and 0.6 wt.%. Human fibroblasts (HFF-1) in 200 ll of the respective medium containing the degraded products were added at a density of 5  104 cells ml1 to the wells of 96-well plate. The wells containing the cells exposed to type I collagenase solutions of the same concentration served as comparisons. The cells were incubated for 24 h. Then, cell viability was assessed using the Alamar blue assay. Briefly, the culture medium was aspirated and replaced with fresh medium, and incubated with Alamar blue dye (10%) for 3 h to measure the viability of cells. A fluorescence measurement was performed with an Infinite M200 (Tecan, Switzerland), where excitation and emission wavelengths were set at 570 and 590 nm, respectively. The results were expressed as a percentage of viability compared with untreated cells. 2.12. In vivo degradation of Gtn–phenol hydrogels and histological evaluation Non-obese diabetic/severe combined immunodeficiency (NOD/ SCID) mice that were supplied by Biological Resource Center (BRC) in Biopolis, Singapore were used at 6–8 weeks of age. Gtn–phenol conjugates were dissolved in PBS at a concentration of 10 wt.%. Prior to injection, 3 ll of HRP and 3 ll of H2O2 were added to 500 ll of Gtn–phenol solutions to give final concentrations of 0.15 units ml1 and 13 mM, respectively. 100 ll of the solution was then injected subcutaneously through a 22-gauge needle into the mice. Following this subcutaneous injection, the in vivo degradability of hydrogel was assessed by measuring weight loss at specific time intervals with reference to initial weight of the implanted hydrogel. The hydrogel implants were collected at 1 and 4 weeks post-injection for a separate histological evaluation study. The samples were fixed with 4% paraformaldehyde for 24 h at 4 °C and immersed in 30% sucrose solution overnight, before they were imbedded in an OCT cryostat embedding medium (Tissue-TekÒ, Sakura Finetek, USA). The cryostat sections were then cut, collected on silane-coated slides, and stored at 20 °C for analysis. The care and use of laboratory animals was performed according to the

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approved protocol of the Institutional Animal Care and Use Committee (IACUC) at the BRC in Biopolis, Singapore.

0 -0.02

72.2 oC

-0.04

72.5 oC

2.13. Statistical analysis All data are expressed as the mean ± standard deviation. Differences between the values were assessed using Student’s unpaired t-test and p < 0.05 was considered statistically significant.

Gtn 200.6 oC

Gtn-HPA -0.06

72.3 oC

197.4 oC Gtn-HPA-Tyr

-0.08

3. Results and discussion

72.4 oC 194.8 oC

Gtn-HPA-Insoluble -0.10

3.1. Synthesis and characterization of Gtn–HPA–Tyr conjugate We have previously reported that Gtn–HPA conjugate was successfully synthesized by a general carbodiimide/active ester-mediated coupling reaction in distilled water [16]. It was found that 90% of the amine group in Gtn was conjugated with HPA. In this study, Gtn–HPA–Tyr conjugate was synthesized to further increase the phenol content into Gtn by using a two-step reaction (Fig. 1). Firstly, the Gtn–HPA conjugate was synthesized by a reaction between amine groups of Gtn and succinimide-activated HPA. After the purification of the Gtn–HPA conjugate through dialysis, Tyr was added to the Gtn–HPA conjugate to synthesize Gtn–HPA–Tyr conjugate. As described in Section 2, the conjugation of the phenol was optimized so as not to compromise the water-solubility of the resultant conjugate. 1 H NMR measurements of synthesized bioconjugates revealed that the integrated phenol intensities (6.8 ppm and 7.1 ppm) of the Gtn–HPA–Tyr conjugate were higher than those of the Gtn– HPA conjugate, confirming the conjugation of both HPA and Tyr with Gtn (Fig. 2). Furthermore, the conjugation of phenol molecules was quantitatively analyzed by measuring the absorbance values at 276 nm. The total phenol content of Gtn–HPA and Gtn– HPA–Tyr conjugate was determined to be 4.44  107 and 7.11  107 mol mg1 conjugate, respectively. The DSC curves of Gtn, Gtn–HPA, Gtn–HPA–Tyr and Gtn-insoluble conjugates were compared in Fig. 3. Two glass transition temperatures (Tgs) (first Tg; 72.2 and second Tg; 200.6 °C) were observed in the thermograms for Gtn. The conjugation of HPA to Gtn seemed to have a strong influence on the second Tg. With such conjugation, the second Tg was shifted by 3.2 °C from 200.6 °C. The further conjugation of Tyr to Gtn–HPA conjugate decreased the second Tg even further by another 2.6 °C to 194.8 °C in the case of the Gtn–HPA–Tyr conjugate. This shift of the second Tg was largely pronounced in the Gtn-insoluble conjugate with overly conju-

-0.12 Exo up 50

157.9 oC

70

90

110

130

170

190

210

o

Temperature ( C) Fig. 3. DSC thermograms of Gtn, Gtn–HPA, Gtn–HPA–Tyr and Gtn-insoluble conjugates.

gated Tyr moiety. Its second Tg appeared at 157.9 °C. However, the first Tg was only marginally different among all the samples. As Fraga et al. have reported, the first Tg of Gtn is a minor one, observed at around 80–100 °C and associated with the glass transition of a-amino acid blocks (soft blocks) [26]. The other more intense Tg is observed at around 180–200 °C, and represents the blocks of imino acids, proline, hydroxyproline with glycine (rigid blocks). Therefore, the shift of the second Tg by phenol conjugation suggests that the majority of the conjugation occurred near the rigid block of the Gtn with minimal impact on the soft block of Gtn. It is considered that the conjugation of the phenol molecules disrupted the crystalline structure, mainly formed by the rigid blocks, and drove the second Tg down. 3.2. Rheological measurement As previously reported, the hydrogel composed of Gtn–HPA conjugate was formed using the oxidative coupling of HPA moieties catalyzed by H2O2 and HRP [16,17]. The storage modulus (G0 ) of the Gtn–HPA hydrogels prepared using 5 wt.% of Gtn–HPA conjugate was readily tuned from 600 to 8000 Pa by varying the H2O2 concentration with G0 peaking at 8.5 mM of H2O2 [17]. In this study, the concentration of both Gtn–HPA and Gtn–HPA–Tyr conjugates for the hydrogel preparation was 10 wt.%. A further

Gtn

Gtn-HPA

Gtn-HPA-Tyr 7.5

150

7.0 Fig. 2. 1H NMR spectra of Gtn, Gtn–HPA and Gtn–HPA–Tyr conjugates.

6.5 ppm

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increase in polymer concentrations in both polymer conjugate systems significantly increased the viscosity of their solutions and resulted in difficulties in rheological measurement and administration, and clinically relevant procedures such as a subcutaneous injection of Gtn–phenol solutions through syringes. Therefore, the concentration of Gtn–HPA or Gtn–HPA–Tyr conjugates was capped at 10 wt.%. The Gtn–HPA–Tyr hydrogel was formed using the same oxidative coupling of both HPA and Tyr moieties catalyzed by H2O2 and HRP (Fig. 4). It is well known that phenols cross-link through either a more common C–C linkage between the ortho-carbons of the aromatic ring or a C–O linkage between the ortho-carbon and the phenolic oxygen [27]. Fig. 5 summarizes the rheological properties of Gtn–phenol hydrogels formed with varied H2O2 concentrations. The same trend where the G0 was significantly increased with the increase of the H2O2 concentration was found in both Gtn–HPA and Gtn–HPA– Tyr hydrogels (Fig. 5a). The further increase of H2O2 concentrations to 17 mM resulted in a decline of G0 , which was likely due to deactivation of the HRP by an excess amount of H2O2 [28]. The dependence of H2O2 concentration indicates that with more phenol moiety available in the system, more H2O2 was needed to maximize the cross-linking density. H2O2 decomposes to water after oxidizing HRP, which in turn oxidizes the HPA. Thus, the percentage of phenol moieties that actually participated in the cross-linking reaction would depend on the amount of H2O2 available. The G0 of Gtn–HPA hydrogel ranged from 998 ± 14 to 13,556 ± 665 Pa. Previously, we reported on the rheological properties of Gtn–HPA hydrogels prepared using 5 wt.% of Gtn–HPA conjugates and found that the highest G0 was 8000 Pa [17]. It was also found that in contrast to the effect of H2O2 on hydrogel stiffness, an increase in HRP concentration did not result in an increase in hydrogel stiffness. However, the gelation rate was in direct correlation with the HRP concentration. The HRP concentration was optimized at 0.15 unit ml1 as the gel point was less than 160 s, which is very efficient for gel formation as an injectable system. In this study, we found that the G0 of Gtn–HPA hydrogel increased to 13,500 Pa when 10 wt.% of Gtn–HPA conjugate was utilized. The increase in G0 achieved with a higher concentration of Gtn–HPA conjugate (10 wt.%) indicates that a higher number of HPA moieties participated in the cross-linking reaction. For the Gtn–HPA–Tyr hydrogel, a much higher G0 (26,830 ± 471 Pa) was achieved when 13 mM of H2O2 was utilized. The highest G0 achieved in a Gtn–HPA–Tyr hydrogel was almost double compared with the Gtn–HPA hydrogel. It indicates that the further conjugation of Tyr to the Gtn–HPA conjugate significantly increased the stiffness of the resultant hydrogel. In addition, the time

CO NH

(a) 30000 Gtn-HPA-Tyr

25000 20000 15000

Gtn-HPA

10000 5000 0 0

5

10 H2O2 (mM)

15

20

(b) 10000 Gtn-HPA 8000 6000 4000 2000 Gtn-HPA-Tyr 0 0

5

10 H2O2 (mM)

15

20

Fig. 5. Effects of H2O2 on (a) G0 and (b) the time needed for G0 to reach a plateau. HRP concentration is fixed at 0.15 units ml–1. Results are shown as the mean values ± standard deviation (n = 3).

required for G0 to reach a plateau increased with an increase of H2O2 concentration for both the Gtn–HPA and Gtn–HPA–Tyr hydrogels (Fig. 5b). This result is in good agreement with our earlier reports of enzyme-mediated injectable hydrogel systems [14,16]. Interestingly, the time required for G0 to reach a plateau in the Gtn–HPA–Tyr hydrogel was lower compared to that of Gtn–HPA hydrogel when higher H2O2 concentration (13 and 17 mM) was utilized. This faster cross-linking achieved in the Gtn–HPA–Tyr hydrogel is most likely attributed to an increase in

NH CO Cell HRP/H2O2 Gtn-HPA-Tyrhydrogel

Gtn-HPA-Tyrconjugate Fig. 4. Formation of Gtn–HPA–Tyr hydrogel by enzyme-catalyzed oxidation for 2-D cell growth.

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the concentration of local phenol moiety as a result of the higher phenol content in the Gtn–HPA–Tyr hydrogel. It is highly important in tissue engineering to design a hydrogel system which can provide a broader range of stiffness control, especially when the role of substrate stiffness is considered on cell functions such as cell morphology, adhesion, migration, proliferation, apoptosis and differentiation [22,29–34]. In our previous studies using the Gtn–HPA hydrogel system, the hMSCs did clearly demonstrate their capability to differentiate into specific lineages based on the stiffness of the hydrogel. The hMSCs on a softer hydrogel (600 Pa) expressed more neurogenic protein markers, while cells on a stiffer hydrogel (12,800 Pa) showed a higher upregulation of myogenic protein [17]. However, osteogenic differentiation was not observed using the Gtn–HPA hydrogel system. This was most likely due to insufficient stiffness as implied in the previous report, for a substrate with much higher stiffness was required in order to stimulate osteogenic differentiation in contrast to the stiffness needed for myogenic differentiation [22]. Our improved Gtn–phenol hydrogel system offers G0 up to 26,800 Pa, and in order to define the range of stiffness in G0 needed for stimulation of osteogenic differentiation, hydrogels with varied stiffness ranging from 13,500 Pa to 26,800 Pa were prepared for this study. Their rheological properties were summarized in Table 1. The G0 of Gtn–HPA–Tyr hydrogels was significantly greater than that of the Gtn–HPA hydrogel. The gel point is defined as the time at which crossover of G0 and loss modulus (G00 ) occurred, and is used as an indicator of the gelation rate. From the measurements of the gel point of hydrogels, the gel point was less than 160 s, indicating an efficient gel formation. From the perspective of an injectable hydrogel system, the rapid gel formation of the Gtn–HPA–Tyr hydrogel may minimize uncontrolled diffusion of the gel precursors and bioactive agents to the surrounding tissues in potential clinical applications. 3.3. Cell attachment It is generally understood that cells growing on stiffer surfaces have a larger spreading area, more organized cytoskeletons and a more stable focal adhesion, although there is cell-to-cell variability. In our previous study on the Gtn–HPA hydrogel system, the same correlation was found between the hMSCs focal contact and the hydrogel stiffness across the range of stiffness studied (600 to 12,800 Pa) [17]. In this study, Gtn–HPA hydrogel (G0 = 13,500 Pa) and Gtn–HPA–Tyr-26k hydrogel (G0 = 26,800 Pa) were chosen to investigate the effect of further increase in hydrogel stiffness on the cell attachment and spreading as a result of the further conjugation of the Tyr moieties. The number of adherent hMSCs increased over incubation time (Fig. 6a). The cell attachment on both Gtn–HPA and Gtn–HPA–Tyr-26k hydrogels in the first 2 h was significantly higher than on the plastic cell culture plate, while no significant difference between Gtn–HPA and Gtn– HPA–Tyr-26k hydrogels was observed. After 4 h, more than 95% of cells attached to all the surfaces including the culture plate. Within the range of hydrogel stiffness studied in this report, the effect of the further increase in hydrogel stiffness on cell attach-

Fig. 6. (a) hMSCs attachment and (b) its confocal fluorescence microscopy of focal adhesion and actin cytoskeleton on the surface of Gtn–phenol hydrogels. Results are shown as the mean values ± standard deviation (n = 4). ⁄p < 0.001.

ment and its focal contact was less pronounced when compared with the findings in a previous study using hydrogels with stiffness in lower range from 600 Pa to 8000 Pa. Fig. 6b shows the confocal fluorescence images of focal adhesion and actin cytoskeleton in hMSCs when the cells were cultured using Gtn–phenol hydrogels. These images reveal focal contacts in green using an anti-vinculin monoclonal antibody. Also, F-actin was detected in red. No significant difference was observed in cell focal contact between the surfaces of Gtn–HPA and Gtn–HPA–Tyr-26k hydrogels when the stiffness of the hydrogel ranged from 13,500 to 26,800 Pa. The cells were found to adhere tightly onto both hydrogel surfaces with organized structural arrangements of F-actin and focal contact, indicating that the further conjugation of the Tyr moiety to the Gtn–HPA conjugate did not affect cell attachment behavior. The cell adhesion was strong and stable enough to induce steady cell growth. 3.4. Osteogenic differentiation of hMSCs The osteogenic differentiation of hMSCs in relation to the stiffness of hydrogel was studied after the cells were cultured on the surfaces of Gtn–phenol hydrogels of varied stiffness. Real-time

Table 1 Rheological property of Gtn–HPA and Gtn–HPA–Tyr hydrogels used in the cell culture study.a

a b c

Hydrogel

Gtn–HPA conjugate (wt.%)

Gtn–HPA–Tyr conjugate (wt.%)

HRP (units ml1)

H2O2 (mM)

G0 (Pa)

Gel point (s)b

Time needed for G0 to reach plateau (s)

Gtn–HPA Gtn–HPA–Tyr-20k Gtn–HPA–Tyr-26k

10 0 0

0 10 10

0.15 0.15 0.15

13 11 13

13557 ± 665c 20078 ± 576c 26830 ± 47c

74 ± 14 142 ± 24 151 ± 15

6412 ± 235 1345 ± 45 1590 ± 29

Measurement was taken with constant deformation of 1% at 1 Hz and 37 °C (n = 3). Results are shown as the values ± standard deviation. Gel point is defined as the time at which the crossover of storage modulus (G0 ) and loss modulus (G00 ) occurred. It is used herein as an indicator of the rate of gelation. Significantly different from each other (p < 0.001).

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Fig. 7. (a) Gene expression, (b) immunofluorescence images of OC and RUNX2 and (c) alkaline phosphatase staining of hMSCs after 3 weeks of culture on Gtn–phenol hydrogels of varied stiffness. For the gene expression experiment, undifferentiated hMSCs were used as a reference sample and all results were normalized with respect to the expression of b-actin levels. Results are shown as the mean values ± standard deviation (n = 3). ⁄p < 0.01, ⁄⁄p < 0.001.

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PCR analysis, immunocytochemistry, colorimetric assay on alkaline phosphatase (ALP) and calcium accumulation were employed to determine such differentiation in this study. Although not much difference was observed on the cell attachment and focal adhesion when they were cultured on the Gtn–phenol hydrogels with G0 ranging from 13,500 to 26,000 Pa, we observed a significant difference of relative gene expressions from hMSCs cultured on hydrogels of different stiffness (Fig. 7a). Gtn– HPA–Tyr-20k (G0 = 20,078 Pa) with intermediate stiffness between 13,500 and 26,000 Pa was also included in this study to determine threshold stiffness for osteogenic differentiation. A significant increase in expressions of runt-related transcription factor 2 (RUNX2) and osteocalcin (OC), two commonly studied osteogenic markers, was detected on the cells cultured on the hydrogels with G0 higher than 20,000 Pa in real-time PCR analysis in comparison to those cultured on the Gtn–HPA hydrogel (G0 = 13,000 Pa). The level of up-regulation of such gene expresssions was directly correlated to the hydrogel stiffness. The Gtn–HPA–Tyr-26k showed a ten-fold and seven-fold increase in the OC and RUNX2 gene expressions, respectively, while Gtn–HPA–Tyr-20k showed a three-fold increase in both genes. It suggests that the hydrogel with a stiffness higher

than 20,000 Pa was more likely to stimulate the osteogenic differentiation, although some up-regulation of myogenic transcription factor 1 (MYOD1), a well-studied myogenic differentiation factor, was also found in cells cultured on Gtn–HPA–20k. For the neural transcription factor, Gtn–HPA–Tyr-26k showed a significant down-regulation of enolase2 (ENO2). The two osteogenic differentiation markers, OC and RUNX2, were also shown on the protein level. Images on immunostaining (Fig. 7b) revealed a positive staining for these two markers. The color intensity was higher on the cells cultured on stiffer Gtn– HPA–Tyr-26k when the setting for confocal imaging remained unchanged for all the images shown in Fig. 7b. Alkaline phosphatase (ALP), another well-known osteognic marker, was also detected in red on cells cultured on stiffer hydrogels by colorimetric assay (Fig. 7c). The intensity of the color was also visually higher when the cells were cultured on the hydrogels with higher stiffness. Besides the expressions of osteogenic markers on both gene and protein levels, a mineralized matrix rich in calcium would also be considered as one of the characteristics of osteogenic differentiation. Thus, both qualitative and quantitative measurements of calcium deposits were performed using Alizarin Red assay. The cells

(a)

0.5

(b)

** **

**

0.4

0.3

0.2

0.1

0 Gtn-HPA

Gtn-HPA-Tyr20k

Gtn-HPA-Tyr26k

(c)

0

1

2 3 Energy (keV)

4

5

Fig. 8. (a) Alizarin Red staining of hMSCs after 3 weeks of culture on Gtn–HPA and Gtn–HPA–Tyr hydrogels. (b) Quantitative measurement of Alizarin Red. The concentration is normalized to cell number. Results are shown as the mean values ± standard deviation (n = 4). ⁄⁄p < 0.001. (c) EDX and (d) SEM analysis on the surfaces of Gtn–phenol hydrogels.

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cultured on the surfaces of Gtn–HPA–Tyr hydrogels with G0 greater than 20,000 Pa showed characteristic staining for calcium with Alizarin Red, while those on Gtn–HPA hydrogel did not (Fig. 8a). The concentration of Alizarin Red extracted from the cells cultured on Gtn–HPA–Tyr hydrogels was much higher that from cells cultured on Gtn–HPA hydrogel (Fig. 8b). The calcium accumulation by the cells cultured on the Gtn–HPA–Tyr hydrogel with higher stiffness was further confirmed by EDX-SEM analysis (Fig. 8c and d). A higher level of Ca on Gtn–HPA–Tyr hydrogels was detected, while it was not detected on the Gtn–HPA hydrogel. The collective results indicated a successful osteogenic differentiation through the hydrogel stiffness when the cells were cultured on the Gtn–HPA–Tyr hydrogel of G0 higher than 20,000 Pa, as evidenced by the presence of the expressed osteogenic markers, ALP and calcium. The above observation has validated the role of substrate stiffness in directing hMSC differentiation in the absence of other biochemical factors. The new Gtn–phenol (Gtn–HPA–Tyr) hydrogel system has allowed us to make observations on osteogenic differentiation of hMSCs. Such differentiation was not achieved with the predecessor Gtn–HPA hydrogel system of the same conjugate concentration. This was most likely due to the lack of the required stiffness for such differentiation. Our newly improved Gtn–HPA– Tyr hydrogel could serve as an appropriate platform to achieve the full benefits of stem cell differentiation for tissue regeneration, in the light of current research that directs stem cell differentiation solely by substrate stiffness. 3.5. Enzymatic degradation of Gtn–HPA–Tyr hydrogel and cytotoxicity of degraded products Enzymatic degradability of Gtn–HPA–Tyr hydrogels was examined in the presence of type-I collagenase. Type-1 collagenase is a member of the matrix metalloproteases (MMP) family, which was found to degrade the extracellular matrix, and leads to cell migration and growth in the body [35]. Accordingly, they can digest proteolysis-sensitive hydrogels [9]. It is understood that the enzymatic degradability of biopolymers is lowered by conjugation to the polymers [36]. Also, the enzymatic degradability of hydrogel could be diminished by an increase in hydrogel stiffness. Such a low hydrogel degradability is not ideal for a scaffold in tissue engineering applications. Therefore, it is crucial to assess the degradability of hydrogels especially if the stiffness of the hydrogel was altered. Fig. 9 shows enzymatic

100 80 60 40 20 0 0

5

10

15 Time (h)

20

25

Fig. 9. Enzymatic degradation of Gtn–HPA–Tyr hydrogels with different stiffness; 600 (s), 3200 (h), 13,500 (}), 14,600 (4) and 26,800 Pa (5). The experiment was carried out in the presence of 6.7 units ml1 of type I collagenase at 37 °C. Results are shown as the mean values ± standard deviation (n = 3).

* * 100 Gtn-HPA 80 Gtn-HPA-Tyr 60 40 20 0

collagenase 0.25 0.6 alone ConcentrationofGtn-phenolconjugate(wt.%)

Fig. 10. Cytotoxicity study of fragmented Gtn–phenol hydrogels. The fragmented products were obtained by the enzymatic degradation of hydrogels in the presence of 6.7 units ml1 of type I collagenase at 37 °C. HFF-1 containing the degraded products was incubated for 24 h before the Almar blue assay. Results are shown as the mean values ± standard deviation (n = 8). ⁄p < 0.01.

degradation of Gtn–HPA–Tyr hydrogels with different stiffness. All the Gtn–HPA–Tyr hydrogels could be completely degraded by type-I collagenase although the phenol content and stiffness were higher compared to the Gtn–HPA hydrogel system. It was also found that the hydrogels with higher stiffness degraded slower when compared to the ones with lower stiffness. This result suggests that the enzymatic degradability of Gtn–HPA hydrogels can be controlled by hydrogel stiffness, and the Gtn–HPA–Tyr hydrogels would be useful for various biomedical applications. The fragmented product of the hydrogel was prepared by enzymatic degradation of Gtn–phenol hydrogels in an attempt to simulate the clinically relevant conditions when hydrogel degradation occurs upon implantation in the body. The cytotoxicity of the degraded product from Gtn–phenol hydrogel was evaluated by incubating it with the cells for 24 h at 37 °C. The viability of human fibroblasts (HFF-1) was 95 and 80% when the degraded product of 0.25 and 0.6 wt.% was utilized, respectively (Fig. 10). Type I collagenase at this test concentration did not have an adverse effect on cell viability. 3.6. In vivo degradation of Gtn–phenol hydrogels and histological evaluation It was also observed that the in vivo degradation of the Gtn– phenol hydrogel was hydrogel stiffness-dependent (Fig. 11a). The degradation of Gtn–HPA–Tyr hydrogels was significantly slower than that of the Gtn–HPA hydrogel. In addition, a much slower degradation was observed when hydrogels of the same compositions were injected subcutaneously in mice compared to our previous enzymatic degradation study in vitro largely due to the much lower enzyme concentration present under the skin. The hydrogels formed readily in situ and 70% still remained under the subcutaneous layer after 4 weeks post-injection for the hydrogel with lowest stiffness. The optical microscopy-based histological evaluation of tissues surrounding the implant was performed to evaluate the biocompatibility and toxicity of the Gtn–phenol hydrogel. Representative photomicrographs of hematoxylin and eosin stained hydrogels and surrounding tissue are shown in Fig. 11b. The hydrogel implants with a minimal capsule surrounding the materials was evident both 1 and 4 weeks after implantation. The thickness of the capsule seemed to be in a trend of decreasing over implantation time. No prominent inflammatory response (giant cells, monocytes, and polymorphonuclear leukocytes) or significant histological

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(a) 100

80

Gtn-HPA-Tyr-26k

*

* Gtn-HPA-Tyr-20k Gtn-HPA

60 0

1

2

3

4

Time (week)

(b)

i

ii

Interface

Gtn-HPA

Interface 200 µm

200 µm iii

iv

Gtn-HPA-Tyr-26k

Interface Interface 200 µm

200 µm

Fig. 11. (a) In vivo degradation profile of Gtn–phenol hydrogels in mice. Results are shown as the mean values ± standard deviation (n = 6). ⁄p < 0.05. (b) Representative histological images of the hydrogels and surrounding tissue 1 week (i, iii) and 4 weeks (ii, iv) after subcutaneous injection in vivo.

abnormality was detected at both 1 and 4 weeks post-implantation. The collective results from these experiments demonstrated that the Gtn–phenol hydrogels and their degraded products showed minimal cytotoxicity. These results indicate that the Gtn–HPA–Tyr hydrogels were capable of providing biologically compatible support in potential biomedical applications. 4. Conclusions Gtn–HPA–Tyr conjugate was successfully synthesized by the further conjugation of Tyr to the Gtn–HPA conjugate. The phenol content of the Gtn–HPA–Tyr conjugate was enhanced when compared to that of the Gtn–HPA conjugate. We successfully formulated the Gtn–HPA–Tyr hydrogels with a broader range of stiffness by using an enzyme-mediated oxidation reaction. A significant high level of gene and protein expressions of OC and RUNX2, which are closely associated with osteogenic differentiation of hMSCs, was observed in the cells when they were cultured on the hydrogel with stiffness higher than 20,000 Pa. The level of such expressions was directly correlated to the hydrogel stiffness. The presence of alkaline phosphatase and calcium were also evident in cells grown on the hydrogels with G0 higher than 20,000 Pa. This observation was only made possible in this newly improved Gtn–HPA–Tyr hydrogel system. Such differentiation

was not achieved using the predecessor Gtn–HPA hydrogel system with the same conjugate concentration. Furthermore, the in vivo degradation of Gtn–phenol hydrogels was also stiffness-dependent and the hydrogels also did not show prominent inflammatory response in vivo. Thus, our new design of an injectable hydrogel system with tunable mechanical strength and rapid gelation rate would be beneficial to repair bone defects. In the light of current research on the role of substrate stiffness in many physiological processes, our hydrogel system could serve as an appropriate platform to study the effect of substrate stiffness for tissue regeneration on a broad range of various cell functions. Acknowledgements This work was supported by the Institute of Bioengineering and Nanotechnology (Biomedical Research Council, Agency for Science, Technology and Research, Singapore). We thank Dr Yang Xianfen for his help with the SEM and EDX characterizations.

Appendix A. Figures with essential color discrimination Certain figures in this article, particularly Figs. 4, 6, 7, 8 and 11, are difficult to interpret in black and white. The full color images

L.-S. Wang et al. / Acta Biomaterialia 8 (2012) 1826–1837

can be found in the on-line version, at doi:10.1016/j.actbio.2012. 02.002

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