Article pubs.acs.org/Biomac
Synthesis and Characterizations of In Situ Cross-Linkable Gelatin and 4-Arm-PPO-PEO Hybrid Hydrogels via Enzymatic Reaction for Tissue Regenerative Medicine Kyung Min Park, Yunki Lee, Joo Young Son, Dong Hwan Oh, Jung Seok Lee, and Ki Dong Park* Department of Molecular Science and Technology, Ajou University, 5 Woncheon, Yeongtong, Suwon 443-749, Republic of Korea ABSTRACT: In situ cross-linkable hybrid hydrogels composed of gelatin and 4-arm-polypropylene oxide−polyethylene oxide (Tetronic) was developed as an injectable scaffold for tissue regeneration. The gelatin was modified by hydroxyphenyl propionic acid (HPA) and the Tetronic was conjugated with tyramines (Tet−TA). The hydrogels were rapidly formed by mixing the polymer solutions containing horseradish peroxidase (HRP) and hydrogen peroxide (H2O2). The gelation time and mechanical properties of the hydrogels could be controlled by varying the HRP and H2O2 concentrations. In vitro degradation study of the hybrid hydrogels was carried out using collagenase and the prolonged proteolytic degradation was obtained due to the presence of the Tetronic. Human dermal fibroblast (hDFB) was cultured in the hydrogel matrices to evaluate the cyto-compatibility. The encapsulated cells were shown to be highly viable and spread over the gel matrices, suggesting that the hybrid hydrogels have an excellent cyto-compatibility. The hydrogels were also subcutaneously injected in the back of mice and the results demonstrated that the hydrogels were rapidly formed at the injected site. From these results, we demonstrate that the in situ cross-linkable hydrogels formed by hybridization of gelatin and Tetronic via enzyme-mediated reactions hold great promise for use as injectable matrices for tissue regenerative medicine due to their tunable physico-chemical properties and excellent bioactivity.
1. INTRODUCTION In situ forming hydrogel systems have attracted considerable interest as injectable scaffolds for tissue engineering and drug delivery due to their easy applications and minimally invasive injection procedure.1,2 The encapsulation of the therapeutic agents like drugs, genes, peptides, proteins, and cells can be easily performed and it is possible to acquire the hydrogels that fit in any shapes of the defects and surrounding tissue. Over past decades, various types of in situ forming hydrogels have been studied. These hydrogel systems can be classified into physical cross-linking systems, such as those based on the charge attractions,3−5 hydrophobic interactions,6−8 and stereocomplexation,9−11 and chemical cross-linking systems using Michael−type addition reaction12,13 and UV irradiation.14 However, these systems encounter some drawbacks including a multistep process, relatively long reaction times, and insufficient mechanical properties. An enzyme-mediated cross-linking reaction using horseradish peroxidase (HRP) and hydrogen peroxide (H2O2) provides an alternative method for the in situ formation of the hydrogels. The HRP-mediated cross-linking reaction has several advantages including tunable reaction rate, mild reaction conditions, and good biocompatibility. The enzyme-mediated formable hydrogels are easy to handle and capable of loading with cells or bioactive molecules homogeneously because the viscosity of precursor polymer solutions is relatively low. In recent years, several studies have applied the enzymatic reaction to prepared © 2012 American Chemical Society
in situ forming hydrogels composed of phenol-conjugated synthetic polymer, polysaccharides, and proteins, including dextran,15 hyaluronic acid,8,16 alginate,17 cellulose,18 gelatin,19−21 and albumin,22 as injectable materials. These hydrogels have been shown to enable control of the gelation time and mechanical strength by varying the reaction conditions and to exhibit a great biocompatibility. Gelatin, a natural polymer derived from the partial hydrolysis of collagen, is commonly used for various biomedical applications involving the pharmaceutical and medical applications due to its excellent biocompatibility and the proteolytic degradability.23−25 However, gelatin-based hydrogels have some limitations that the degradation is often too fast and the mechanical strength is not sufficient in vivo.26 Although a number of studies have been reported to improve the mechanical stability,19,27 it still remains a major challenge. One of the most reasonable approaches is to use the hybrid systems that consist of natural components, providing the tissue and cell compatible environments, and synthetic materials, which can improve the mechanical strength and stability. Therefore, we employed gelatin as the natural/degradable component and Tetronic for the synthetic hydrogel backbone in the study. Received: July 5, 2011 Revised: January 19, 2012 Published: January 21, 2012 604
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
Table 1. Sample Preparation and Sample Codes samples
Tet−TA conc. (wt %)
GHPA conc. (wt %)
phenol conc. (μmol/ mL)
HRP conc. (mg/mL)
H2O2 conc. (wt %)
elastic modulus (G′, Pa)
viscous modulus (G″, Pa)
Tet-TA T5G1 T5G3 T5G5 GHPA
5 5 5 5 0
0 1 3 5 5
11.0 12.2 14.6 17.0 6.0
0.0025 0.0025 0.0025 0.0025 0.0025
0.025 0.025 0.025 0.025 0.025
1540 2130 4220 8470 2010
108 102 195 204 6
The Tet−TA was synthesized by the conjugation of TA with Tet− PNC. The Tet−PNC (20 g, 1.11 mmol) dissolved in 200 mL of dimethyl sulfoxide (DMSO) was added to the TA (0.761 g, 5.55 mmol) in 150 mL of DMSO, and the conjugation was carried out at RT under nitrogen atmosphere for 24 h. The reacted solution was dialyzed sequentially against methanol and acetone (molecular weight cut off (MWCO) = 3500 Da) for 3 days and filtered using ammonium oxide. The filtered solution was evaporated and precipitated twice in cold diethyl ether. The precipitate was isolated and dried under vacuum, yielding a white powder. The Tet−TA polymer (10 mg/mL) was dissolved in CDCl3 for the 1H NMR measurement to determine the chemical structure (Varian, 400 MHz spectrometer). 2.3. Synthesis of Gelatin−Hydroxyphenyl Propionic Acid (GHPA) Conjugates. Gelatin−hydroxyphenyl propionic acid (GHPA) was synthesized using EDC and NHS as coupling reagents. Gelatin (5 g/150 mL) was dissolved in water at 40 °C. HPA (1.66 g, 10 mmol) was dissolved in a mixture of water and dimethylformamide (DMF) with the volume ratio of 3:2 and reacted with EDC (1.94 g, 10 mmol) and NHS (1.6 g, 13.9 mmol) at RT for 15 min to activate the terminal carboxyl groups of the HPA. The solution was then applied to the gelatin solution and the reaction was conducted at 40 °C for 24 h. The resulting solution was subjected to the sequential dialysis in 100 mM of NaCl solution for 2 days, distilled water/ethanol for 1 day, and distilled water for 1 day (MWCO = 3500 Da). The purified solution was filtered and lyophilized to obtain the GHPA. The degree of substitution (DS) of the HPA was measured using an UV spectrometer (Jasco, V-750 UV/vis/NIR, Japan). The GHPA polymer was dissolved in distilled water (1 mg/mL), and the absorbance was measured at 275 nm wavelengths. The concentration of the conjugated HPA molecules was calculated from a calibration curve given by monitoring the absorbance of a known concentration of HPA, and standardized with the baseline measured using gelatin solution (1 mg/ mL). The chemical structure of the GHPA conjugate was characterized using 1H NMR spectroscopy and the GHPA polymer (10 mg/mL) was dissolved in D2O for the measurement. 2.4. Gelation Time Measurement and Rheological Experiments. The GHPA and Tet−TA hybrid hydrogels (200 μL) were prepared in a 1 mL vial at RT. A total of 100 μL of the GHPA in H2O2 solution (0.05 wt % of stock solution) and 100 μL of the Tet−TA in HRP solutions (0.0025−0.025 mg/mL of stock solutions) were mixed and gently shaken. The hydrogels were prepared at different concentrations of the GHPA. The gelation time of the hybrid hydrogels in 0.01 M PBS (pH 7.4) was determined through the vial tilting methods.8,29 The gel state was regarded when no flow was observed within a minute after inversion of the vial. The experiments were carried out in triplicate. The elastic (G′) and viscos modulus (G″) of the hybrid hydrogels were measured using Advanced Rheometer GEM-150-050 (Bohlin Instruments, U.S.A.) in the oscillatory mode at 37 °C. For the rheological experiments, 100 μL of the Tet−TA dissolved in 0.005 mg/mL of HRP solution and the same volume of the different concentrations of GHPA (2−10 wt %) dissolved in different concentrations of H2O2 solutions (0.02−0.05 wt %) were placed on the plate of the instrument. The Tet−TA solution was gently mixed with the GHPA solutions on the plate and the upper plate was immediately lowered down. The G′ and G″ values were recorded at a frequency of 0.1 Hz and a strain of 0.1% (strain control) using parallel plate geometry (diameter = 25 mm, gap = 0.5 mm). Table 1 summarized the prepared hydrogel samples with different concen-
In situ cross-linkable hybrid hydrogels composed tyramineconjugated 4-arm-polypropylene oxide−polyethylene oxide (Tet−TA)28 and gelatin−hydroxyphenyl propionic acid (GHPA) were prepared. Tet−TA and GHPA conjugates were synthesized using a carbodiimide/active ester-mediated coupling reaction. The chemical structures of the conjugates were characterized by 1H NMR and UV−Vis spectroscopic measurements. The physico-chemical properties, such as gelation time, mechanical properties, and in vitro proteolytic degradation behavior were evaluated. The cyto-compatibility of the hybrid hydrogels was studied by 3D culture of human dermal fibroblasts in the hydrogels. Moreover, the hydrogels were also subcutaneously injected in the back of mice to verify that the hydrogels can be rapidly formed at the injected site and may exhibit good tissue compatibility.
2. MATERIALS AND METHODS 2.1. Materials. Tetronic 1307 (4-arm-polypropylene oxide (PPO)−polyethylene oxide (PEO), MW = 18000 g/mol) was obtained from BASF corp. Gelatin (type A from porcine skin, >300 bloom), hydroxyphenyl propionic acid (HPA), 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), N-hydroxysuccinimide (NHS), horseradish peroxidase (HRP, type IV, essentially salt-free, lyophilized powder, 250−330 units/mg solid), hydrogen peroxide (H2O2), 4dimethylamino pyridine (DMAP), p-nitrophenylchloroformate (PNC), and collagenase type II were supplied from Aldrich-Sigma. Tyramine (TA) was purchased from Acros. Triethylamine (TEA) was supplied from Kanto Chemical Corp. Aluminum oxide was purchased from Strem Chemical Inc. All other chemicals and solvents were used without further purification. For the cell study, Dulbecco’s modified Eagle medium (DMEM), fetal bovine serum (FBS), trypsin/EDTA, penicillin−streptomycin (PS), phosphate-buffered saline (PBS, pH 7.4), and Dulbecco’s phosphate buffered saline (DPBS) were purchased from Gipco BRL. Live/dead cell viability/cytotoxicity kit was obtained from Invitrogen Corp. For the in vivo subcutaneous injection study, the rats were purchased from Orient Bio Corp. 2.2. Synthesis of Tyramine Conjugated 4-Arm-PPO−PEO (Tet−TA). Tet−TA linked through a urethane bond was synthesized using PNC, DMAP, and TEA as coupling reagents, as previously reported.28 Briefly, the Tet−TA was synthesized by activating the terminal hydroxyl groups of the 4-arm-PPO−PEO (Tetronic) copolymer with PNC, DMAP, and TEA, and subsequently, conjugating TA molecules to the amine reactive 4-arm-PPO−PEO− PNC (Tet−PNC) to produce Tet−TA. Tetronic (30 g, 1.67 mmol) was dissolved in 300 mL of dichloromethane (DCM) at room temperature (RT) under a nitrogen atmosphere. DMAP (1.018 g, 8.33 mmol) and TEA (0.843 g, 8.33 mmol) dissolved in 20 mL of DCM were added to the Tetronic solution to activate the terminal hydroxyl groups. The mixture was reacted at RT for 15 min. PNC (1.679 g, 8.33 mmol) dissolved in 30 mL of DCM was then added to the activated Tetronic solution and reacted under a nitrogen atmosphere at RT. After 24 h of reaction, the solution was filtered by glass filtration using aluminum oxide to remove the PNC salts and precipitated in cold diethyl ether twice. The precipitate was filtered and dried overnight under vacuum. 605
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
Figure 1. Synthesis of the polymer conjugates: (a) Tet−TA and (b) GHPA conjugates. trations of H2O2 and composition of Tet−TA and GHPA. The numbers indicated the final concentrations of the polymer solutions. 2.5. In Vitro Proteolytic Degradation Study of the Hydrogels. The hydrogels (300 μL) based on the different concentrations of GHPA and H2O2 were prepared in microtubes, and subsequently incubated at 37 °C in 1 mL of PBS (0.01 M, pH 7.4) containing 0.02 wt % of collagenase for predetermined time intervals. At the time point, the PBS solutions were removed from the microtubes, and the weight of the hydrogels (Wd) was measured. Fresh media were added to the vials after the weighting. The percentage of the remaining hydrogels was calculated using the following equation:
glucose DMEM, which included 10% FBS and 1% PS. The cell viability and morphology in the hydrogels were observed by Live/ Dead viability/cytotoxicity kit. After 3 d of incubation, 500 μL of 2 mM of the acetomethoxy derivate of calcein (calcein AM) and 4 mM of ethidium homodimer-1 (EthD-1) mixture was treated to each well at 37 °C for 30 min. The stained samples were then observed by fluorescent microscopy (TE2000, Nikon, Japan). For the cell study, the hydrogels were prepared by using Dulbecco’s phosphate buffered saline (DPBS). 2.7. In Vivo Subcutaneous Injection of the Hybrid Hydrogels. To confirm the in situ formation of the hybrid hydrogels and tissue invasion into the hydrogel matrix from the surrounding tissue, in vivo subcutaneous injection of the hybrid hydrogels was performed using different concentrations of the GHPA. The in vivo study was approved by institutional animal care and use committee of the Korea food and drug administration (KFDA). The polymer solutions were sterilized by filtering using 0.2 μm syringe filters prior to the experiment and injected using a dual syringe with a 26-gauge needle into the back of mice. The two syringes were filled up with Tet−TA in 0.05 mg/mL of HRP solutions and GHPA in 0.05 wt % of H2O2. These polymer solutions (400 μL) were then subcutaneously injected into the mice (5−6 weeks aged female). At 2 weeks postimplantation, the cutaneous tissues were harvested and fixed using glutaraldehyde solution. The fixed samples were then dehydrated and embedded in paraffin. The samples were stained with hematoxylin and eosin (H&E).
weight of hydrogels(%) = (Wd/Wi) × 100% where Wd and Wi are the weights of the degraded and initial hydrogels, respectively. 2.6. Cyto-Compatibility of the Hybrid Hydrogels. In vitro cyto-compatibility of the hybrid hydrogels with the different compositions of Tet−TA and GHPA (table 1) were evaluated using hDFBs. The polymer solutions were filtered using a syringe filter with a pore size of 0.2 μm (cellulose acetate disposable filter, DISMIC25cs) for the sterilization and placed in the 48-well plate. Briefly, the cells were mixed with 10 wt % of Tet−TA dissolved in 0.05 mg/mL of HRP solutions. The polymer solutions containing hDFBs (80 μL) were put into the wells and the same volume of H2O2 solutions (0.03 wt %) with the different concentrations of GHPA were added. The cell-encapsulated hydrogels were gently mixed until it was homogeneous. The density of the cells in the hydrogel matrices was 1.5 × 105 cells/mL and the matrices were cultured for 3 d under standard cell culture conditions (37 °C and 5% CO2) with high 606
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
3. RESULTS AND DISCUSSION 3.1. Synthesis of Tet−TA and GHPA Conjugates. A Tet−TA linked with urethane linkage was synthesized by coupling TA to the Tet−PNC without any coupling reagents. The synthetic route of the Tet−TA is shown in Figure 1a. The terminal hydroxyl groups of Tetronic were activated using excess PNC to form amine−reactive groups. The Tet−PNC conjugate was then mixed with excess TA to produce the Tet− TA conjugate. The feed molar ratio of Tet−PNC to TA was 1:5. The chemical structure of the conjugate was characterized by 1H NMR spectroscopy. The degree of TA substitution to Tet−PNC was calculated by integrating the peaks at δ 1.12 and δ 6.7−7.1 ppm, determined to be approximately 97%. Figure 2a
synthetic route used for the fabrication of the GHPA conjugate is illustrated in Figure 1b. The carboxyl groups of the HPA were activated using EDC/NHS and reacted with the amine groups of gelatin. The chemical structure of the GHPA conjugate was analyzed by 1H NMR and the results are shown in Figure 2b. 1H NMR: δ 4.8 (m, the protons of anomeric carbon of gelatin), δ 0.8−4.6 (m, alkyl protons of gelatin), and δ 6.8−7.1 (m, aromatic protons of TA). The feed amount of the HPA was 4−20 mmol to control the degree of substitution. The content of the HPA in gelatin was determined at the wavelength of 275 nm by UV measurement and the amount of HPA was 80−120 μmol per 1 g of GHPA. This is in agreement with recent results reported for polysaccharides or protein conjugated with phenol derivatives.15,16,19 3.2. In Situ Cross-Linkable Hydrogel Formation and Gelation Time. The hydrogels composed of Tet−TA and GHPA were formed through an enzyme−mediated coupling reaction using HRP and H2O2. In this reaction, HRP allows the phenol moieties both in Tet−TA and GHPA conjugate to couple to each other through the C−C bond at the ortho positions or the C−O bond between the carbons and the phenoxy oxygens under physiological conditions. Figure 3 shows the schematic illustration of the gelation and an image of the in situ formation of the hybrid hydrogel using a dual syringe. The syringe contains solutions of Tet−TA and HRP in channel (a), and GHPA and H2O2 in channel (b). The gelation time of the hybrid hydrogels was determined using the vial tilting method and Figure 4 shows the gelation time depending on the HRP concentrations. The gelation time could be controlled easily by varying the HRP concentration and it ranged from 5 to 300 s. In the case of the T5G5 hydrogel, the gelation time of the hydrogels formed with 0.0125 mg/mL of HRP was 5 s, while that of hydrogels formed with 0.0015 mg/mL of HRP was 70 s. This result is most likely due to an increment in the rate of phenoxy radical generation by the decomposition of H2O2. The gelation rate of the Tet−TA hydrogels was faster than that of the GHPA hydrogels due to the higher phenol concentrations and molecular mobility, which were 11 μmol/mL in the Tet−TA (5 wt %) and 6 μmol/
Figure 2. 1H NMR spectra of the Tet−TA (a) and GHPA (b) conjugates.
shows the 1H NMR spectrum of the Tet−TA conjugate; 1H NMR (CDCl3): δ 1.12 and 3.2−4.4 (m, proton of Tetronic), δ 6.7−7.1 (m, aromatic protons of TA). The GHPA conjugate was synthesized using a common carbodiimide/active ester-mediated coupling reaction. The
Figure 3. Schematic representation of enzyme-mediated cross-linking of Tet−TA and GHPA conjugates and the image of hydrogel formation using a dual syringe: (a) 10 wt % of Tet−TA dissolved in 0.005 mg/mL of HRP and (b) 10 wt % GHPA dissolved in 0.05 wt % of H2O2. 607
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
Figure 4. Gelation time of the Tet−TA and GHPA hydrogels as a function of HRP concentration.
mL in the GHPA (5 wt %), respectively. For instance, the gelation rate of the Tet−TA hydrogels formed with 0.0015 mg/ mL of HRP was 200 s, whereas that of the GHPA hydrogels prepared with the same concentration of HRP was 300 s. Moreover, the concentrations of GHPA have an effect on the gelation time. As the GHPA concentration was increased from 0 to 5 wt %, the gelation time of the hybrid hydrogels formed with 0.0015 mg/mL of HRP remarkably decreased from 200 to 70 s. This faster gelation process may be likely attributed to a higher amount of phenol residues in the hydrogels by introducing the GHPA, which can be cross-linked within the gel matrices. The concentration of phenol molecules in the matrices increased from 11 to 17 μmol/mL. A similar result was obtained by Kurisawa et al., who found that increasing HRP and phenol concentrations decreased the gelation times of tyramine-conjugated hyaluronic acid and phenol−conjugated gelatin hydrogels.16,19 It should be noted that a gelation time is important when the in situ cross-linkable hydrogels are applied for various biomedical applications. For the tissue regenerative medicine, an adequate gelation time is required to fit the regular or irregular defect site and to allow easy handling. On the other hand, for in situ drug delivery, a fast gelation time is necessary to prevent the diffusion of flowing hydrogel precursor solution and loaded bioactive molecules into the target site and surrounding tissue. From this aspect, the hydrogels prepared by HRP and H2O2 can be utilized as an injectable scaffold for tissue engineering and delivery of growth factors or proteins. 3.3. Controllable Mechanical Properties of the Hybrid Hydrogels. The mechanical properties of the hybrid hydrogels were measured in time−controlled oscillatory mode with different GHPA and H2O2 concentration. Figure 5 shows the G′ and G″ values as a function of time for the hybrid hydrogels fabricated with various conditions. In our previous report, we demonstrated that the G′ values of the Tet−TA hydrogels could be controlled by changing the H2O2 concentration and the values ranged from 200 to 15700 Pa.29 The G′ and G″ values of the hybrid hydrogels fabricated with different concentrations of GHPA and H2O2 were measured on the mechanical strength. An increase in the concentration of GHPA led to higher elastic and viscos moduli due to the higher crosslinking density owing to the high concentration of phenol groups in GHPA (Figure 5a). For instance, the G′ and G″ of the Tet−TA hydrogel was 1540 and 108 Pa, while the values of the T5G5 hydrogel were 8470 and 200 Pa. The concentrations of phenol moieties, which can be cross-linked in the gel matrices, were 11 μmol/mL in the Tet−TA and 17 μmol/mL in T5G5 hydrogels. Accordingly, the mechanical properties of the T5G5
Figure 5. Elastic (G′) and viscous (G″) moduli of the hybrid hydrogels: (a) the effect of the GHPA concentration on the mechanical properties of the hydrogels prepared with 0.005 mg/mL of HRP and 0.05 wt % of H2O2, and (b) the effect of the H2O2 concentration on the mechanical properties of the hydrogels fabricated with different concentration of H2O2 (0.02−0.05 wt %) and 0.005 mg/ mL of HRP.
hydrogels were higher than that of the Tet−TA hydrogels. When the H2O2 concentration was increased from 0.02 to 0.05 wt %, the G′ and G″ of the T5G5 hydrogels increased from 710 and 60 Pa to 8470 and 200 Pa, respectively, because the higher concentration of H2O2 generated more phenoxy radicals in enzyme-mediated reaction (Figure 5b). Interestingly, the mechanical strength of the GHPA hydrogel (G′ = 2010 Pa) was higher than that of Tet−TA hydrogel (G′ = 1540 Pa), although the phenol concentration of the GHPA polymer (6 μmol/mL) was lower than that of the Tet−TA polymer (11 μmol/mL). This result may be attributed to the inherent characteristic of the polymer backbone, which demonstrated that the GHPA moiety could influence on the mechanical properties of the hydrogels through increase of cross-linking density and synergistic effect. 3.4. In Vitro Proteolytic Degradation Behavior of the Hybrid Hydrogels. The in vitro proteolytic degradation behavior of the hybrid hydrogels was investigated using collagenase in PBS. The method was adequate for determining the degradation behavior of the hydrogel, although the experimental method did not closely simulate the degradation behavior in the in vivo environment. Figure 6 shows the in vitro proteolytic degradation behaviors of the hybrid hydrogels. The GHPA hydrogels were completely degraded within 6 h through 608
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
This result may be caused by the enhancement of the hydrogels stability due to the slowly-degradable Tet−TA moiety in the matrix. Tet−TA hydrogel was also degraded for 60% in 3 days. Tetronic is composed of poly(propylene oxide) and poly(ethylene oxide), which is well-known as nondegradable polymers. However, the Tet−TA hydrogel was slowly degradable due to the urethane linkages, which is formed during the conjugation of TA and Tet−PNC. Generally, the urethane bond is gradually degraded through the hydrolytic and proteolytic degradation.30 The degradation was dependent on the use of Tet-TA, showing only slightly different degradation kinetics when measured as functions of GHPA or H2O2 concentrations (Figure 6a,b). The hydrogels formed with a lower concentration of H2O2 showed a slightly faster degradation rate compared with the hydrogels prepared with a higher concentration of H2O2, retaining about 30% of its initial weight after 3 d of incubation. A similar result has been reported that the lower cross-linking density resulted in a fast degradation rate for the gelatin−based hydrogels, which are cross-linked with glutaraldehyde and via enzyme-mediated reaction.19,27 The proteolytic degradable properties of the synthetic matrix were important factors because the remodeling of the artificial extracellular matrix is necessary for the tissue or cell invasion. A number of proteolytically degradable hydrogels have been widely studied.31−33 Rizzi et al. reported that the cells cultured in the MMP-sensitive PEG hydrogels were well spread and showed more proliferation than those cultured in the nondegradable hydrogels.33 3.5. Cell Viability of the hDFBs in the Hybrid Hydrogel Matrices. To assess cyto-compatibility of the hybrid hydrogels, hDFBs were cultured within the hydrogel matrices at the different polymer components for 3 days. The cultured cells were observed using fluorescent microscopy by live/dead staining. Figure 7 shows the fluorescence images of the hDFBs cultured in the hydrogels matrices. Most cells cultured in the Tet−TA hydrogel were viable, but the cell spreading and proliferation were limited, resulting that the morphology of the cells adopted a round shape due to the low cyto-compatibility of the Tet−TA hydrogels as shown in Figure 7a,f. In contrast, most cells were much more viable in the hybrid hydrogels and the cells were well−spread (Figure 7b−d and g−i). For instance, the cells cultured within the hybrid hydrogels (T5G5) were increasingly spindle-shaped and polygonal, and a similar result was shown in the GHPA hydrogels. Interestingly, as
Figure 6. In vitro proteolytic degradation of the hybrid hydrogels in 0.01 M of PBS (pH 7.4) containing 0.02 wt % of collagenase; the effect of the GHPA (a) and H2O2 (b) concentrations on the degradation profiles of the T5G5 hydrogels.
the decomposition of the peptide chains via proteolytic degradation as shown in Figure 6a. A similar study was investigated by Sakai et al., showing that the gelatin hydrogels were completely degraded within 24 h in the media containing the papain.19 However, the fast degradation of the hydrogel matrices may be disadvantageous when implanted in vivo. A prolonged degradation of the hydrogels were obtained as a result of hybridization of GHPA and Tet−TA in the hydrogels, reaching 40% of their initial weights after 3 d of incubation.
Figure 7. Fluorescent microscopic images of the hDFBs in the hydrogels for the live/dead assay: (a, f) Tet−TA, (b, g) T5G1, (c, h) T5G3, (d, i) T5G5, and (e, j) GHPA hydrogels. Scale bar (100 μm) was reprinted. 609
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
Figure 8. Histological images of the implanted hydrogels on the back of the mice: the cutaneous tissues containing the injected hydrogels were analyzed at day 14. Star marks and arrows indicated the hydrogels and tissue infiltration from the surrounded tissue, respectively; (a) Tet−TA, (b) T5G1, (c) T5G3, and (d) T5G5.
gels. The hDFBs exhibited a high viability after the encapsulation in the hydrogels via enzyme-mediated crosslinking process and 3 days of incubation. In the subcutaneous injection study, the hybrid hydrogels rapidly formed at the injected site and in-growth of natural tissues into the matrices. From the in vivo and in vitro evaluations, we demonstrated that the hybrid hydrogel had biocompatibility and bioactivity. The combined results demonstrate that the in situ forming Tet−TA and GHPA hybrid hydrogels with tunable properties, good stability, and a superior bioactivity hold great promise for use as injectable materials for various biomedical applications.
increased the GHPA concentration, the cells were more spread within the hydrogel matrices, suggesting that the introduction of the gelatin moiety into the Tet−TA hydrogel matrices enhanced cellular activity and the in situ forming Tet−TA and GHPA hybrid hydrogels formed via an enzyme-mediated reaction were highly cyto-compatible materials. 3.6. In Vivo Subcutaneous Injection. To confirm the feasibility of using the in situ cross-linkable Tet−TA and GHPA hybrid hydrogels as an injectable matrix, the hydrogels were injected subcutaneously in the back of mice. These in vivo injection experiments demonstrated successful in situ gelation of the hybrid hydrogels and prolonged in vivo stability was observed over 2 weeks. The GHPA hydrogels were completely degraded within 2 weeks in vivo. To observe the histological reaction to the hydrogels, the surrounding tissues and the hydrogels were stained with hematoxylin and eosin (H&E) 2 weeks after implantation. Figure 8 shows the histological images (H&E staining) of the implanted hydrogels containing the cutaneous tissue at the different compositions. The invasion of the surrounded tissue was observed when the GHPA was incorporated within the gel matrices, while no tissue invasion into the Tet−TA gel matrix was observed due to the low tissue compatibility. In addition, the concentration of GHPA in the gel matrices also influenced the tissue infiltration. Only minimal infiltration of tissues or cells in the hydrogels were observed when the concentration of the GHPA conjugate was low (T5G1), whereas significant infiltration of the surrounding tissues or cells into the hydrogel was observed at high GHPA concentrations (T5G5). Although the surrounding tissue was infiltrated into the gel matrices, inflammatory reactions were also observed in the samples surrounded by a fibrous encapsulation. However, the formation of fibrous layers over implanted materials in vivo has been frequently reported for a variety of feasible biomaterials.34,35 This limitation of the hybrid hydrogels can be passed by incorporating the anti-inflammatory drug like dexamethasone within the gel matrix. Actually, we are trying to incorporate the drugs within the hybrid hydrogel matrix to overcome the problems in our future work.
■
AUTHOR INFORMATION
Corresponding Author
*Fax: +82-31-219-1592. Tel.: +82-31-219-1846. E-mail: kdp@ ajou.ac.kr. Notes
The authors declare no competing financial interest.
■
ACKNOWLEDGMENTS This work was supported by the grants from the NanoBiotechnology Project (Regenomics) funded by the Ministry of Education, Science and Technology (MEST; B020214) and the NRF grant funded by the Korea government (2010-0027776).
■
REFERENCES
(1) Burkoth, A. K.; Anserth, K. S. Biomaterials 2000, 21, 2395. (2) Klouda, L.; Mikos, A. G. Eur. J. Pharm. Biopharm. 2008, 68, 34. (3) Rowley, J. A.; Madlambayan, G.; Mooney, D. J. Biomaterials 1999, 20, 45. (4) Doria-Serrano, M. C.; Ruiz-Trevino, F. A.; Rios-Arciga, C.; Hernandez-Esparza, M.; Santiago, P. Biomacromolecules 2001, 2, 568. (5) Kuo, C. K.; Ma, P. X. Biomaterials 2001, 22, 511. (6) Bhattarai, N.; Ramay, H. R.; Gunn, J.; Matsen, F. A.; Zhang, M. Q. J. Controlled Release 2005, 103, 609. (7) Gong, C. Y.; Shi, S. A.; Dong, P. W.; Kan, B.; Gou, M. L.; Wang, X. H.; Li, X. Y.; Luo, F.; Zhao, X.; Wei, Y. Q.; Qian, Z. Y. Int. J. Pharm. 2009, 365, 89. (8) Park, K. M.; Joung, Y. K.; Na, J. S.; Lee, M. C.; Park, K. D. Acta Biomater. 2009, 5, 1956. (9) Hennink, W. E.; De Jong, S. J.; Bos, G. W.; Veldhuis, T. F.; van Nostrum, C. F. Int. J. Pharm. 2004, 277, 99. (10) Bos, G. W.; Hennink, W. E.; Brouwer, L. A.; den Otter, W.; Veldhuis, T. F.; van Nostrum, C. F.; van Luyn, M. J. Biomaterials 2005, 26, 3901. (11) Jun, Y. J.; Park, K. M.; Joung, Y. K.; Park, K. D.; Lee, S. J. Macromol. Res. 2008, 16, 704. (12) Pratt, A. B.; Weber, F. E.; Schmoekel, H. G.; Muller, R.; Hubbell, J. A. Biotechnol. Bioeng. 2004, 86, 27. (13) Shu, X. Z.; Liu, Y. C.; Palumbo, F. S.; Lu, Y.; Prestwich, G. D. Biomaterials 2004, 25, 1339.
4. CONCLUSION In situ cross-linkable hybrid hydrogels were developed as a bioactive and injectable matrix with high tissue compatibility as well as a sufficiently long-term stability in vivo. The hybrid hydrogels were rapidly formed by an enzyme-mediated networking of Tet−TA and GHPA using HRP and H2O2. The gelation time and mechanical properties of the hybrid hydrogels were easily and reliably adjusted by changing the concentration of the HRP and H2O2. The hybrid hydrogels showed the enhanced stability comparing with GHPA hydro610
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
Biomacromolecules
Article
(14) Sharifi, S.; Mirzadeh, H.; Imani, M.; Atai, M.; Ziaee, F. J. Biomed. Mater. Res., Part A 2008, 84A, 545. (15) Jin, R.; Hiemstra, C.; Zhong, Z.; Feijen, J. Biomaterials 2007, 28, 2791. (16) Kurisawa, M.; Chung, J. E.; Yang, Y. Y.; Gao, S. J.; Uyama, H. Chem. Commun. 2005, 4312. (17) Sakai, S.; Kawakami, K. J. Biomed. Mater. Res., Part. A 2008, 85A, 345. (18) Sakai, S.; Ogushi, Y.; Kawakami, K. Acta Biomater. 2009, 5, 554. (19) Sakai, S.; Hirose, K.; Taguchi, K.; Ogushi, Y.; Kawakami, K. Biomaterials 2009, 30, 3371. (20) Wang, L.-S.; Boulaire, J.; Chan, P. P. -Y.; Chung, J. E.; Kurisawa, M. Biomaterials 2010, 31, 8608. (21) Wang, L.-S.; Chung, J. E.; Chan, P. P.-Y.; Kurisawa, M. Biomaterials 2009, 31, 1148. (22) Sakai, S.; Matsuyama, T.; Hirose, K.; Kawakami, K. Biomacromolecules 2010, 11, 1370. (23) Karim, A. A.; Bhat, R. Trends Food Sci. Technol. 2008, 19, 644. (24) Mano, J.; Silva, G.; Azevedo, H.; Malafaya, P.; Sousa, R.; Silva, S.; Boesel, L.; Oliveira, J.; Santos, T.; Marques, A. J. R. Soc. Interface 2007, 4, 999. (25) Young, S.; Wong, M.; Tabata, Y.; Mikos, A. G. J. Controlled Release 2005, 109, 256. (26) Chawla, K.; Yu, T. B.; Liao, S. W.; Guan, Z. Biomacromolecules 2011, 12, 560. (27) Tabata, Y.; Nagano, A.; Ikada, Y. Tissue Eng. 1999, 5, 127. (28) Park, K. M.; Jun, I.; Joung, Y. K.; Shin, H.; Park, K. D. Soft Matter 2011, 7, 986. (29) Park, K. M.; Shin, Y. M.; Joung, Y. K.; Shin, H.; Park, K. D. Biomacromolecules 2010, 11, 706. (30) Guelcher, S. A. Tissue Eng., Part B 2008, 14, 3. (31) Lee, S. H.; Moon, J. J.; Miller, J. S. Biomaterials 2007, 28, 3163. (32) Lévesque, S. G.; Shoichet, M. S. Bioconjugate chem. 2007, 18, 874. (33) Rizzi, S. C.; Hubbell, J. A. Biomacromolecules 2006, 7, 3019. (34) Lemperle, G.; Morhenn, V.; Charrier, U. Aesthet. Plast. Surg. 2003, 27, 354. (35) Lemperle, G.; Morhenn, V.; Pestonjamasti, V.; Gallo, R. L. Plast. Reconstr. Surg. 2004, 113, 1380.
611
dx.doi.org/10.1021/bm201712z | Biomacromolecules 2012, 13, 604−611
