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Preparation and characterization of dual-crosslinked gelatin hydrogel via Dopa-Fe3+ complexation and fenton reaction
Accepted Manuscript Title: Preparation and Characterization of Dual-crosslinked Gelatin Hydrogel via Dopa-Fe3+ Complexation and Fenton reaction Authors: Ji Yeon Kim, Seung Bae Ryu, Ki Dong Park PII: DOI: Reference:
S1226-086X(17)30487-2 https://doi.org/10.1016/j.jiec.2017.09.014 JIEC 3617
To appear in: Received date: Revised date: Accepted date:
22-6-2017 25-8-2017 8-9-2017
Please cite this article as: Ji Yeon Kim, Seung Bae Ryu, Ki Dong Park, Preparation and Characterization of Dual-crosslinked Gelatin Hydrogel via DopaFe3+ Complexation and Fenton reaction, Journal of Industrial and Engineering Chemistry https://doi.org/10.1016/j.jiec.2017.09.014 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Preparation and Characterization of Dual-crosslinked Gelatin Hydrogel via Dopa-Fe3+ Complexation and Fenton reaction
Ji Yeon Kim*, Seung Bae Ryu* and Ki Dong Park#
Department of Molecular Science and Technology, Ajou University, Suwon, 443-749, Republic of Korea
#Corresponding
author: Department of Molecular Science and Technology, Ajou University, 5
Woncheon, Yeongtong, Suwon, 443-749, Republic of Korea; Fax: +82 31 219 1592; Tel: +82 31 219 1846; E-mail address: [email protected]
* These authors contributed equally to this work.
Graphical abstract
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Abstract Gelatin hydrogel via Dopa-Fe3+ complexation and Fenton reaction was developed as a dual-crosslinked hydrogel fabrication method for biomedical application. In order to prepare these hydrogels, methacrylated gelatin dopamine (GMD) conjugates were synthesized. The hydrogels were formed rapidly by using hydrogen peroxide (H2O2) and iron(II) chloride (FeCl2), the acryl moieties at the gelatin conjugate react via Fenton reaction and created Fe3+ ions as a byproduct. These Fe3+ ions were used to form Dopa-Fe3+ complexes. The mechanical strength could be controlled by varying the H2O2 concentration and exhibited 2–3 times higher than that single-crosslinked hydrogels. Furthermore, the swelling ratio and degradation behavior of the hydrogels were controllable depending on the crosslinking density. It was also found that the GMD hydrogels have tissue adhesive ability because of the dopamine content. The developed GMD hydrogels without oxidizing agents under mild conditions can be used as injectable materials in tissue regenerative medicine and various biomedical applications.
Keywords: dopamine, Fenton reaction, dual crosslinking, gelatin, Dopa-Fe3+ coordination complex
1. Introduction Biomedical carriers have been widely used to deliver drugs with sustained and controlled release at the target cells or tissues. Several novel drug carriers have been employed in the biomedical field, including liposomes, micelles, and nanoparticles [1, 2]. Compared to these known carriers, injectable hydrogels have attracted significant
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attention in recent years. They also exhibit several advantages over traditional carriers. First, hydrogels with 3D networks can load large amounts of drugs and release drugs in a sustained and controlled manner, which can lead to better therapeutic results [3, 4]. Second, injectable hydrogels can release drugs directly to the target cells or tissues, avoiding circulation of the drug in the entire body [3]. In this manner, they can quickly deliver a large dose of drugs [5]. Moreover, for cancer patients whose physical conditions are not suitable for intensive surgery, the use of injectable hydrogels can significantly reduce treatment risks as their application is minimally invasive operation. In order to prepare an injectable hydrogel, various methods such as enzyme-catalyzed reactions, Schiff base reactions, and photo-crosslinking reactions have been employed [6-8]. From these, Fenton reaction induced hydrogel crosslinking using ferrous salt (Fe2+) and hydrogen peroxide (H2O2) that generate hydroxyl radical could initiating polymerization of vinyl monomers. This reaction have some advantages such as rapid crosslinking time, easy drug/cell loading, and prolong the drug release profiles compared with photo-crosslinked hydrogels[9-11] Among these injectable hydrogels, dopamine-conjugated hydrogels offer a biomimetic approach toward efficient and safe drug delivery [12]. The dopamine moiety in the hydrogel helps it to attach itself to various organic and inorganic surfaces, even in wet conditions [13-16]. Hydrogels with catechol-containing polymers can be formed by adding oxidizing agents such as periodate ([IO]4−) [17, 18] and metal (e.g., Fe(III)) [1923] ions, which can provide a highly biocompatible 3D matrix. The organic oxidizing agent [IO]4− helps to form covalently cross-linked (i.e., irreversible) hydrogels, although the use of an excess amount of [IO]4− may result in toxicity. A transition-metal ion such as Fe(III) assists in the formation of Dopa-Fe3+ coordination-based (i.e., reversible) hydrogels. However, the stoichiometric ratio of Dopa-Fe3+ coordination is 1:3 when the catechol moiety is exposed to relatively high pH values (>8). When the pH value is decreased, the tri-coordinated Dopa-Fe3+ complex transforms to a mono- or bicoordinated complex, resulting in either weak mechanical strength or no gelation.
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In this study, methacrylated gelatin dopamine (GMD) hydrogels were prepared that were dually cross-linked via Dopa-Fe3+ complexation and Fenton reaction under mild conditions. Such crosslinks improved the mechanical strength and stability of the dopamine-conjugated hydrogels. The Fenton reaction typically relies on the reduction of hydrogen peroxide at the expense of Fe2+ ions and the Fenton reagent has been widely used as a radical initiator in vinylic polymerization or grafting. The chemical structures of the GMD conjugates were characterized by
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H NMR and UV
spectroscopies. The physicochemical properties such as mechanical strength, adhesive ability, and degradation behavior were evaluated for the hydrogels formed by using different concentrations of H2O2 in the Fenton cross-linking process. In addition, the stability of the hydrogels in a simulated body fluid (SBF) was also investigated in order to understand the effect of body fluids on the hydrogels.
2. EXPERIMENTAL 2.1 Materials Gelatin (type A from porcine skin, >300 bloom), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), N-hydroxysuccinimide (NHS), dopamine hydrochloride (DA), methacrylic anhydride (MA), iron(II) chloride (FeCl2), hydrogen peroxide (H2O2), and collagenase type II were obtained from Sigma Aldrich (St. Louis, MO, USA). Fetal bovine serum (FBS) was purchased from Wisent (Saint-Bruno, QC, Canada) and hydrochloric acid, 35.0–37.0% (HCl) was obtained from Samchun (Seoul, Korea). Human dermal fibroblasts (DFBs) were supplied from Lonza Inc. (Walkersville, MD, USA). Dulbecco's modified Eagle medium (DMEM), fetal bovine serum (FBS), penicillin–streptomycin (P–S), trypsin–EDTA, and Dulbecco's phosphate buffered saline (DPBS) were purchased from Gibco BRL (Grand Island, NY, USA). EZ-Cytox Enhanced Cell Viability Assay Kit was purchased from ITSbio Inc. (Seoul, Korea).
2.2 Synthesis and characterization of methacrylated gelatin dopamine conjugates
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GMD conjugates were synthesized by using MA and DA, as shown in Figure 1. First, methacrylated gelatin (GMA) was synthesized by dissolving 5 g of gelatin in 100 mL of deionized water at 60 °C. Then, MA (0.625 μL) was added to the gelatin solution and allowed to react for 1 h. Finally, 400 mL of warm (40 °C) phosphate buffered saline (PBS, pH 7.4) was added to stop the reaction. The mixture was dialyzed against distilled water using 12–14 kDa cutoff dialysis tubing for three days at 40 °C to remove the salts and residual methacrylic acid. The synthesized GMA was lyophilized for 3 days to generate a white porous foam. Second, GMD and gelatin dopamine (GDA) were synthesized by adding dopamine hydrochloride to GMA and gelatin solution, respectively. The detailed reaction conditions are provided in Table 1. Briefly, 2 g of GMA and gelatin were each dissolved in 100 mL of PBS solution. Next, EDC, NHS, and dopamine hydrochloride were added and the pH of the mixtures was adjusted to 5.0. The reactions were left to stir for 7 h at 40 °C. Subsequently, the resulting solutions were each transferred into a dialysis bag (MWCO = 3.5 kDa) and dialyzed against deionized water at pH ~4 for 2 days and at pH = 7 for 1 day, filtered, and lyophilized to obtain the sponge-type GMD and GDA conjugates. The chemical structure of GMD, GDA, and GMA conjugates were characterized by 1H NMR spectroscopy (AS400, OXFORD instruments, UK). The dopamine contents of
the conjugates were measured quantitatively using the solution of GMD in distilled water at 1mg/mL and dopamine hydrochloride (0.01 - 0.1 mg/mL) as a standard curve at 280 nm with a UV visible spectrophotometer in agreement with the previously reported [24] .(V-750 UV/Vis/NIR, Jasco, Japan).
2.3 Preparation of hydrogels The GMD, GDA, and GMA hydrogels were formed by mixing PBS (pH 7.4) solution with FeCl2 and H2O2 (Figure 2). The GMD, GDA, and GMA polymer solutions (13.5 wt%) in PBS were divided into two aliquots, each containing either FeCl2 (2.5 – 10 mM)
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or H2O2 (0.005 – 0.05%). The volume ratio of polymer solution : FeCl2 or H2O2 was 9 : 1. In order to obtain hydrogels of these polymers, equal volumes from the two aliquots were mixed and shaken (Figure 3). The pH of all solutions used to fabricate hydrogels were 7.4.
2.4. Rheological measurement The elastic modulus (G') of the hydrogels were measured by using a rheometer (Advanced Rheometer GEM-150-050, Bohlin 80 Instruments, USA) in oscillation mode. 150 μL of each GMD, GMA, GDA polymer solutions containing either FeCl2 or H2O2 were mixed together and applied to the bottom plate by varying H2O2 concentration. The FeCl2 concentration dependent stiffness of GMD hydrogels were assessed as the same procedure. The mechanical strength of the formed hydrogel was then recorded for 10 min at 37 °C with frequency of 0.1 Hz and a strain of 0.01% using a parallel plate geometry (diameter = 25 mm, gap = 0.5 mm)
2.5. Swelling ratio Three replicas of each dried hydrogel were swollen in PBS at 37 °C for 6 h to achieve equilibrium swelling. The degree of swelling of the hydrogels was measured until 72 h. The standard deviations are marked as error bars in the swelling profile charts. The degree of swelling was calculated as follows: Degree of swelling = [(Wet weight – Initial weight) / Initial weight] ×100%
2.6 In vitro proteolytic degradation A total of 300 µL of GMD, GDA, and GMA hydrogels with different elastic modulus were prepared. After equilibrium for 30 min, the hydrogels were incubated at 37 °C in 1 mL of PBS (0.01 M, pH 7.4) with and without 0.005 mg mL−1 of collagenase. At predetermined time intervals, the media were removed and changes in the weight of
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the hydrogels were recorded. Fresh media were then added for the next time interval. The degree of degradation was calculated using the following equation: Weight of hydrogels (%) = (Wd / Wi) x 100 where Wi and Wd refer to the initial weight and weight of the degraded hydrogels, respectively.
2.7. Stability of hydrogels The GMD, GDA, and GMA hydrogels (50 µL) were fabricated in 1.5 mL microtubes, as described above. The GMD, GDA, and GMA polymer solutions (13.5 wt%) in PBS were divided into two aliquots, each containing either FeCl2 (10 mM) or H2O2 (0.05%) as follow mechanical strength test. Subsequently, 0.8 mL of simulated body fluid (SBF) supplemented with 0.5% (v/v) fetal bovine serum was added into each tube. The changes in the appearance of each sample were recorded as a function of incubation time at 37 °C.
2.8. Tissue adhesive strength Tissue adhesive strength of the hydrogels was investigated by using a universal testing machine (UTM; Instron 3343, Instron, USA) in accordance with the modified ASTM method F2255-05. Pre-dehydrated porcine skins (diameter 25 mm, in 0.01 M PBS, pH 7.4 for 1 h) were attached to the rubber in the testing machine by using cyanoacrylate glue for 1 h. The hydrogels formed with the FeCl2 (10 mM) and H2O2 solutions (10 µL, 0.005–0.05%) were placed between the substrates and cured for 15 s with 1N pressure. The adhesive strength was measured by equipping the UTM with a load cell of 100 N at a crosshead of 10 mm min−1. Each measurement was repeated five times.
2.9. Cytotoxicity test of hydrogels Cytotoxicity of GMD hydrogels against human dermal fibroblasts (hDFBs) were evaluated with indirect contact method, according to ISO10993 standard test.[25].
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Briefly, the hDFBs were cultured with DMEM containing 10% FBS, 1% P-S on the tissue culture plate (TCPs) at 37 °C and 5% CO2. The hydrogels were prepared (200 µL) with 1 mL of culture medium and kept at 37 °C and 5% CO2 for 24h to fully extract the unreacted residues. The hDFBs were seeded on the 48-well plate (2 ×105 cells / well) and cultured with extracted media for 24 h and the WST-1 assay is carried out. The cells were incubated at 37 °C and 5% CO2 with 1 mL of culture medium containing WST-1 reagents for 2 h. The absorbance of samples were measured at 450 nm using a microplate spectrophotometer (VersaMax tunable microplate reader; Molecular Devices, Sunnyvale, CA, USA). The cell viability of hydrogel extracted media was calculated using the following equation Cell viability (%) = (O.D. of the sample / O.D. of the control) x 100 All experiments were performed in triplicate.
3. RESULTS AND DISCUSSION 3.1. Synthesis and characterization of polymers First, the GMA polymers were synthesized by using MA to enable Fenton reaction as the first crosslinking strategy. The fraction of lysine groups reacted was modified by altering the concentration of MA. GMA was synthesized with 53.8±0.5% of methacrylation degree corresponding to 1.25% volume percentage of MA [26]. Next, the GMD and GDA conjugates were synthesized by grafting dopamine onto GMA and gelatin backbone via EDC/NHS chemistry, which served to enable Dopa-Fe3+ complexation as the second crosslinking strategy. The GMD, GMA, and gelatin conjugates were characterized by
1
H NMR spectroscopy (AS400, OXFORD
instruments, UK). The peaks observed in the range 6.7–6.9 ppm were attributed to the aromatic protons of dopamine. Compared to unmodified gelatin, a new signal was observed in spectra of the GMD and GMA conjugates at 5.3–5.6 ppm and 1.8 ppm, which was assigned to the methacrylate group (Figure 4a) [25]. Compared with gelatin solution, the enhanced absorption of UV light at λ = 280 nm is observed in the solution
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of the resultant GMA–dopamine conjugate and gelatin-dopamine conjugate, confirming the successful conjugation of dopamine (namely, catechol groups) (Figure 4b) [27]. Moreover, no peaks appear at λ = 395 nm, which indicate that the catechol groups in the gelatin–dopamine conjugate are not oxidized to quinones during the synthetic process [28]. It was found that the dopamine content of these gelatin derivatives reached its maximum amount (142.4 µmol g−1 of GMD and GDA) by taking advantage of the available functional group (–NH2) on GMA and gelatin. The two gelatin derivatives GMD and GDA, which represent the highest dopamine contents, were chosen for further study.
3.2. Rheological analysis of gelatin derivatives hydrogels The elastic modulus (G') against the shear strain is an indication of cohesive strength and was determined by varying either the concentrations of H2O2 or FeCl2 concentration at fixed FeCl2 (10 mM) or H2O2 (0.05%). The fixed concentration of FeCl2 (10 mM) and H2O2 (0.05%) were selected by measuring rheological analysis, in which mechanical strength of hydrogels were highest at each condition. The increasing concentration of H2O2 led to higher G' values because of increased cross-linking density of the hydrogels. The elastic modulus of the GMD hydrogels was found to be higher than that of GDA and GMA hydrogels at the same H2O2 concentration (Figure 5a). These results can be explained that the concentration of H2O2 increases, the formation of OH radicals and Fe3+ ions also increases that lead more crosslinking as both OH radicals and Fe3+ ions are involved in the dual-crosslinking of GMD hydrogels. As shown Figure 5b, FeCl2 concentration dependent mechanical strength increased with increasing FeCl2 concentration but decreased at more than 10 mM. This may due to the binding form of Fe-Dopa complexation is converted from the tri- and bis-form to the mono-form by increasing Fe ion concentration [29].
3.3 Swelling ratio of hydrogels on the crosslinking density
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Swelling ratio has been widely used as a simple method to characterize water absorption and the stability of biomaterials [30]. The swelling behavior of GMD, GDA, and GMA hydrogels in PBS as a function of time is depicted in Figure 6. It was found that most of the hydrogels attained equilibrium swelling in 12 h. After 12 h of swelling in PBS, the swelling ratio of the GMD hydrogel, which had high crosslinking density, reached 158±12%, while that of the GDA hydrogel having low crosslinking density was 280±14%. These results imply that the formation of a weak network makes these hydrogels porous and the strength of the elastic restoring forces is dependent on the extent of interconnectedness (crosslinking density) of the polymer network.
3.4. Proteolytic degradation behavior The gelatin backbone can be degraded by the action of proteases and its degradability is a very important factor in in situ forming hydrogels during interaction with the tissues in the body. Enzymatic degradation of the GMD, GDA, and GMA hydrogels was investigated in the presence of collagenase, a member of the matrix metalloproteinase family that degrades extracellular matrix components and enhances in vivo cell migration and growth. Figure 7 shows the in vitro proteolytic degradation behavior of the synthesized hydrogels. The GMD hydrogel incubated in PBS without collagenase remained stable for six days. In the presence of collagenase, the GMD hydrogel was completely degraded after six days. However, the GDA and GMA hydrogels were completely degraded within three and five days, respectively, because of the decomposition of the peptide chains by collagenase. Therefore, it was concluded that the degradation rates of the hydrogels depended on their cross-linking density because of limited accessibility of collagenase to the cleavage sites in crosslinked hydrogels. These results confirm that the dual-crosslinked hydrogels synthesized in this work can also be used as a biomimetic scaffold and their degradation can be controlled.
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3.5. Stability test in the simulated body fluid As internal human organs/tissues are always wet with body fluid, the stability of the hydrogels in simulated body fluid (SBF) solution containing 0.5% (v/v) of FBS was investigated. As seen in Figure 8, the freshly prepared hydrogels at the bottom of 1.5 mL microtubes were brown in color and the SBF solution was nearly colorless. After immersion for 6 h, all hydrogels were intact. Interestingly, the GDA hydrogel was completely dissolved after incubation for 24 h and at the same time, the color of the SBF solution became deeper compared with that after immersion for 6 h. These observations suggest that the Fe3+ ions are gradually lost from the gel system into the SBF solution containing FBS. The dissolution of GDA hydrogels might be attributed to Fe3+ binding to serum proteins in the SBF solution. The serum transferrin present in body fluids is known to strongly bind to Fe3+ ions [31]. Therefore, albeit with rapid gelation, the GDA hydrogels cannot serve as an ideal tissue adhesive owing to lack of long-term effectiveness under in vivo conditions. Actually, the natural MAPs solution which is secreted from a mussel, is also solidified through intermolecular covalent crosslinking after oxidation of its catechol groups to quinones takes place, catalyzed by the polyphenol oxidase enzyme [28, 32]. By learning from mussels, GMD can be further developed by adding a stable, second crosslinking moiety, namely, methacrylated gelatin–dopamine chains achieved from the Fenton reaction.
3.6. Effect of dopamine on the adhesive strength of hydrogels The adhesive properties of the GMD and GDA hydrogels having various mechanical strengths were measured by lap-shear test that assesses the adhesive strength of different materials to biological tissues (Figure 9). Considering the biological similarity to human tissues, fresh porcine skin was chosen as an adherend to closely mimic clinical conditions. The tissue adhesive strength of GMD hydrogels (15 kPa) was similar to that of GDA hydrogels at 0.05% of H2O2, although the mechanical strength of the latter was lower than that of the former. While
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the GMA hydrogels exhibited low tissue adhesion, the adhesive strength of GMD hydrogels was found to be 2-fold higher than that of the commercially available fibrin glue. These results can be attributed to the strong non-covalent binding interactions generated by the dopamine moieties to the tissue. Furthermore, the dopaminecontaining biopolymer is crosslinked by various means, among which a natural way involves the conversion of catechol groups into quinones by polyphenol oxidase enzyme in mussels, thus achieving covalent self-crosslinking.
3.7 Cytotoxicity test of the GMD hydrogels To assess the in vitro cytotoxicity of GMD hydrogels with different H2O2 concentrations (0.01 – 0.05%), we used the modified ISO 10993 test to the viability of hDFBs cultured in media extracted from hydrogels. The hydrogel extracts were obtained after 24 h of incubation with GMD hydrogels (13.5%) formed by FeCl2 (10 mM) and H2O2 (0.01%, 0.03%, and 0.05%). Culture without hydrogels were used as a control. As shown in Figure 10, the WST-1 assay shown the quantitatively analyze the cell viability. We observed no toxic effects on hDFBs, and overall metabolic rates of the cells exposed to the extracts from the hydrogels were comparable with control.
4. CONCLUSIONS In this study, a dual-crosslinked gelatin hydrogel was developed by combining DopaFe3+ complexation and Fenton reaction. The hydrogel was formed rapidly in the presence of H2O2 and FeCl2. Its physicochemical properties could be easily controlled by varying the concentrations of either H2O2 or FeCl2. It was found that the mechanical strength of the hydrogel ranged from approximately 100 to over 5,000 Pa, which is 2fold and 4-fold higher than that of GMA and GDA (single crosslink), respectively. Thus, the GMD hydrogels were more stable than GMA and GDA in the collagenase solutions. The introduction of dopamine moieties into the gelatin hydrogels led to their 4-fold and 2-fold higher adhesive strength as compared to methacrylated gelatin and fibrin glue,
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respectively. The dual crosslinking strategy is expected to be extensively employed to fabricate biomedical devices with polymers that contain both amino and carboxyl groups, such as chitosan, polyamino acids, and proteins. In conclusion, these dualcrosslinked gelatin-based hydrogels showing increased mechanical strength, biodegradability, and stability in simulated body fluid have a great potential for use as injectable materials in tissue regenerative medicine and various other biomedical applications with good biocompatibility. We will study further with this hydrogel platform to be used as adhesive material and dressings for wound healing.
Acknowledgements This work was supported by the National Research Foundation of Korea(NRF) grant funded by the Korea government(MSIP)(NRF-2015R1A2A1A14027221) and the Materials and Components Technology Development Program (Strategic Core Material
Technology
Development
Program)
of
MOTIE/KEIT.
[10053595,
Development of functionalized hydrogel scaffold based on medical grade biomaterials with 30% or less of molecular weight reduction]
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Figure captions Figure 1. Synthetic route of GMD. Figure 2. Schematic representation of in situ dual-crosslinked GMD composite hydrogels via Dopa-Fe3+ complexation and Fenton reaction. Figure 3. Preparation of GMD hydrogels via reaction between H2O2 and FeCl2. Figure 4. (a) 1H NMR and (b) UV-Vis spectra of GMD hydrogels. Figure 5. (a) Mechanical strength of hydrogels by varying H2O2 concentration and (b) FeCl2 concentration dependent elastic modulus (G’) of GMD hydrogels. Figure 6. Swelling ratio of hydrogels as a function of time. Figure 7. In vitro proteolytic degradation of hydrogels (n = 3). Figure 8. Stability test of GMD, GDA, and GMA hydrogels. Figure 9. Adhesive strength of hydrogels at 0.05% H2O2 concentration. Figure 10. Cell viability test of GMD hydrogels by varying H2O2 concentrations (n = 3).
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Figure 1.
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Figure 2.
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Figure 4.
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Figure 5.
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Figure 6.
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Figure 7.
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Figure 8.
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Figure 9.
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Figure 10.
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Table 1. Synthetic conditions of GMD conjugates and their dopamine content.
Sample
GMA (g)
Dopamine (mmol)
EDC (mmol)
NHS (mmol)
GMD #1
2.0
13
15
20.8
Dopamine content (μmol g−1 of polymer) 132.4±10.1
GMD #2
2.0
8
10
13.8
98.2±5.3
GMD #3
2.0
3
5
6.9
57.2±4.6
28
S1226-086X(17)30487-2 https://doi.org/10.1016/j.jiec.2017.09.014 JIEC 3617
To appear in: Received date: Revised date: Accepted date:
22-6-2017 25-8-2017 8-9-2017
Please cite this article as: Ji Yeon Kim, Seung Bae Ryu, Ki Dong Park, Preparation and Characterization of Dual-crosslinked Gelatin Hydrogel via DopaFe3+ Complexation and Fenton reaction, Journal of Industrial and Engineering Chemistry https://doi.org/10.1016/j.jiec.2017.09.014 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Preparation and Characterization of Dual-crosslinked Gelatin Hydrogel via Dopa-Fe3+ Complexation and Fenton reaction
Ji Yeon Kim*, Seung Bae Ryu* and Ki Dong Park#
Department of Molecular Science and Technology, Ajou University, Suwon, 443-749, Republic of Korea
#Corresponding
author: Department of Molecular Science and Technology, Ajou University, 5
Woncheon, Yeongtong, Suwon, 443-749, Republic of Korea; Fax: +82 31 219 1592; Tel: +82 31 219 1846; E-mail address: [email protected]
* These authors contributed equally to this work.
Graphical abstract
1
Abstract Gelatin hydrogel via Dopa-Fe3+ complexation and Fenton reaction was developed as a dual-crosslinked hydrogel fabrication method for biomedical application. In order to prepare these hydrogels, methacrylated gelatin dopamine (GMD) conjugates were synthesized. The hydrogels were formed rapidly by using hydrogen peroxide (H2O2) and iron(II) chloride (FeCl2), the acryl moieties at the gelatin conjugate react via Fenton reaction and created Fe3+ ions as a byproduct. These Fe3+ ions were used to form Dopa-Fe3+ complexes. The mechanical strength could be controlled by varying the H2O2 concentration and exhibited 2–3 times higher than that single-crosslinked hydrogels. Furthermore, the swelling ratio and degradation behavior of the hydrogels were controllable depending on the crosslinking density. It was also found that the GMD hydrogels have tissue adhesive ability because of the dopamine content. The developed GMD hydrogels without oxidizing agents under mild conditions can be used as injectable materials in tissue regenerative medicine and various biomedical applications.
Keywords: dopamine, Fenton reaction, dual crosslinking, gelatin, Dopa-Fe3+ coordination complex
1. Introduction Biomedical carriers have been widely used to deliver drugs with sustained and controlled release at the target cells or tissues. Several novel drug carriers have been employed in the biomedical field, including liposomes, micelles, and nanoparticles [1, 2]. Compared to these known carriers, injectable hydrogels have attracted significant
2
attention in recent years. They also exhibit several advantages over traditional carriers. First, hydrogels with 3D networks can load large amounts of drugs and release drugs in a sustained and controlled manner, which can lead to better therapeutic results [3, 4]. Second, injectable hydrogels can release drugs directly to the target cells or tissues, avoiding circulation of the drug in the entire body [3]. In this manner, they can quickly deliver a large dose of drugs [5]. Moreover, for cancer patients whose physical conditions are not suitable for intensive surgery, the use of injectable hydrogels can significantly reduce treatment risks as their application is minimally invasive operation. In order to prepare an injectable hydrogel, various methods such as enzyme-catalyzed reactions, Schiff base reactions, and photo-crosslinking reactions have been employed [6-8]. From these, Fenton reaction induced hydrogel crosslinking using ferrous salt (Fe2+) and hydrogen peroxide (H2O2) that generate hydroxyl radical could initiating polymerization of vinyl monomers. This reaction have some advantages such as rapid crosslinking time, easy drug/cell loading, and prolong the drug release profiles compared with photo-crosslinked hydrogels[9-11] Among these injectable hydrogels, dopamine-conjugated hydrogels offer a biomimetic approach toward efficient and safe drug delivery [12]. The dopamine moiety in the hydrogel helps it to attach itself to various organic and inorganic surfaces, even in wet conditions [13-16]. Hydrogels with catechol-containing polymers can be formed by adding oxidizing agents such as periodate ([IO]4−) [17, 18] and metal (e.g., Fe(III)) [1923] ions, which can provide a highly biocompatible 3D matrix. The organic oxidizing agent [IO]4− helps to form covalently cross-linked (i.e., irreversible) hydrogels, although the use of an excess amount of [IO]4− may result in toxicity. A transition-metal ion such as Fe(III) assists in the formation of Dopa-Fe3+ coordination-based (i.e., reversible) hydrogels. However, the stoichiometric ratio of Dopa-Fe3+ coordination is 1:3 when the catechol moiety is exposed to relatively high pH values (>8). When the pH value is decreased, the tri-coordinated Dopa-Fe3+ complex transforms to a mono- or bicoordinated complex, resulting in either weak mechanical strength or no gelation.
3
In this study, methacrylated gelatin dopamine (GMD) hydrogels were prepared that were dually cross-linked via Dopa-Fe3+ complexation and Fenton reaction under mild conditions. Such crosslinks improved the mechanical strength and stability of the dopamine-conjugated hydrogels. The Fenton reaction typically relies on the reduction of hydrogen peroxide at the expense of Fe2+ ions and the Fenton reagent has been widely used as a radical initiator in vinylic polymerization or grafting. The chemical structures of the GMD conjugates were characterized by
1
H NMR and UV
spectroscopies. The physicochemical properties such as mechanical strength, adhesive ability, and degradation behavior were evaluated for the hydrogels formed by using different concentrations of H2O2 in the Fenton cross-linking process. In addition, the stability of the hydrogels in a simulated body fluid (SBF) was also investigated in order to understand the effect of body fluids on the hydrogels.
2. EXPERIMENTAL 2.1 Materials Gelatin (type A from porcine skin, >300 bloom), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), N-hydroxysuccinimide (NHS), dopamine hydrochloride (DA), methacrylic anhydride (MA), iron(II) chloride (FeCl2), hydrogen peroxide (H2O2), and collagenase type II were obtained from Sigma Aldrich (St. Louis, MO, USA). Fetal bovine serum (FBS) was purchased from Wisent (Saint-Bruno, QC, Canada) and hydrochloric acid, 35.0–37.0% (HCl) was obtained from Samchun (Seoul, Korea). Human dermal fibroblasts (DFBs) were supplied from Lonza Inc. (Walkersville, MD, USA). Dulbecco's modified Eagle medium (DMEM), fetal bovine serum (FBS), penicillin–streptomycin (P–S), trypsin–EDTA, and Dulbecco's phosphate buffered saline (DPBS) were purchased from Gibco BRL (Grand Island, NY, USA). EZ-Cytox Enhanced Cell Viability Assay Kit was purchased from ITSbio Inc. (Seoul, Korea).
2.2 Synthesis and characterization of methacrylated gelatin dopamine conjugates
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GMD conjugates were synthesized by using MA and DA, as shown in Figure 1. First, methacrylated gelatin (GMA) was synthesized by dissolving 5 g of gelatin in 100 mL of deionized water at 60 °C. Then, MA (0.625 μL) was added to the gelatin solution and allowed to react for 1 h. Finally, 400 mL of warm (40 °C) phosphate buffered saline (PBS, pH 7.4) was added to stop the reaction. The mixture was dialyzed against distilled water using 12–14 kDa cutoff dialysis tubing for three days at 40 °C to remove the salts and residual methacrylic acid. The synthesized GMA was lyophilized for 3 days to generate a white porous foam. Second, GMD and gelatin dopamine (GDA) were synthesized by adding dopamine hydrochloride to GMA and gelatin solution, respectively. The detailed reaction conditions are provided in Table 1. Briefly, 2 g of GMA and gelatin were each dissolved in 100 mL of PBS solution. Next, EDC, NHS, and dopamine hydrochloride were added and the pH of the mixtures was adjusted to 5.0. The reactions were left to stir for 7 h at 40 °C. Subsequently, the resulting solutions were each transferred into a dialysis bag (MWCO = 3.5 kDa) and dialyzed against deionized water at pH ~4 for 2 days and at pH = 7 for 1 day, filtered, and lyophilized to obtain the sponge-type GMD and GDA conjugates. The chemical structure of GMD, GDA, and GMA conjugates were characterized by 1H NMR spectroscopy (AS400, OXFORD instruments, UK). The dopamine contents of
the conjugates were measured quantitatively using the solution of GMD in distilled water at 1mg/mL and dopamine hydrochloride (0.01 - 0.1 mg/mL) as a standard curve at 280 nm with a UV visible spectrophotometer in agreement with the previously reported [24] .(V-750 UV/Vis/NIR, Jasco, Japan).
2.3 Preparation of hydrogels The GMD, GDA, and GMA hydrogels were formed by mixing PBS (pH 7.4) solution with FeCl2 and H2O2 (Figure 2). The GMD, GDA, and GMA polymer solutions (13.5 wt%) in PBS were divided into two aliquots, each containing either FeCl2 (2.5 – 10 mM)
5
or H2O2 (0.005 – 0.05%). The volume ratio of polymer solution : FeCl2 or H2O2 was 9 : 1. In order to obtain hydrogels of these polymers, equal volumes from the two aliquots were mixed and shaken (Figure 3). The pH of all solutions used to fabricate hydrogels were 7.4.
2.4. Rheological measurement The elastic modulus (G') of the hydrogels were measured by using a rheometer (Advanced Rheometer GEM-150-050, Bohlin 80 Instruments, USA) in oscillation mode. 150 μL of each GMD, GMA, GDA polymer solutions containing either FeCl2 or H2O2 were mixed together and applied to the bottom plate by varying H2O2 concentration. The FeCl2 concentration dependent stiffness of GMD hydrogels were assessed as the same procedure. The mechanical strength of the formed hydrogel was then recorded for 10 min at 37 °C with frequency of 0.1 Hz and a strain of 0.01% using a parallel plate geometry (diameter = 25 mm, gap = 0.5 mm)
2.5. Swelling ratio Three replicas of each dried hydrogel were swollen in PBS at 37 °C for 6 h to achieve equilibrium swelling. The degree of swelling of the hydrogels was measured until 72 h. The standard deviations are marked as error bars in the swelling profile charts. The degree of swelling was calculated as follows: Degree of swelling = [(Wet weight – Initial weight) / Initial weight] ×100%
2.6 In vitro proteolytic degradation A total of 300 µL of GMD, GDA, and GMA hydrogels with different elastic modulus were prepared. After equilibrium for 30 min, the hydrogels were incubated at 37 °C in 1 mL of PBS (0.01 M, pH 7.4) with and without 0.005 mg mL−1 of collagenase. At predetermined time intervals, the media were removed and changes in the weight of
6
the hydrogels were recorded. Fresh media were then added for the next time interval. The degree of degradation was calculated using the following equation: Weight of hydrogels (%) = (Wd / Wi) x 100 where Wi and Wd refer to the initial weight and weight of the degraded hydrogels, respectively.
2.7. Stability of hydrogels The GMD, GDA, and GMA hydrogels (50 µL) were fabricated in 1.5 mL microtubes, as described above. The GMD, GDA, and GMA polymer solutions (13.5 wt%) in PBS were divided into two aliquots, each containing either FeCl2 (10 mM) or H2O2 (0.05%) as follow mechanical strength test. Subsequently, 0.8 mL of simulated body fluid (SBF) supplemented with 0.5% (v/v) fetal bovine serum was added into each tube. The changes in the appearance of each sample were recorded as a function of incubation time at 37 °C.
2.8. Tissue adhesive strength Tissue adhesive strength of the hydrogels was investigated by using a universal testing machine (UTM; Instron 3343, Instron, USA) in accordance with the modified ASTM method F2255-05. Pre-dehydrated porcine skins (diameter 25 mm, in 0.01 M PBS, pH 7.4 for 1 h) were attached to the rubber in the testing machine by using cyanoacrylate glue for 1 h. The hydrogels formed with the FeCl2 (10 mM) and H2O2 solutions (10 µL, 0.005–0.05%) were placed between the substrates and cured for 15 s with 1N pressure. The adhesive strength was measured by equipping the UTM with a load cell of 100 N at a crosshead of 10 mm min−1. Each measurement was repeated five times.
2.9. Cytotoxicity test of hydrogels Cytotoxicity of GMD hydrogels against human dermal fibroblasts (hDFBs) were evaluated with indirect contact method, according to ISO10993 standard test.[25].
7
Briefly, the hDFBs were cultured with DMEM containing 10% FBS, 1% P-S on the tissue culture plate (TCPs) at 37 °C and 5% CO2. The hydrogels were prepared (200 µL) with 1 mL of culture medium and kept at 37 °C and 5% CO2 for 24h to fully extract the unreacted residues. The hDFBs were seeded on the 48-well plate (2 ×105 cells / well) and cultured with extracted media for 24 h and the WST-1 assay is carried out. The cells were incubated at 37 °C and 5% CO2 with 1 mL of culture medium containing WST-1 reagents for 2 h. The absorbance of samples were measured at 450 nm using a microplate spectrophotometer (VersaMax tunable microplate reader; Molecular Devices, Sunnyvale, CA, USA). The cell viability of hydrogel extracted media was calculated using the following equation Cell viability (%) = (O.D. of the sample / O.D. of the control) x 100 All experiments were performed in triplicate.
3. RESULTS AND DISCUSSION 3.1. Synthesis and characterization of polymers First, the GMA polymers were synthesized by using MA to enable Fenton reaction as the first crosslinking strategy. The fraction of lysine groups reacted was modified by altering the concentration of MA. GMA was synthesized with 53.8±0.5% of methacrylation degree corresponding to 1.25% volume percentage of MA [26]. Next, the GMD and GDA conjugates were synthesized by grafting dopamine onto GMA and gelatin backbone via EDC/NHS chemistry, which served to enable Dopa-Fe3+ complexation as the second crosslinking strategy. The GMD, GMA, and gelatin conjugates were characterized by
1
H NMR spectroscopy (AS400, OXFORD
instruments, UK). The peaks observed in the range 6.7–6.9 ppm were attributed to the aromatic protons of dopamine. Compared to unmodified gelatin, a new signal was observed in spectra of the GMD and GMA conjugates at 5.3–5.6 ppm and 1.8 ppm, which was assigned to the methacrylate group (Figure 4a) [25]. Compared with gelatin solution, the enhanced absorption of UV light at λ = 280 nm is observed in the solution
8
of the resultant GMA–dopamine conjugate and gelatin-dopamine conjugate, confirming the successful conjugation of dopamine (namely, catechol groups) (Figure 4b) [27]. Moreover, no peaks appear at λ = 395 nm, which indicate that the catechol groups in the gelatin–dopamine conjugate are not oxidized to quinones during the synthetic process [28]. It was found that the dopamine content of these gelatin derivatives reached its maximum amount (142.4 µmol g−1 of GMD and GDA) by taking advantage of the available functional group (–NH2) on GMA and gelatin. The two gelatin derivatives GMD and GDA, which represent the highest dopamine contents, were chosen for further study.
3.2. Rheological analysis of gelatin derivatives hydrogels The elastic modulus (G') against the shear strain is an indication of cohesive strength and was determined by varying either the concentrations of H2O2 or FeCl2 concentration at fixed FeCl2 (10 mM) or H2O2 (0.05%). The fixed concentration of FeCl2 (10 mM) and H2O2 (0.05%) were selected by measuring rheological analysis, in which mechanical strength of hydrogels were highest at each condition. The increasing concentration of H2O2 led to higher G' values because of increased cross-linking density of the hydrogels. The elastic modulus of the GMD hydrogels was found to be higher than that of GDA and GMA hydrogels at the same H2O2 concentration (Figure 5a). These results can be explained that the concentration of H2O2 increases, the formation of OH radicals and Fe3+ ions also increases that lead more crosslinking as both OH radicals and Fe3+ ions are involved in the dual-crosslinking of GMD hydrogels. As shown Figure 5b, FeCl2 concentration dependent mechanical strength increased with increasing FeCl2 concentration but decreased at more than 10 mM. This may due to the binding form of Fe-Dopa complexation is converted from the tri- and bis-form to the mono-form by increasing Fe ion concentration [29].
3.3 Swelling ratio of hydrogels on the crosslinking density
9
Swelling ratio has been widely used as a simple method to characterize water absorption and the stability of biomaterials [30]. The swelling behavior of GMD, GDA, and GMA hydrogels in PBS as a function of time is depicted in Figure 6. It was found that most of the hydrogels attained equilibrium swelling in 12 h. After 12 h of swelling in PBS, the swelling ratio of the GMD hydrogel, which had high crosslinking density, reached 158±12%, while that of the GDA hydrogel having low crosslinking density was 280±14%. These results imply that the formation of a weak network makes these hydrogels porous and the strength of the elastic restoring forces is dependent on the extent of interconnectedness (crosslinking density) of the polymer network.
3.4. Proteolytic degradation behavior The gelatin backbone can be degraded by the action of proteases and its degradability is a very important factor in in situ forming hydrogels during interaction with the tissues in the body. Enzymatic degradation of the GMD, GDA, and GMA hydrogels was investigated in the presence of collagenase, a member of the matrix metalloproteinase family that degrades extracellular matrix components and enhances in vivo cell migration and growth. Figure 7 shows the in vitro proteolytic degradation behavior of the synthesized hydrogels. The GMD hydrogel incubated in PBS without collagenase remained stable for six days. In the presence of collagenase, the GMD hydrogel was completely degraded after six days. However, the GDA and GMA hydrogels were completely degraded within three and five days, respectively, because of the decomposition of the peptide chains by collagenase. Therefore, it was concluded that the degradation rates of the hydrogels depended on their cross-linking density because of limited accessibility of collagenase to the cleavage sites in crosslinked hydrogels. These results confirm that the dual-crosslinked hydrogels synthesized in this work can also be used as a biomimetic scaffold and their degradation can be controlled.
10
3.5. Stability test in the simulated body fluid As internal human organs/tissues are always wet with body fluid, the stability of the hydrogels in simulated body fluid (SBF) solution containing 0.5% (v/v) of FBS was investigated. As seen in Figure 8, the freshly prepared hydrogels at the bottom of 1.5 mL microtubes were brown in color and the SBF solution was nearly colorless. After immersion for 6 h, all hydrogels were intact. Interestingly, the GDA hydrogel was completely dissolved after incubation for 24 h and at the same time, the color of the SBF solution became deeper compared with that after immersion for 6 h. These observations suggest that the Fe3+ ions are gradually lost from the gel system into the SBF solution containing FBS. The dissolution of GDA hydrogels might be attributed to Fe3+ binding to serum proteins in the SBF solution. The serum transferrin present in body fluids is known to strongly bind to Fe3+ ions [31]. Therefore, albeit with rapid gelation, the GDA hydrogels cannot serve as an ideal tissue adhesive owing to lack of long-term effectiveness under in vivo conditions. Actually, the natural MAPs solution which is secreted from a mussel, is also solidified through intermolecular covalent crosslinking after oxidation of its catechol groups to quinones takes place, catalyzed by the polyphenol oxidase enzyme [28, 32]. By learning from mussels, GMD can be further developed by adding a stable, second crosslinking moiety, namely, methacrylated gelatin–dopamine chains achieved from the Fenton reaction.
3.6. Effect of dopamine on the adhesive strength of hydrogels The adhesive properties of the GMD and GDA hydrogels having various mechanical strengths were measured by lap-shear test that assesses the adhesive strength of different materials to biological tissues (Figure 9). Considering the biological similarity to human tissues, fresh porcine skin was chosen as an adherend to closely mimic clinical conditions. The tissue adhesive strength of GMD hydrogels (15 kPa) was similar to that of GDA hydrogels at 0.05% of H2O2, although the mechanical strength of the latter was lower than that of the former. While
11
the GMA hydrogels exhibited low tissue adhesion, the adhesive strength of GMD hydrogels was found to be 2-fold higher than that of the commercially available fibrin glue. These results can be attributed to the strong non-covalent binding interactions generated by the dopamine moieties to the tissue. Furthermore, the dopaminecontaining biopolymer is crosslinked by various means, among which a natural way involves the conversion of catechol groups into quinones by polyphenol oxidase enzyme in mussels, thus achieving covalent self-crosslinking.
3.7 Cytotoxicity test of the GMD hydrogels To assess the in vitro cytotoxicity of GMD hydrogels with different H2O2 concentrations (0.01 – 0.05%), we used the modified ISO 10993 test to the viability of hDFBs cultured in media extracted from hydrogels. The hydrogel extracts were obtained after 24 h of incubation with GMD hydrogels (13.5%) formed by FeCl2 (10 mM) and H2O2 (0.01%, 0.03%, and 0.05%). Culture without hydrogels were used as a control. As shown in Figure 10, the WST-1 assay shown the quantitatively analyze the cell viability. We observed no toxic effects on hDFBs, and overall metabolic rates of the cells exposed to the extracts from the hydrogels were comparable with control.
4. CONCLUSIONS In this study, a dual-crosslinked gelatin hydrogel was developed by combining DopaFe3+ complexation and Fenton reaction. The hydrogel was formed rapidly in the presence of H2O2 and FeCl2. Its physicochemical properties could be easily controlled by varying the concentrations of either H2O2 or FeCl2. It was found that the mechanical strength of the hydrogel ranged from approximately 100 to over 5,000 Pa, which is 2fold and 4-fold higher than that of GMA and GDA (single crosslink), respectively. Thus, the GMD hydrogels were more stable than GMA and GDA in the collagenase solutions. The introduction of dopamine moieties into the gelatin hydrogels led to their 4-fold and 2-fold higher adhesive strength as compared to methacrylated gelatin and fibrin glue,
12
respectively. The dual crosslinking strategy is expected to be extensively employed to fabricate biomedical devices with polymers that contain both amino and carboxyl groups, such as chitosan, polyamino acids, and proteins. In conclusion, these dualcrosslinked gelatin-based hydrogels showing increased mechanical strength, biodegradability, and stability in simulated body fluid have a great potential for use as injectable materials in tissue regenerative medicine and various other biomedical applications with good biocompatibility. We will study further with this hydrogel platform to be used as adhesive material and dressings for wound healing.
Acknowledgements This work was supported by the National Research Foundation of Korea(NRF) grant funded by the Korea government(MSIP)(NRF-2015R1A2A1A14027221) and the Materials and Components Technology Development Program (Strategic Core Material
Technology
Development
Program)
of
MOTIE/KEIT.
[10053595,
Development of functionalized hydrogel scaffold based on medical grade biomaterials with 30% or less of molecular weight reduction]
13
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Figure captions Figure 1. Synthetic route of GMD. Figure 2. Schematic representation of in situ dual-crosslinked GMD composite hydrogels via Dopa-Fe3+ complexation and Fenton reaction. Figure 3. Preparation of GMD hydrogels via reaction between H2O2 and FeCl2. Figure 4. (a) 1H NMR and (b) UV-Vis spectra of GMD hydrogels. Figure 5. (a) Mechanical strength of hydrogels by varying H2O2 concentration and (b) FeCl2 concentration dependent elastic modulus (G’) of GMD hydrogels. Figure 6. Swelling ratio of hydrogels as a function of time. Figure 7. In vitro proteolytic degradation of hydrogels (n = 3). Figure 8. Stability test of GMD, GDA, and GMA hydrogels. Figure 9. Adhesive strength of hydrogels at 0.05% H2O2 concentration. Figure 10. Cell viability test of GMD hydrogels by varying H2O2 concentrations (n = 3).
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Figure 1.
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Figure 2.
19
Figure 3.
20
Figure 4.
21
Figure 5.
22
Figure 6.
23
Figure 7.
24
Figure 8.
25
Figure 9.
26
Figure 10.
27
Table 1. Synthetic conditions of GMD conjugates and their dopamine content.
Sample
GMA (g)
Dopamine (mmol)
EDC (mmol)
NHS (mmol)
GMD #1
2.0
13
15
20.8
Dopamine content (μmol g−1 of polymer) 132.4±10.1
GMD #2
2.0
8
10
13.8
98.2±5.3
GMD #3
2.0
3
5
6.9
57.2±4.6
28