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carboxymethyl chitosan hydrogel with tunable biomechanical properties has application potential as cartilage scaffold
International Journal of Biological Macromolecules 137 (2019) 382–391
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International Journal of Biological Macromolecules journal homepage: http://www.elsevier.com/locate/ijbiomac
Silk fibroin/carboxymethyl chitosan hydrogel with tunable biomechanical properties has application potential as cartilage scaffold Tao Li a,1, Xiongbo Song b,1, Changmei Weng b, Xin Wang a, Liling Gu c, Xiaoyuan Gong a, Quanfang Wei d, Xiaojun Duan a, Liu Yang a,⁎, Cheng Chen a,⁎ a
Center for Joint Surgery, Southwest Hospital, Third Military Medical University (Army Medical University), Chongqing 400038, China Research Institute of Surgery, Daping Hospital, Third Military Medical University (Army Medical University), Chongqing 400038, China Department of Rehabilitation, Guizhou Provincial People's Hospital, Guiyang 550002, China d Biomedical Analysis Center, Third Military Medical University (Army Medical University), Chongqing 400038, China b c
a r t i c l e
i n f o
Article history: Received 22 March 2019 Received in revised form 27 June 2019 Accepted 30 June 2019 Available online 02 July 2019 Keywords: Silk fibroin Carboxymethyl chitosan Hydrogel Cartilage scaffold
a b s t r a c t Tissue engineering is a promising strategy for cartilage repair and regeneration. However, an ideal scaffolding material that not only mimics the biomechanical properties of the native cartilage, but also supports the chondrogenic phenotype of the seeding cells is in need. In this study, we developed a silk fibroin (SF) and carboxymethyl chitosan (CMCS) composite hydrogel with enzymatic cross-links (horseradish peroxidase and hydrogen peroxide) and β-sheet cross-links (ethanol treatment). Results of Fourier transform infrared (FTIR), thermal gravimetric analysis (TGA), and X-ray diffraction (XRD) verified that SF/CMCS composite hydrogels had a tunable β-sheet structure. Therefore, by increasing the time of ethanol treatment from 0 h to 8 h, a series of parameters including pore size (from 50 to 300 μm), equilibrium swelling (from 78.1 ± 2.6% to 91.9 ± 0.9%), degradation (from 100% to 9% reduction in mass over 56 days), rheological properties (storage modulus from 177 Pa to 88,904 Pa), and mechanical properties (compressive modulus from 13 to 829 kPa) of the hydrogels were adjusted. In particular, the material parameters of the hydrogels with 2 h ethanol treatment appeared most suitable for engineered cartilage. Furthermore, the in vitro cellular experiments showed that the hydrogels supported the adhesion, proliferation, glycosaminoglycan synthesis, and chondrogenic phenotype of rabbit articular chondrocytes. Finally, subcutaneous implantation of the hydrogels in mice showed no infections or local inflammatory responses, indicating a good biocompatibility in vivo. In conclusion, the chemical-physical crosslinking SF/CMCS composite hydrogels, with tunable material properties and degradation rate, good biocompatibility, are promising scaffolds for cartilage tissue engineering. © 2019 Published by Elsevier B.V.
1. Introduction Articular cartilage defect, usually caused by trauma, disease, or sports-related injury, results in pain, impaired function, and limited movement of the joint, leading to poor life quality and heavy socioeconomic burden [1,2]. Current clinical treatment methods for cartilage lesion, such as microfracture and autologous chondrocyte implantation have shown much efficacy, but they are restricted by limitations including poor integration to the surrounding tissue, and formation of inferior fibrocartilage [3,4]. In recent years, tissue engineering (TE) has emerged as a promising strategy for cartilage repair and regeneration [5–8]. Silk fibroin (SF), a natural biopolymer extracted from Bombyx mori cocoons, has been ⁎ Corresponding authors at: Center for Joint Surgery, Southwest Hospital, Third Military Medical University, Chongqing 400038, China. E-mail addresses: [email protected] (L. Yang), [email protected] (C. Chen). 1 These authors contributed equally to this work.
https://doi.org/10.1016/j.ijbiomac.2019.06.245 0141-8130/© 2019 Published by Elsevier B.V.
widely used in TE due to its good biocompatibility, excellent mechanical strength, and slow degradation [9–11]. These properties also make SF a suitable scaffold in cartilage repair [12]. However, in SF based hydrogels, the SF molecular chains tend to form large β-sheet aggregates via conformation transition from random coil to β-sheet structure, which affects the mechanical properties of the hydrogels during use [13]. Considering that SF protein contains some reactive amino acid residues of tyrosine (5%), SF hydrogels with chemical cross-links between dityrosine residues can be prepared by reactions with horseradish peroxidase (HRP) and hydrogen peroxide (H2O2) [14]. However, the conformation transition of SF chains from random coil to β-sheet can still be observed in such hydrogels after 4 weeks in vitro [11]. Recently, Numata et al. and Su et al. developed SF hydrogels by inducing the conformation transition from random coil to β-sheet by ethanol treatment in a restricted SF network pre-crosslinked by HRP/H2O2 [15,16]. Although these hydrogels exhibit excellent mechanical properties, they are not suitable for cell growth due to the lack of cell-specific epitopes for SF. Therefore, SF is usually used in combination with other
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biomaterials, such as chitosan, hyaluronic acid, sodium alginate, and pullulan. In this way, the favorable properties of each polymer could be retained, meanwhile the limitations of each component could be remedied [17–20]. Chitosan (CS) is a semisynthetic polymer derived from the partial deacetylation of the chitin, and has excellent biocompatibility and biodegradability [21–23]. Moreover, CS could interact electrostatically with negatively charged glycosaminoglycans (GAGs), and has been shown to promote chondrogenic activity and cartilage-specific protein expression [24,25]. Carboxymethyl chitosan (CMCS) is a derivative of CS, but has a higher water solubility and a better bioactivity than CS [26–29]. The aim of this study was to combine SF and CMCS into a composite hydrogel that has cell adhesion sites and excellent mechanical properties. In addition, treatment time of ethanol to SF/CMCS composite hydrogel was tested to avoid the conformation transition of SF chains from random coil to β-sheet structure of SF chains in hydrogel. In this study, SF/CMCS composite hydrogels were prepared by chemical cross-links with HRP/H2O2, followed by physical cross-links with ethanol treatment of different durations. The physicochemical properties such as morphology, equilibrium swelling, mechanical properties, and degradation of SF/CMCS composite hydrogel were studied. To evaluate the potential of SF/CMCS composite hydrogel in cartilage TE, rabbit articular chondrocytes were cultured within the hydrogels, and cell viability, proliferation, GAGs production, and gene expression profiles were tested. 2. Materials and methods 2.1. Preparation of SF solution SF solutions were prepared according to a previous protocol [11]. Briefly, SF cocoons (kindly provided by Institute of Biotechnology of Southwest University, Chongqing, China) were cut into pieces, degummed in boiling 0.02 M sodium carbonate (Macklin, Shanghai, China) solution for 30 min, washed with deionized (DI) water to remove sericin and wax. The degumming procedure was repeated once, then the degummed SF was dried overnight at 40 °C, and dissolved in 9.3 M lithium bromide (Macklin) at 60 °C for 2 h. The dissolved SF solution was dialyzed with DI water using a 3.5 kD molecular weight cutoff (MWCO) dialysis membrane for 3 days. After dialysis, the solution was centrifuged at 4500g for 20 min under 4 °C to remove the insoluble SF particulates. The concentration of the obtained SF solution was approximately 3.2 wt%, as determined by weighing a dried sample of a known volume. The final concentration of SF was adjusted to 2 wt% with DI water. 2.2. Preparation of SF/CMCS composite hydrogels The tyramine (TA, J&K, Beijing, China) substituted CMCS was synthesized by the coupling reaction of amine groups of TA to carboxylic acid groups of CMCS using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC)/N-hydroxysuccinimide (NHS) activation [20]. Briefly, 1.0 g of CMCS (MW 10 kD, degree of substitution ≥80%, Macklin) was dissolved in 100 mL of 4-Morpholinoethanesulfonic acid buffer (0.1 M, pH 6.0, Macklin). Then, 6 mM of EDC (J&K) and 6 mM of NHS (J&K) were added to the solution. After 30 min stirring, 10 mL of N,N-Dimethylformamide (J&K) solution containing TA (1.5 mM) was added, and the resulting solution was stirred under nitrogen for 3 days at room temperature. The reaction mixture was neutralized with 1 M sodium hydroxide (Macklin) and then purified by dialyzing (MWCO 3500) exhaustively against 0.1 M sodium chloride for 1 day and DI water for another 2 days. The final product was collected after freeze-drying, and stored at 4 °C until use. The percentage of TA that conjugated to the CMCS was determined using UV-spectrophotometer [30,31]. Briefly, the aqueous solution of CMCS-TA was determined at
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the absorbance of 275 nm, and the concentration of the phenolic group in the solution was calculated from a standard curve of TA. SF/CMCS composite hydrogels were prepared by chemical crosslinks with HRP/H2O2, followed by physical cross-links with ethanol treatment as previously described [15]. Briefly, CMCS-TA was dissolved in SF solutions (20 mg/mL) with final CMCS-TA concentration at 2 mg/mL. 10 μL of HRP solution (1000 U/mL) was added to 1 mL of mixed solution of SF and CMCS-TA, then 10 μL of H2O2 (1% v/v) was added and mixed by gentle pipetting. After incubation of the mixture at 37 °C for 1 h, the SF/CMCS composite hydrogels with chemical cross-links were formed. Afterward, the hydrogels were immersed in ethanol (75% v/v) for different time periods (1, 2, 4, or 8 h) to induce the formation of β-sheet structures and physical molecular networks. Then the SF/CMCS composite hydrogels with chemical-physical crosslinks were obtained by washing with DI water to remove ethanol and other residues. The chemically cross-linked hydrogels without ethanol treatment were used as control. Depending on the time of ethanol treatment, the groups of SF/CMCS composite hydrogels were marked as EA0 (control), EA1 (1 h), EA2 (2 h), EA4 (4 h), and EA8 (8 h), respectively. The detailed information of each hydrogel was shown in Table 1.
2.3. Characterization of SF/CMCS composite hydrogels The structure of the hydrogels was analyzed with Fourier transform infrared (FTIR). The hydrogels were quickly frozen in −80 °C followed by freeze-drying, and were cut into slices. The FTIR spectra were obtained in the wavenumber range of 400–4000 cm−1 using FTIR spectrophotometer (PerkinElmer S100, Massachusetts, USA). The thermal properties of the hydrogels were evaluated by thermal gravimetric analysis (TGA). The hydrogels were quickly frozen in −80 °C followed by freeze-drying, and then grinded into powder. TGA was performed on a 1600LF thermal gravimetric/differential scanning calorimetry (TGA/DSC1) instrument (Mettler Toledo, Bern, Switzerland) by heating rate of 2 °C/min under a nitrogen gas flow of 50 mL/min. The crystalline structures of the hydrogels were determined with Xray diffraction (XRD). The sample preparation procedure was the same as for TGA. XRD patterns were obtained on X'Pert Rro MPD powder multi-functional X-ray diffractometer (PANalytical, Almelo, Netherlands), using Cu Kα radiation (40 mA, 40 kV) with a scanning speed of 6°/min. The two-theta range was collected from 5° to 60° and analyzed by JADE6 software. The morphologies of the hydrogels were characterized using scanning electron microscopy (SEM). The hydrogels were pre-frozen at −80 °C, then freeze-dried for 24 h, crosscut with the blade, adhered on the conductive board, and sputter-coated with gold. The morphologies of the surface and section were viewed using Hitachi S3400NIISEM (Hitachi, Tokyo, Japan) operated at 2 kV accelerating voltage.
Table 1 contains detailed information of EA0-EA8 hydrogels including the concentrations of SF and CMCS-TA, the contents of HRP and H2O2 for chemical cross-links, the ethanol treatment time for physical cross-links, the water contents and compressive modulus for each hydrogel. Name SF (mg/mL) CMCS-TA (mg/mL) HRP (1000 U/mL, % v/v) H2O2 (1% v/v, % v/v) Ethanol treatment time (75% v/v, h) Water (wt%) Compressive modulus (kPa)
EA0
EA1
EA2
EA4
EA8
20 2 1 1 0
20 2 1 1 1
20 2 1 1 2
20 2 1 1 4
20 2 1 1 8
91.9 ± 1.5 13.3 ± 4.0
85.6 ± 1.5 59.5 ± 20.6
83.3 ± 2.0 430.0 ± 62.0
77.7 ± 2.7 491.8 ± 36.2
78.1 ± 4.5 828.8 ± 145.2
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2.4. Properties of SF/CMCS composite hydrogels
2.7. GAG quantification
The swelling ratios of the hydrogels were measured in DI water using gravimetric method. The freeze-dried hydrogels (weighed as W1) were immersed in DI water, and kept at 37 °C for 2 days until equilibrium of swelling was reached. The weight of fully swollen hydrogels was recorded immediately after the excess of water lying on the surfaces was absorbed with filter paper, and weighed as W2. The equilibrium water uptake was calculated using the equation: (W2 − W1) / W2 × 100%. The in vitro enzymatic degradation of the hydrogels was measured gravimetrically at 37 °C. The hydrogels (weighed as W0) were immersed in 2 mL PBS with 0.2 units of protease (Type XIV). Enzyme solution was replaced every 3 days for a period of 56 days. Sodium azide (0.05 w/v%) was added to the solution to prevent microbial growth. At specified time points, the hydrogels were removed and weighed (Wt). Samples without the protease were taken as control. The weight remaining ratio was defined as [1 − (W0 − Wt) / W0] × 100%. The rheological properties of the hydrogels were performed with DHR-1 rheometer (TA Instruments, Delaware, USA) by using parallel plates (diameter of plates: 20 mm, gap between plates: 500 nm) in the oscillatory mode at 37 °C. The frequency sweeps were carried out in the range of 0.1–10 Hz at a constant strain of 1%. The mechanical properties of the hydrogels with known sizes (height and diameter) were measured using Instron 5969 testing frame (Instron Instruments, Massachusetts, USA). The compression test was conducted with a pre-strain force of 0.06 N load cell at room temperature. Then the constant strain force was maintained at 2 mm/min till the rupture of hydrogels. The compressive modulus was obtained from the linear elastic region of the stress-strain curves [32].
GAG content of the chondrocyte-seeded hydrogels was evaluated using Blyscan Sulfated Glycosaminoglycan Assay (Biocolor, Antrim, UK) according to the manufacturer's instructions. Cell-seeded hydrogels at 7 and 14 days were digested by papain (Sigma) solution at 60 °C for 6 h. 200 μL of dye solution was transferred to individual wells of a 96well plate, and the absorbance at 656 nm was measured using a microplate reader (Thermo Scientific, Massachusetts, USA). The GAG content was normalized to the dry weight of the remained gel and to DNA content of each sample, respectively.
2.5. Chondrocyte isolation and culture Chondrocytes were harvested from the femoral condyles of 4 week old New Zealand rabbits in accordance with the guidelines approved by the Institutional Animal Care and Use Committee of Third Military Medical University (Chongqing, China). The articular cartilages were cut into slices and digested by 2 mg/mL of type II collagenase (Sigma, Missouri, USA) in DMEM/F12 (Gibco, New York, USA) at 37 °C in a 5% CO2 incubator for 12 h. After that, cells were collected from the digestion solution through a 40 μm cell strainer (BD Bioscience, New Jersey, USA), centrifuged at 500g for 5 min. Then, the collected cells were resuspended in complete medium comprised of DMEM/F12, 10 v/v% fetal bovine serum (FBS, Gibco), and 1 v/v% of penicillin-streptomycin (Gibco). Culture medium was changed every 3 days. The fabricated hydrogels (8 mm in diameter and 3 mm in height) were preconditioned with complete medium in a 24-well plate at 37 °C for 6 h before cell seeding. Chondrocytes were seeded on each hydrogel at a density of 1 × 104 cells/well. After 4 h, the hydrogels were transferred to another 24-well plate containing chondrogenic medium. The medium was changed every 2 days.
2.8. Gene expression and histological analysis First, total RNA was extracted using Trizol (Roche, Basel, Switzerland). Subsequently, RNA was reverse transcribed into cDNA using the Transcriptor cDNA Synth. kit 2 (Roche). Real-time polymerase chain reaction (RT-PCR) was carried out using FS Essential DNA Green Master (Roche). Collagen type II (COL2A1), aggrecan (ACAN), collagen type I (COL1A1), collagen type X (COL10A1), and SOX9 genes were quantified. The results were analyzed according to the 2−ΔΔCT method and normalized to the housekeeping gene GAPDH. The target gene primers (Sangon Biotech, Shanghai, China) were designed as follows: COL2A1, forward, CAACAACCAGATCGAGAGCA, reverse, CCAGTAGTCAC CGCTCTTCC; ACAN, forward, CGACATCAGTGGAGACCTCA, reverse, AGACTGCCAGAGTCCAGCTC; COL1A1, forward, CAATCACGCCTCTCAG AACA, reverse, TCGGCAACAAGTTCAACATC; COL10A1, forward, GAAT GGCACGCCTGTAATGT, reverse, CCATTGGACTCAGCGTTAGG; SOX9, forward, GGAAGCTCTGGAGACTGCTG, reverse, GCGGCTGGTACTTGTAGTC C; GAPDH, forward, CGACATCAAGAAGGTGGTGA, reverse, ATCGAA GGTGGAGGAGTGG. The chondrocyte-seeded hydrogels were fixed in 4% paraformaldehyde for 2 h at room temperature. For immunohistochemistry, the Anti-collagen type I and Anti-collagen type II (Abcam) primary antibodies were employed under 4 °C for overnight. Then goat anti-rabbit secondary antibodies were used. The images were observed using microscope (Olympus BX50, Japan). 2.9. Subcutaneous implantation The in vivo biological response of the SF/CMCS composite hydrogels was evaluated by subcutaneous implantation. The Kunming mice (weight 20–30 g) were anesthetized with 1 w/v% of pentobarbital (50 mg/kg) by intraperitoneal injection. Then, the SF/CMCS hydrogels (6 mm in diameter and 4 mm in height) were then implanted into lateral incisions on the dorsal region. The mice were euthanized after implantation for 2 and 4 weeks. The specimens along with the adjacent tissues were fixed in 4% paraformaldehyde at 4 °C overnight, dehydrated in sucrose, embedded in frozen section media, and frozen sectioned at a thickness of 10 μm. The histological sections were observed with hematoxylin & eosin (H&E) staining.
2.6. Cell viability and proliferation
2.10. Statistical analysis
Cell viability was determined using a Live/Dead cell double stain kit (Yeasen, Shanghai, China) at 7 and 14 days. In brief, cell-seeded hydrogels were incubated in Assay Buffer containing 2 mM CalceinAM and 1.5 mM propidium iodine for 15 min at 37 °C. Labeled cells were visualized using a fluorescent microscope. Live cells were stained green, whereas dead cells were stained red. Cell proliferation at 7 and 14 days was evaluated by quantifying DNA. DNA content was detected with AccuGreenTM High Sensitivity dsDNA Quantitation Kits (Biotium, California, USA) according to the manufacturer's instructions. DNA content of cell-seeded hydrogels was expressed as DNA amount normalized to the dry weight of the remained gel.
All data were presented as the mean ± standard deviation (SD). Statistical significance was determined by one-way ANOVA followed by Tukey's post hoc analysis (SPSS statistics 23, IBM). p values b 0.05 were considered statistically significant. 3. Results and discussion 3.1. Preparation of SF/CMCS composite hydrogel The synthetic scheme of CMCS-TA was shown in Fig. 1A. CMCS-TA was synthesized by the coupling reaction of the amino group of TA to the carboxylic acid groups of CMCS using EDC/NHS activation. The
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Fig. 1. The scheme of hydrogel formation and characterization of CMCS-TA. (A) CMCS was partially decorated with TA. (B) SF/CMCS composite hydrogel was prepared with chemical crosslinks by HRP/H2O2, followed by physical cross-links with ethanol treatment. (C) 1H NMR spectra of TA, CMCS, and CMCS-TA. (D) UV–Vis absorbance spectrum of H2O, CMCS, CMCS-TA, and TA.
structure of the CMCS-TA conjugates was confirmed by 1H NMR (Fig. 1C). When compared with CMCS spectrum, 1H NMR spectra of CMCS-TA exhibited new signals at δ 6.67 and δ 6.98, which indicated the presence of the aromatic protons of TA. The degree of TA molecules per 100 repeating units of CMCS was about 10.08%, calculated from a standard curve of TA (Fig. 1D). SF/CMCS composite hydrogels were prepared by chemical crosslinks with HRP/H2O2, followed by physical cross-links with ethanol treatment (Fig. 1B). The CMCS-TA polymer was freeze-dried and then dissolved in SF solution. The enzyme-mediated polymerization strategy was applied to form chemical cross-links with HRP and H2O2. SF contains tyrosine residues with a content of approximately 5 mol% [33], therefore the TA groups of CMCS can react with tyrosine residues of SF to form covalent bonds. HRP can facilitate the cross-links between SF and CMCS-TA via forming free radical species in the presence of H2O2 [34]. Then the chemically cross-linked hydrogels were treated with ethanol to induce the β-sheet formation via the hydrophobic interactions of the (GAGAGS)n sequences in relatively smaller network structures of SF [15]. 3.2. Structure, swelling and degradation of SF/CMCS composite hydrogel Structural changes of the SF/CMCS composite hydrogels were detected by FTIR. In general, the structure of SF at 1625 cm−1 and 1515 cm−1 in the FTIR spectrum is recognized as β-sheet conformation, meanwhile the absorption band at 1260 cm−1 inside the amide III region is also associated with the β-sheet conformation [16,35]. According to Fig. 2A, it was found that a new peak appeared at 1625 and 1515 cm−1 after ethanol treatment, which could be attributed to the conformational transition of SF from random coil to β-sheet [36,37]. The absorption band changed very slightly when the ethanol treatment
time increased from 2 h to 8 h, indicating that the conformational transition of SF from random coil to β-sheet had basically completed after 2 h. Then, the conformational changes of the SF/CMCS composite hydrogels were evaluated by DSC (Fig. 2B). Compared with EA0 hydrogel, the degradation peak [38] appeared at 250–280 °C in all the ethanol-treated hydrogels. This result indicated the presence of βsheet structure in ethanol-treated hydrogels. Similarly, there were only slight changes in the degradation peak between EA2, EA4, and EA8 hydrogels, which was consistent with the FTIR results. Moreover, XRD analysis provided a detailed information for the microstructural difference among the ethanol-treated hydrogels. As was shown in Fig. 2C, compared with EA0 hydrogel, the ethanol-treated hydrogels presented a sharper peak around 20°, which is the character peak of β-sheet [16]. When the time of ethanol treatment increased from 2 h to 8 h, there were no obvious changes in XRD curves. Therefore, the XRD results together with FTIR and DSC results, indicated that the SF/ CMCS composite hydrogels possessed a tunable β-sheet structure that had been basically formed after 2 h ethanol treatment. The swelling ability of the hydrogel is used to character its ability to maintain structure and retain enough water. Water absorption and retention is important when hydrogels are used in TE, because it reflects the ability of the hydrogel to transfer oxygen and nutrient [39]. The swelling ratio of the SF/CMCS composite hydrogels was investigated by immersing the freeze-dried hydrogels in PBS to reach equilibrium at 37 °C. As was shown in Fig. 2D and Table 1, all hydrogels showed good swelling ratio, ranging from 78.1 ± 2.6% to 91.9 ± 0.9%. The water uptake ability significantly decreased from 91.9 ± 0.9% to 78.1 ± 2.6%, as the ethanol treatment time increased from 0 to 8 h. These results could be ascribed to the increase of ethanol treatment time that leads to more the formation of β-sheet structure in SF. The recurring hydrophobic regions of alanine and glycine residues in SF form β-sheet
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Fig. 2. The structural analysis, swelling, and degradation of the SF/CMCS composite hydrogels. (A) FTIR, (B) TGA, and (C) XRD curves of the hydrogels. (D) Water uptake, (B) Degradation of the hydrogels in PBS, (C) Degradation of the hydrogels in protease (Type XIV). Data were presented as mean ± SD (n = 3, *p b 0.05, **p b 0.01, ***p b 0.001).
structures [40], and the higher β-sheet content contribute to the higher hydrophobicity. The degradation rate is another key parameter of hydrogels in TE. Here the in vitro degradation was assessed by monitoring the weight loss of the hydrogel post 56 days of incubation with PBS and protease XIV solution, respectively. In Fig. 2E, the remaining masses of EA0 and EA1 hydrogels were 56.6 ± 4.5% and 84.5 ± 1.5% respectively after incubating with PBS up to 56 days. The increase of ethanol treatment time resulted in a significant decrease in degradation rate. However, there was almost no difference between the EA2 (96.9 ± 0.9%), EA4 (98.4 ± 0.1%), and EA8 (97.5 ± 0.8%) hydrogels in degradation, indicating that the β-sheet structure of SF had been basically formed after 2 h ethanol treatment. While in Fig. 2F, the degradation rate of hydrogels
was much faster in the presence of protease XIV. The EA0 and EA1 hydrogels completely degraded after 2 and 10 days in protease XIV, respectively. The degradation of EA2 hydrogel accelerated after 42 days, and eventually EA2 (73.1 ± 5.4%) hydrogel showed 44% reduction in mass over 56 days. However, the EA4 (87.9 ± 1.3%) and EA8 (91.0 ± 1.3%) hydrogels only showed 12% and 9% reduction in mass over 56 days, respectively. The reason is that the hydrogels with higher βsheet content could better protect themselves against enzymatic attack [41]. In cartilage TE, if a scaffold degrades too fast, it might collapse under continual mechanical loading; if a scaffold degrades too slowly, it hinders the growth of the new cartilage tissue. In this study, we showed that the degradation of the SF/CMCS composite hydrogels could be tuned into a proper rate by varying the ethanol treatment time.
Fig. 3. Morphology of the SF/CMCS composite hydrogels. (A) Macroscopic observation of the hydrogels. (B) SEM images of the cross-section of the hydrogels. (C) SEM images of the surface of the hydrogels.
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3.3. Morphologies of SF/CMCS composite hydrogel The macroscopic observation of the SF/CMCS composite hydrogels was shown in Fig. 3A. The hydrogels became darker as ethanol treatment increased from 0 h to 8 h. As was shown in Fig. 3B, the inner morphology of SF/CMCS composite hydrogels was determined by crosssectional SEM images. These images revealed that the composite hydrogels had a pore size range of 50–300 μm with well interconnected three-dimensional porous structure. The EA2, EA4, and EA8 hydrogels had similar pore sizes, because the conformational transition of SF from random coil to β-sheet had basically completed after 2 h ethanol treatment. It has been shown that SF scaffolds with pore sizes of 90–250 μm provide the best environment for adhesion and proliferation of chondrocytes [42]. This pore size range almost overlaps with that of the SF/CMCS composite hydrogels, therefore we expected that the SF/ CMCS composite hydrogels likewise support chondrocyte growth. 3.4. Mechanical properties of SF/CMCS composite hydrogel Generally, the rheological properties are studied via oscillatory rheology to assess the stability of three-dimensional cross-linked networks such as hydrogels. Storage modulus (G′) and loss modulus (G″) represent elastic and viscous behavior of the hydrogels, respectively. As was shown in Fig. 4, all samples exhibited higher G′ values than G″ values in the frequency sweep tests (from 0.1 Hz to 100 Hz), indicating a predominantly elastic behavior of the hydrogels. At a fixed frequency of 1 Hz, the G′ values for EA0, EA1, EA2, EA4, and EA8 hydrogels were 177 Pa, 20,443 Pa, 36,495 Pa, 56,282 Pa, and 88,904 Pa, respectively. There was about a 500-fold increase in G′ values (from 177 Pa to 88,904 Pa) of the hydrogels when the time of ethanol treatment increased from 0 h to 8 h, demonstrating an improvement of the mechanical properties with longer treatment of ethanol. Elasticity of the scaffolding materials has been recognized as a dominating factor of cell fate in TE [43]. Here, the compression test was conducted to verify the elastic properties of the SF/CMCS composite hydrogels. As was shown in Fig. 5A, the stress-at-failure and strain-atfailure values were determined from the resulting compressive stressstrain curves. The ethanol treatment time had a great effect on the stress and strain of these hydrogels at break. Fig. 5B showed that 8 h of ethanol treatment led to a 30-fold increase in compressive modulus (from
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13 kPa of EA0 to 829 kPa of EA8). This was in accordance with a previous study showing that the SF molecular chains could be further immobilized by forming extensive β-sheet structures after ethanol treatment, which in turn greatly improves both strength and elasticity of the SF hydrogel. Note that the compressive modulus of EA2 (430 ± 62 kPa), EA4 (492 ± 36 kPa) and EA8 (829 ± 145 kPa) hydrogels was in the range of compressive moduli of natural cartilage (300–800 kPa) [44]. Therefore, the results of the compression tests together with the rheological tests, suggested that the mechanical performance of the SF/CMCS composite hydrogels could be adjusted to suit cartilage TE. Taken together the parameters of pore size, swelling, degradation, and mechanical performance, we chose EA2 hydrogel as experimental group and EA0 hydrogel as control in the cellular experiments. 3.5. Cell viability, proliferation, and GAG production Fluorescent images of live/dead cell staining were used to investigate the viability of rabbit chondrocytes seeded on the SF/CMCS composite hydrogels (Fig. 6A). The EA0 and EA2 hydrogels showed uniform distribution and displayed a spherical morphology of cells throughout the culture conditions. Few dead cells were observed within the hydrogels. Next, chondrocyte proliferation was monitored by quantifying DNA content of the hydrogels (Fig. 6B). The DNA content of EA2 hydrogel at day 14 was significantly higher than that at day 7. Furthermore, there was no significant difference in DNA content between EA2 and EA0 hydrogel after 7 and 14 days of culture. These results indicated that both the EA2 and EA0 hydrogels had good biocompatibility with chondrocytes. GAG is an important ECM component of cartilage. When the GAG production of the rabbit chondrocytes seeded on the SF/CMCS composite hydrogels was normalized to the dry weight of remained gel (Fig. 6C), there was no statistical difference between EA0 and EA2 hydrogel. In EA2 hydrogel, it was found that the GAG/dry gel ratio had a significant increase at day 14, in comparison to that at day 7. When the GAG production from the cells seeded on the SF/CMCS composite hydrogels was normalized by their respective DNA content (Fig. 6D), the normalized value of EA2 hydrogel was higher that of EA0 hydrogel at both day 7 and day 14, indicating that averagely each single cell in EA2 hydrogel had a better ability to synthesize GAG. Similarly, a study showed that SF scaffolds with β-sheet conformational, derived from HRP-mediated cross-linking SF solution in combination with salt-
Fig. 4. Rheology of the SF/CMCS composite hydrogels.
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A
B EA1
EA2
EA4
EA8
Stress (kPa)
EA0
Strain (%) Fig. 5. Mechanical properties of the SF/CMCS hydrogels. (A) Compressive stress–strain curves, (B) Compressive modulus (n = 3, ***p b 0.001).
leaching and freeze-drying methodologies, facilitated cartilage-specific ECM production as scaffolds for cartilage regeneration [45]. This finding together with our results, suggests that the appropriate β-sheet structures in the SF/CMCS composite hydrogels enhance cartilaginous matrix production. 3.6. Gene expression and immunohistochemical staining To compare the ability of EA0 and EA2 hydrogel in maintaining chondrocyte phenotype, the expression levels of cartilage related genes including chondrogenic markers COL2A1 and ACAN, fibrotic marker COL1A1, hypertrophic marker COL10A1, and early chondrogenic transcription factor SOX9 were quantified by RT-PCR after 7 and 14 days of culture. As was shown in Fig. 7A, the COL2A1 expression of chondrocytes in EA2 hydrogel increased significantly from day 7 to day 14 (55.2 folds); moreover, the chondrocytes in EA2 hydrogel expressed about 22.8 folds of COL2A1 compared to those in EA0 hydrogel at day 14. For ACAN expression (Fig. 7B), there was a higher
EA0
EA2
3.7. In vivo hydrogel biocompatibility The in vivo biological response of SF/CMCS composite hydrogels was evaluated by subcutaneous implantation in a mouse model. The
B
day 14
day 7
A
expression in EA2 hydrogel than in EA0 hydrogel at day 14 (1.7 folds). As for expression of SOX9, COL1A1, and COL10A1 (Fig. 7C, D, and E), there was no difference either between groups (EA0 vs. EA2) or with time (day 7 vs. day 14). The higher expression of chondrogenic markers COL2A1 and ACAN in EA2 group than in EA0 group, indicated that the SF/CMCS composite hydrogels with chemical-physical cross-links had a beneficial impact on chondrogenic phenotype. The immunohistochemical staining of chondrocytes-seeded SF/ CMCS composite hydrogels were shown in Fig. 8. The EA0 and EA2 hydrogels exhibited limited staining for collagen type I (fibrotic marker) at Day 7. Compared to EA0 hydrogels, EA2 hydrogels exhibited a stronger staining of collagen type II, indicating that EA2 hydrogels were more suitable in maintaining chondrocyte phenotype.
C
D
Fig. 6. Cell viability, proliferation, and GAG production on the SF/CMCS composite hydrogels. (A) Live-dead staining (green: live, red: dead). (B) DNA content of cell-seeded hydrogels was expressed as DNA amount normalized to the dry weight of the remained gel after 7 and 14 days of culture. (C) GAG production normalized to the dry weight of the remained gel, (D) GAG production normalized to the DNA content of chondrocytes on SF/CMCS hydrogels after 7 and 14 days of culture (n = 3, *p b 0.05, **p b 0.01). (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
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A
B
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C
COL2A1
SOX9
ACAN
E
D COL1A1
COL10A1
Fig. 7. Gene expression analysis of the rabbit chondrocytes on the SF/CMCS composite hydrogels. (A) COL2A1, (B) ACAN, (C) SOX9, (D) COL1A1, and (E) COL10A1 were normalized to GAPDH and expressed as relative values to day 7. (n = 3, *p b 0.05, ***p b 0.001).
macroscopic images of the explants showed that EA0 and EA1 hydrogels mostly degraded, EA2 hydrogels slightly degraded, while EA4 and EA8 hydrogels hardly degraded (Fig. 9A). This was consistent with the in vitro degradation result, and further confirmed that the SF/CMCS composite hydrogels had tunable degradation through changing ethanol treatment time. The macroscopic images of the explants in situ showed that hydrogels were integrated in the subcutaneous tissue
after 14 and 28 days of implantation (Fig. 9B and C). There was no evident sign of infections or inflammatory responses in the surrounding tissue. The hydrogels and surrounding tissues were collected and stained with H&E for qualitative histological examination. From the H&E staining images (Fig. 9D), it was observed that a thick layer of connective tissue adhered on the entire surface of the hydrogels after 14 and 28 days of implantation. The absence of edema or signs of
Fig. 8. Immunohistochemical staining of chondrocytes-seeded SF/CMCS composite hydrogels at 7 days.
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A EA1
EA4
EA2
EA8
B
EA0
C
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Day 28
Day 14
EA0
EA0
EA1
EA2
EA4
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D
Fig. 9. In vivo degradation and biocompatibility of the SF/CMCS composite hydrogels. (A) Macroscopic images of the hydrogels after 14 days and 28 days of implantation. (B, C) Macroscopic images of the EA0 hydrogel in situ after implantation for 14 days and 28 days. (D) H&E staining of the hydrogel explants after 14 days and 28 days of implantation.
neutrophils after implantation, revealed that no acute inflammation was induced by the implanted biomaterial. These results suggested that the SF/CMCS composite hydrogels exhibited an acceptable tissue compatibility in vivo. 4. Conclusions In this study, SF/CMCS composite hydrogels were prepared by chemical cross-links with HRP/H2O2, followed by physical cross-links with ethanol treatment. Frist, FTIR, DSC, and XRD verified that SF/CMCS composite hydrogels had a tunable β-sheet structure. Next, the physicochemical properties such as morphology, equilibrium swelling, degradation, rheological properties, and mechanical properties of the hydrogels could be adjusted by varying the ethanol treatment time. Then, rabbit chondrocytes were seeded on the selected hydrogels (EA2) for cellular, biochemical, and gene expression assays, and such hydrogels maintained and even promoted chondrogenic phenotype. Last, SF/CMCS composite hydrogels exhibited an excellent tissue compatibility in vivo. Overall, our studies suggest that the chemical-physical cross-linking SF/CMCS composite hydrogels with tunable mechanical properties and biodegradability are a promising scaffolding material for cartilage TE. Declaration of Competing Interest We declare that we have no conflict of interest. Acknowledgements This work was supported by National Natural Science Foundation of China (Grant no. 81572107). References [1] D. Correa, S.A. Lietman, Articular cartilage repair: current needs, methods and research directions, Semin. Cell Dev. Biol. 62 (2017) 67–77.
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Contents lists available at ScienceDirect
International Journal of Biological Macromolecules journal homepage: http://www.elsevier.com/locate/ijbiomac
Silk fibroin/carboxymethyl chitosan hydrogel with tunable biomechanical properties has application potential as cartilage scaffold Tao Li a,1, Xiongbo Song b,1, Changmei Weng b, Xin Wang a, Liling Gu c, Xiaoyuan Gong a, Quanfang Wei d, Xiaojun Duan a, Liu Yang a,⁎, Cheng Chen a,⁎ a
Center for Joint Surgery, Southwest Hospital, Third Military Medical University (Army Medical University), Chongqing 400038, China Research Institute of Surgery, Daping Hospital, Third Military Medical University (Army Medical University), Chongqing 400038, China Department of Rehabilitation, Guizhou Provincial People's Hospital, Guiyang 550002, China d Biomedical Analysis Center, Third Military Medical University (Army Medical University), Chongqing 400038, China b c
a r t i c l e
i n f o
Article history: Received 22 March 2019 Received in revised form 27 June 2019 Accepted 30 June 2019 Available online 02 July 2019 Keywords: Silk fibroin Carboxymethyl chitosan Hydrogel Cartilage scaffold
a b s t r a c t Tissue engineering is a promising strategy for cartilage repair and regeneration. However, an ideal scaffolding material that not only mimics the biomechanical properties of the native cartilage, but also supports the chondrogenic phenotype of the seeding cells is in need. In this study, we developed a silk fibroin (SF) and carboxymethyl chitosan (CMCS) composite hydrogel with enzymatic cross-links (horseradish peroxidase and hydrogen peroxide) and β-sheet cross-links (ethanol treatment). Results of Fourier transform infrared (FTIR), thermal gravimetric analysis (TGA), and X-ray diffraction (XRD) verified that SF/CMCS composite hydrogels had a tunable β-sheet structure. Therefore, by increasing the time of ethanol treatment from 0 h to 8 h, a series of parameters including pore size (from 50 to 300 μm), equilibrium swelling (from 78.1 ± 2.6% to 91.9 ± 0.9%), degradation (from 100% to 9% reduction in mass over 56 days), rheological properties (storage modulus from 177 Pa to 88,904 Pa), and mechanical properties (compressive modulus from 13 to 829 kPa) of the hydrogels were adjusted. In particular, the material parameters of the hydrogels with 2 h ethanol treatment appeared most suitable for engineered cartilage. Furthermore, the in vitro cellular experiments showed that the hydrogels supported the adhesion, proliferation, glycosaminoglycan synthesis, and chondrogenic phenotype of rabbit articular chondrocytes. Finally, subcutaneous implantation of the hydrogels in mice showed no infections or local inflammatory responses, indicating a good biocompatibility in vivo. In conclusion, the chemical-physical crosslinking SF/CMCS composite hydrogels, with tunable material properties and degradation rate, good biocompatibility, are promising scaffolds for cartilage tissue engineering. © 2019 Published by Elsevier B.V.
1. Introduction Articular cartilage defect, usually caused by trauma, disease, or sports-related injury, results in pain, impaired function, and limited movement of the joint, leading to poor life quality and heavy socioeconomic burden [1,2]. Current clinical treatment methods for cartilage lesion, such as microfracture and autologous chondrocyte implantation have shown much efficacy, but they are restricted by limitations including poor integration to the surrounding tissue, and formation of inferior fibrocartilage [3,4]. In recent years, tissue engineering (TE) has emerged as a promising strategy for cartilage repair and regeneration [5–8]. Silk fibroin (SF), a natural biopolymer extracted from Bombyx mori cocoons, has been ⁎ Corresponding authors at: Center for Joint Surgery, Southwest Hospital, Third Military Medical University, Chongqing 400038, China. E-mail addresses: [email protected] (L. Yang), [email protected] (C. Chen). 1 These authors contributed equally to this work.
https://doi.org/10.1016/j.ijbiomac.2019.06.245 0141-8130/© 2019 Published by Elsevier B.V.
widely used in TE due to its good biocompatibility, excellent mechanical strength, and slow degradation [9–11]. These properties also make SF a suitable scaffold in cartilage repair [12]. However, in SF based hydrogels, the SF molecular chains tend to form large β-sheet aggregates via conformation transition from random coil to β-sheet structure, which affects the mechanical properties of the hydrogels during use [13]. Considering that SF protein contains some reactive amino acid residues of tyrosine (5%), SF hydrogels with chemical cross-links between dityrosine residues can be prepared by reactions with horseradish peroxidase (HRP) and hydrogen peroxide (H2O2) [14]. However, the conformation transition of SF chains from random coil to β-sheet can still be observed in such hydrogels after 4 weeks in vitro [11]. Recently, Numata et al. and Su et al. developed SF hydrogels by inducing the conformation transition from random coil to β-sheet by ethanol treatment in a restricted SF network pre-crosslinked by HRP/H2O2 [15,16]. Although these hydrogels exhibit excellent mechanical properties, they are not suitable for cell growth due to the lack of cell-specific epitopes for SF. Therefore, SF is usually used in combination with other
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biomaterials, such as chitosan, hyaluronic acid, sodium alginate, and pullulan. In this way, the favorable properties of each polymer could be retained, meanwhile the limitations of each component could be remedied [17–20]. Chitosan (CS) is a semisynthetic polymer derived from the partial deacetylation of the chitin, and has excellent biocompatibility and biodegradability [21–23]. Moreover, CS could interact electrostatically with negatively charged glycosaminoglycans (GAGs), and has been shown to promote chondrogenic activity and cartilage-specific protein expression [24,25]. Carboxymethyl chitosan (CMCS) is a derivative of CS, but has a higher water solubility and a better bioactivity than CS [26–29]. The aim of this study was to combine SF and CMCS into a composite hydrogel that has cell adhesion sites and excellent mechanical properties. In addition, treatment time of ethanol to SF/CMCS composite hydrogel was tested to avoid the conformation transition of SF chains from random coil to β-sheet structure of SF chains in hydrogel. In this study, SF/CMCS composite hydrogels were prepared by chemical cross-links with HRP/H2O2, followed by physical cross-links with ethanol treatment of different durations. The physicochemical properties such as morphology, equilibrium swelling, mechanical properties, and degradation of SF/CMCS composite hydrogel were studied. To evaluate the potential of SF/CMCS composite hydrogel in cartilage TE, rabbit articular chondrocytes were cultured within the hydrogels, and cell viability, proliferation, GAGs production, and gene expression profiles were tested. 2. Materials and methods 2.1. Preparation of SF solution SF solutions were prepared according to a previous protocol [11]. Briefly, SF cocoons (kindly provided by Institute of Biotechnology of Southwest University, Chongqing, China) were cut into pieces, degummed in boiling 0.02 M sodium carbonate (Macklin, Shanghai, China) solution for 30 min, washed with deionized (DI) water to remove sericin and wax. The degumming procedure was repeated once, then the degummed SF was dried overnight at 40 °C, and dissolved in 9.3 M lithium bromide (Macklin) at 60 °C for 2 h. The dissolved SF solution was dialyzed with DI water using a 3.5 kD molecular weight cutoff (MWCO) dialysis membrane for 3 days. After dialysis, the solution was centrifuged at 4500g for 20 min under 4 °C to remove the insoluble SF particulates. The concentration of the obtained SF solution was approximately 3.2 wt%, as determined by weighing a dried sample of a known volume. The final concentration of SF was adjusted to 2 wt% with DI water. 2.2. Preparation of SF/CMCS composite hydrogels The tyramine (TA, J&K, Beijing, China) substituted CMCS was synthesized by the coupling reaction of amine groups of TA to carboxylic acid groups of CMCS using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC)/N-hydroxysuccinimide (NHS) activation [20]. Briefly, 1.0 g of CMCS (MW 10 kD, degree of substitution ≥80%, Macklin) was dissolved in 100 mL of 4-Morpholinoethanesulfonic acid buffer (0.1 M, pH 6.0, Macklin). Then, 6 mM of EDC (J&K) and 6 mM of NHS (J&K) were added to the solution. After 30 min stirring, 10 mL of N,N-Dimethylformamide (J&K) solution containing TA (1.5 mM) was added, and the resulting solution was stirred under nitrogen for 3 days at room temperature. The reaction mixture was neutralized with 1 M sodium hydroxide (Macklin) and then purified by dialyzing (MWCO 3500) exhaustively against 0.1 M sodium chloride for 1 day and DI water for another 2 days. The final product was collected after freeze-drying, and stored at 4 °C until use. The percentage of TA that conjugated to the CMCS was determined using UV-spectrophotometer [30,31]. Briefly, the aqueous solution of CMCS-TA was determined at
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the absorbance of 275 nm, and the concentration of the phenolic group in the solution was calculated from a standard curve of TA. SF/CMCS composite hydrogels were prepared by chemical crosslinks with HRP/H2O2, followed by physical cross-links with ethanol treatment as previously described [15]. Briefly, CMCS-TA was dissolved in SF solutions (20 mg/mL) with final CMCS-TA concentration at 2 mg/mL. 10 μL of HRP solution (1000 U/mL) was added to 1 mL of mixed solution of SF and CMCS-TA, then 10 μL of H2O2 (1% v/v) was added and mixed by gentle pipetting. After incubation of the mixture at 37 °C for 1 h, the SF/CMCS composite hydrogels with chemical cross-links were formed. Afterward, the hydrogels were immersed in ethanol (75% v/v) for different time periods (1, 2, 4, or 8 h) to induce the formation of β-sheet structures and physical molecular networks. Then the SF/CMCS composite hydrogels with chemical-physical crosslinks were obtained by washing with DI water to remove ethanol and other residues. The chemically cross-linked hydrogels without ethanol treatment were used as control. Depending on the time of ethanol treatment, the groups of SF/CMCS composite hydrogels were marked as EA0 (control), EA1 (1 h), EA2 (2 h), EA4 (4 h), and EA8 (8 h), respectively. The detailed information of each hydrogel was shown in Table 1.
2.3. Characterization of SF/CMCS composite hydrogels The structure of the hydrogels was analyzed with Fourier transform infrared (FTIR). The hydrogels were quickly frozen in −80 °C followed by freeze-drying, and were cut into slices. The FTIR spectra were obtained in the wavenumber range of 400–4000 cm−1 using FTIR spectrophotometer (PerkinElmer S100, Massachusetts, USA). The thermal properties of the hydrogels were evaluated by thermal gravimetric analysis (TGA). The hydrogels were quickly frozen in −80 °C followed by freeze-drying, and then grinded into powder. TGA was performed on a 1600LF thermal gravimetric/differential scanning calorimetry (TGA/DSC1) instrument (Mettler Toledo, Bern, Switzerland) by heating rate of 2 °C/min under a nitrogen gas flow of 50 mL/min. The crystalline structures of the hydrogels were determined with Xray diffraction (XRD). The sample preparation procedure was the same as for TGA. XRD patterns were obtained on X'Pert Rro MPD powder multi-functional X-ray diffractometer (PANalytical, Almelo, Netherlands), using Cu Kα radiation (40 mA, 40 kV) with a scanning speed of 6°/min. The two-theta range was collected from 5° to 60° and analyzed by JADE6 software. The morphologies of the hydrogels were characterized using scanning electron microscopy (SEM). The hydrogels were pre-frozen at −80 °C, then freeze-dried for 24 h, crosscut with the blade, adhered on the conductive board, and sputter-coated with gold. The morphologies of the surface and section were viewed using Hitachi S3400NIISEM (Hitachi, Tokyo, Japan) operated at 2 kV accelerating voltage.
Table 1 contains detailed information of EA0-EA8 hydrogels including the concentrations of SF and CMCS-TA, the contents of HRP and H2O2 for chemical cross-links, the ethanol treatment time for physical cross-links, the water contents and compressive modulus for each hydrogel. Name SF (mg/mL) CMCS-TA (mg/mL) HRP (1000 U/mL, % v/v) H2O2 (1% v/v, % v/v) Ethanol treatment time (75% v/v, h) Water (wt%) Compressive modulus (kPa)
EA0
EA1
EA2
EA4
EA8
20 2 1 1 0
20 2 1 1 1
20 2 1 1 2
20 2 1 1 4
20 2 1 1 8
91.9 ± 1.5 13.3 ± 4.0
85.6 ± 1.5 59.5 ± 20.6
83.3 ± 2.0 430.0 ± 62.0
77.7 ± 2.7 491.8 ± 36.2
78.1 ± 4.5 828.8 ± 145.2
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2.4. Properties of SF/CMCS composite hydrogels
2.7. GAG quantification
The swelling ratios of the hydrogels were measured in DI water using gravimetric method. The freeze-dried hydrogels (weighed as W1) were immersed in DI water, and kept at 37 °C for 2 days until equilibrium of swelling was reached. The weight of fully swollen hydrogels was recorded immediately after the excess of water lying on the surfaces was absorbed with filter paper, and weighed as W2. The equilibrium water uptake was calculated using the equation: (W2 − W1) / W2 × 100%. The in vitro enzymatic degradation of the hydrogels was measured gravimetrically at 37 °C. The hydrogels (weighed as W0) were immersed in 2 mL PBS with 0.2 units of protease (Type XIV). Enzyme solution was replaced every 3 days for a period of 56 days. Sodium azide (0.05 w/v%) was added to the solution to prevent microbial growth. At specified time points, the hydrogels were removed and weighed (Wt). Samples without the protease were taken as control. The weight remaining ratio was defined as [1 − (W0 − Wt) / W0] × 100%. The rheological properties of the hydrogels were performed with DHR-1 rheometer (TA Instruments, Delaware, USA) by using parallel plates (diameter of plates: 20 mm, gap between plates: 500 nm) in the oscillatory mode at 37 °C. The frequency sweeps were carried out in the range of 0.1–10 Hz at a constant strain of 1%. The mechanical properties of the hydrogels with known sizes (height and diameter) were measured using Instron 5969 testing frame (Instron Instruments, Massachusetts, USA). The compression test was conducted with a pre-strain force of 0.06 N load cell at room temperature. Then the constant strain force was maintained at 2 mm/min till the rupture of hydrogels. The compressive modulus was obtained from the linear elastic region of the stress-strain curves [32].
GAG content of the chondrocyte-seeded hydrogels was evaluated using Blyscan Sulfated Glycosaminoglycan Assay (Biocolor, Antrim, UK) according to the manufacturer's instructions. Cell-seeded hydrogels at 7 and 14 days were digested by papain (Sigma) solution at 60 °C for 6 h. 200 μL of dye solution was transferred to individual wells of a 96well plate, and the absorbance at 656 nm was measured using a microplate reader (Thermo Scientific, Massachusetts, USA). The GAG content was normalized to the dry weight of the remained gel and to DNA content of each sample, respectively.
2.5. Chondrocyte isolation and culture Chondrocytes were harvested from the femoral condyles of 4 week old New Zealand rabbits in accordance with the guidelines approved by the Institutional Animal Care and Use Committee of Third Military Medical University (Chongqing, China). The articular cartilages were cut into slices and digested by 2 mg/mL of type II collagenase (Sigma, Missouri, USA) in DMEM/F12 (Gibco, New York, USA) at 37 °C in a 5% CO2 incubator for 12 h. After that, cells were collected from the digestion solution through a 40 μm cell strainer (BD Bioscience, New Jersey, USA), centrifuged at 500g for 5 min. Then, the collected cells were resuspended in complete medium comprised of DMEM/F12, 10 v/v% fetal bovine serum (FBS, Gibco), and 1 v/v% of penicillin-streptomycin (Gibco). Culture medium was changed every 3 days. The fabricated hydrogels (8 mm in diameter and 3 mm in height) were preconditioned with complete medium in a 24-well plate at 37 °C for 6 h before cell seeding. Chondrocytes were seeded on each hydrogel at a density of 1 × 104 cells/well. After 4 h, the hydrogels were transferred to another 24-well plate containing chondrogenic medium. The medium was changed every 2 days.
2.8. Gene expression and histological analysis First, total RNA was extracted using Trizol (Roche, Basel, Switzerland). Subsequently, RNA was reverse transcribed into cDNA using the Transcriptor cDNA Synth. kit 2 (Roche). Real-time polymerase chain reaction (RT-PCR) was carried out using FS Essential DNA Green Master (Roche). Collagen type II (COL2A1), aggrecan (ACAN), collagen type I (COL1A1), collagen type X (COL10A1), and SOX9 genes were quantified. The results were analyzed according to the 2−ΔΔCT method and normalized to the housekeeping gene GAPDH. The target gene primers (Sangon Biotech, Shanghai, China) were designed as follows: COL2A1, forward, CAACAACCAGATCGAGAGCA, reverse, CCAGTAGTCAC CGCTCTTCC; ACAN, forward, CGACATCAGTGGAGACCTCA, reverse, AGACTGCCAGAGTCCAGCTC; COL1A1, forward, CAATCACGCCTCTCAG AACA, reverse, TCGGCAACAAGTTCAACATC; COL10A1, forward, GAAT GGCACGCCTGTAATGT, reverse, CCATTGGACTCAGCGTTAGG; SOX9, forward, GGAAGCTCTGGAGACTGCTG, reverse, GCGGCTGGTACTTGTAGTC C; GAPDH, forward, CGACATCAAGAAGGTGGTGA, reverse, ATCGAA GGTGGAGGAGTGG. The chondrocyte-seeded hydrogels were fixed in 4% paraformaldehyde for 2 h at room temperature. For immunohistochemistry, the Anti-collagen type I and Anti-collagen type II (Abcam) primary antibodies were employed under 4 °C for overnight. Then goat anti-rabbit secondary antibodies were used. The images were observed using microscope (Olympus BX50, Japan). 2.9. Subcutaneous implantation The in vivo biological response of the SF/CMCS composite hydrogels was evaluated by subcutaneous implantation. The Kunming mice (weight 20–30 g) were anesthetized with 1 w/v% of pentobarbital (50 mg/kg) by intraperitoneal injection. Then, the SF/CMCS hydrogels (6 mm in diameter and 4 mm in height) were then implanted into lateral incisions on the dorsal region. The mice were euthanized after implantation for 2 and 4 weeks. The specimens along with the adjacent tissues were fixed in 4% paraformaldehyde at 4 °C overnight, dehydrated in sucrose, embedded in frozen section media, and frozen sectioned at a thickness of 10 μm. The histological sections were observed with hematoxylin & eosin (H&E) staining.
2.6. Cell viability and proliferation
2.10. Statistical analysis
Cell viability was determined using a Live/Dead cell double stain kit (Yeasen, Shanghai, China) at 7 and 14 days. In brief, cell-seeded hydrogels were incubated in Assay Buffer containing 2 mM CalceinAM and 1.5 mM propidium iodine for 15 min at 37 °C. Labeled cells were visualized using a fluorescent microscope. Live cells were stained green, whereas dead cells were stained red. Cell proliferation at 7 and 14 days was evaluated by quantifying DNA. DNA content was detected with AccuGreenTM High Sensitivity dsDNA Quantitation Kits (Biotium, California, USA) according to the manufacturer's instructions. DNA content of cell-seeded hydrogels was expressed as DNA amount normalized to the dry weight of the remained gel.
All data were presented as the mean ± standard deviation (SD). Statistical significance was determined by one-way ANOVA followed by Tukey's post hoc analysis (SPSS statistics 23, IBM). p values b 0.05 were considered statistically significant. 3. Results and discussion 3.1. Preparation of SF/CMCS composite hydrogel The synthetic scheme of CMCS-TA was shown in Fig. 1A. CMCS-TA was synthesized by the coupling reaction of the amino group of TA to the carboxylic acid groups of CMCS using EDC/NHS activation. The
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Fig. 1. The scheme of hydrogel formation and characterization of CMCS-TA. (A) CMCS was partially decorated with TA. (B) SF/CMCS composite hydrogel was prepared with chemical crosslinks by HRP/H2O2, followed by physical cross-links with ethanol treatment. (C) 1H NMR spectra of TA, CMCS, and CMCS-TA. (D) UV–Vis absorbance spectrum of H2O, CMCS, CMCS-TA, and TA.
structure of the CMCS-TA conjugates was confirmed by 1H NMR (Fig. 1C). When compared with CMCS spectrum, 1H NMR spectra of CMCS-TA exhibited new signals at δ 6.67 and δ 6.98, which indicated the presence of the aromatic protons of TA. The degree of TA molecules per 100 repeating units of CMCS was about 10.08%, calculated from a standard curve of TA (Fig. 1D). SF/CMCS composite hydrogels were prepared by chemical crosslinks with HRP/H2O2, followed by physical cross-links with ethanol treatment (Fig. 1B). The CMCS-TA polymer was freeze-dried and then dissolved in SF solution. The enzyme-mediated polymerization strategy was applied to form chemical cross-links with HRP and H2O2. SF contains tyrosine residues with a content of approximately 5 mol% [33], therefore the TA groups of CMCS can react with tyrosine residues of SF to form covalent bonds. HRP can facilitate the cross-links between SF and CMCS-TA via forming free radical species in the presence of H2O2 [34]. Then the chemically cross-linked hydrogels were treated with ethanol to induce the β-sheet formation via the hydrophobic interactions of the (GAGAGS)n sequences in relatively smaller network structures of SF [15]. 3.2. Structure, swelling and degradation of SF/CMCS composite hydrogel Structural changes of the SF/CMCS composite hydrogels were detected by FTIR. In general, the structure of SF at 1625 cm−1 and 1515 cm−1 in the FTIR spectrum is recognized as β-sheet conformation, meanwhile the absorption band at 1260 cm−1 inside the amide III region is also associated with the β-sheet conformation [16,35]. According to Fig. 2A, it was found that a new peak appeared at 1625 and 1515 cm−1 after ethanol treatment, which could be attributed to the conformational transition of SF from random coil to β-sheet [36,37]. The absorption band changed very slightly when the ethanol treatment
time increased from 2 h to 8 h, indicating that the conformational transition of SF from random coil to β-sheet had basically completed after 2 h. Then, the conformational changes of the SF/CMCS composite hydrogels were evaluated by DSC (Fig. 2B). Compared with EA0 hydrogel, the degradation peak [38] appeared at 250–280 °C in all the ethanol-treated hydrogels. This result indicated the presence of βsheet structure in ethanol-treated hydrogels. Similarly, there were only slight changes in the degradation peak between EA2, EA4, and EA8 hydrogels, which was consistent with the FTIR results. Moreover, XRD analysis provided a detailed information for the microstructural difference among the ethanol-treated hydrogels. As was shown in Fig. 2C, compared with EA0 hydrogel, the ethanol-treated hydrogels presented a sharper peak around 20°, which is the character peak of β-sheet [16]. When the time of ethanol treatment increased from 2 h to 8 h, there were no obvious changes in XRD curves. Therefore, the XRD results together with FTIR and DSC results, indicated that the SF/ CMCS composite hydrogels possessed a tunable β-sheet structure that had been basically formed after 2 h ethanol treatment. The swelling ability of the hydrogel is used to character its ability to maintain structure and retain enough water. Water absorption and retention is important when hydrogels are used in TE, because it reflects the ability of the hydrogel to transfer oxygen and nutrient [39]. The swelling ratio of the SF/CMCS composite hydrogels was investigated by immersing the freeze-dried hydrogels in PBS to reach equilibrium at 37 °C. As was shown in Fig. 2D and Table 1, all hydrogels showed good swelling ratio, ranging from 78.1 ± 2.6% to 91.9 ± 0.9%. The water uptake ability significantly decreased from 91.9 ± 0.9% to 78.1 ± 2.6%, as the ethanol treatment time increased from 0 to 8 h. These results could be ascribed to the increase of ethanol treatment time that leads to more the formation of β-sheet structure in SF. The recurring hydrophobic regions of alanine and glycine residues in SF form β-sheet
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Fig. 2. The structural analysis, swelling, and degradation of the SF/CMCS composite hydrogels. (A) FTIR, (B) TGA, and (C) XRD curves of the hydrogels. (D) Water uptake, (B) Degradation of the hydrogels in PBS, (C) Degradation of the hydrogels in protease (Type XIV). Data were presented as mean ± SD (n = 3, *p b 0.05, **p b 0.01, ***p b 0.001).
structures [40], and the higher β-sheet content contribute to the higher hydrophobicity. The degradation rate is another key parameter of hydrogels in TE. Here the in vitro degradation was assessed by monitoring the weight loss of the hydrogel post 56 days of incubation with PBS and protease XIV solution, respectively. In Fig. 2E, the remaining masses of EA0 and EA1 hydrogels were 56.6 ± 4.5% and 84.5 ± 1.5% respectively after incubating with PBS up to 56 days. The increase of ethanol treatment time resulted in a significant decrease in degradation rate. However, there was almost no difference between the EA2 (96.9 ± 0.9%), EA4 (98.4 ± 0.1%), and EA8 (97.5 ± 0.8%) hydrogels in degradation, indicating that the β-sheet structure of SF had been basically formed after 2 h ethanol treatment. While in Fig. 2F, the degradation rate of hydrogels
was much faster in the presence of protease XIV. The EA0 and EA1 hydrogels completely degraded after 2 and 10 days in protease XIV, respectively. The degradation of EA2 hydrogel accelerated after 42 days, and eventually EA2 (73.1 ± 5.4%) hydrogel showed 44% reduction in mass over 56 days. However, the EA4 (87.9 ± 1.3%) and EA8 (91.0 ± 1.3%) hydrogels only showed 12% and 9% reduction in mass over 56 days, respectively. The reason is that the hydrogels with higher βsheet content could better protect themselves against enzymatic attack [41]. In cartilage TE, if a scaffold degrades too fast, it might collapse under continual mechanical loading; if a scaffold degrades too slowly, it hinders the growth of the new cartilage tissue. In this study, we showed that the degradation of the SF/CMCS composite hydrogels could be tuned into a proper rate by varying the ethanol treatment time.
Fig. 3. Morphology of the SF/CMCS composite hydrogels. (A) Macroscopic observation of the hydrogels. (B) SEM images of the cross-section of the hydrogels. (C) SEM images of the surface of the hydrogels.
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3.3. Morphologies of SF/CMCS composite hydrogel The macroscopic observation of the SF/CMCS composite hydrogels was shown in Fig. 3A. The hydrogels became darker as ethanol treatment increased from 0 h to 8 h. As was shown in Fig. 3B, the inner morphology of SF/CMCS composite hydrogels was determined by crosssectional SEM images. These images revealed that the composite hydrogels had a pore size range of 50–300 μm with well interconnected three-dimensional porous structure. The EA2, EA4, and EA8 hydrogels had similar pore sizes, because the conformational transition of SF from random coil to β-sheet had basically completed after 2 h ethanol treatment. It has been shown that SF scaffolds with pore sizes of 90–250 μm provide the best environment for adhesion and proliferation of chondrocytes [42]. This pore size range almost overlaps with that of the SF/CMCS composite hydrogels, therefore we expected that the SF/ CMCS composite hydrogels likewise support chondrocyte growth. 3.4. Mechanical properties of SF/CMCS composite hydrogel Generally, the rheological properties are studied via oscillatory rheology to assess the stability of three-dimensional cross-linked networks such as hydrogels. Storage modulus (G′) and loss modulus (G″) represent elastic and viscous behavior of the hydrogels, respectively. As was shown in Fig. 4, all samples exhibited higher G′ values than G″ values in the frequency sweep tests (from 0.1 Hz to 100 Hz), indicating a predominantly elastic behavior of the hydrogels. At a fixed frequency of 1 Hz, the G′ values for EA0, EA1, EA2, EA4, and EA8 hydrogels were 177 Pa, 20,443 Pa, 36,495 Pa, 56,282 Pa, and 88,904 Pa, respectively. There was about a 500-fold increase in G′ values (from 177 Pa to 88,904 Pa) of the hydrogels when the time of ethanol treatment increased from 0 h to 8 h, demonstrating an improvement of the mechanical properties with longer treatment of ethanol. Elasticity of the scaffolding materials has been recognized as a dominating factor of cell fate in TE [43]. Here, the compression test was conducted to verify the elastic properties of the SF/CMCS composite hydrogels. As was shown in Fig. 5A, the stress-at-failure and strain-atfailure values were determined from the resulting compressive stressstrain curves. The ethanol treatment time had a great effect on the stress and strain of these hydrogels at break. Fig. 5B showed that 8 h of ethanol treatment led to a 30-fold increase in compressive modulus (from
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13 kPa of EA0 to 829 kPa of EA8). This was in accordance with a previous study showing that the SF molecular chains could be further immobilized by forming extensive β-sheet structures after ethanol treatment, which in turn greatly improves both strength and elasticity of the SF hydrogel. Note that the compressive modulus of EA2 (430 ± 62 kPa), EA4 (492 ± 36 kPa) and EA8 (829 ± 145 kPa) hydrogels was in the range of compressive moduli of natural cartilage (300–800 kPa) [44]. Therefore, the results of the compression tests together with the rheological tests, suggested that the mechanical performance of the SF/CMCS composite hydrogels could be adjusted to suit cartilage TE. Taken together the parameters of pore size, swelling, degradation, and mechanical performance, we chose EA2 hydrogel as experimental group and EA0 hydrogel as control in the cellular experiments. 3.5. Cell viability, proliferation, and GAG production Fluorescent images of live/dead cell staining were used to investigate the viability of rabbit chondrocytes seeded on the SF/CMCS composite hydrogels (Fig. 6A). The EA0 and EA2 hydrogels showed uniform distribution and displayed a spherical morphology of cells throughout the culture conditions. Few dead cells were observed within the hydrogels. Next, chondrocyte proliferation was monitored by quantifying DNA content of the hydrogels (Fig. 6B). The DNA content of EA2 hydrogel at day 14 was significantly higher than that at day 7. Furthermore, there was no significant difference in DNA content between EA2 and EA0 hydrogel after 7 and 14 days of culture. These results indicated that both the EA2 and EA0 hydrogels had good biocompatibility with chondrocytes. GAG is an important ECM component of cartilage. When the GAG production of the rabbit chondrocytes seeded on the SF/CMCS composite hydrogels was normalized to the dry weight of remained gel (Fig. 6C), there was no statistical difference between EA0 and EA2 hydrogel. In EA2 hydrogel, it was found that the GAG/dry gel ratio had a significant increase at day 14, in comparison to that at day 7. When the GAG production from the cells seeded on the SF/CMCS composite hydrogels was normalized by their respective DNA content (Fig. 6D), the normalized value of EA2 hydrogel was higher that of EA0 hydrogel at both day 7 and day 14, indicating that averagely each single cell in EA2 hydrogel had a better ability to synthesize GAG. Similarly, a study showed that SF scaffolds with β-sheet conformational, derived from HRP-mediated cross-linking SF solution in combination with salt-
Fig. 4. Rheology of the SF/CMCS composite hydrogels.
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leaching and freeze-drying methodologies, facilitated cartilage-specific ECM production as scaffolds for cartilage regeneration [45]. This finding together with our results, suggests that the appropriate β-sheet structures in the SF/CMCS composite hydrogels enhance cartilaginous matrix production. 3.6. Gene expression and immunohistochemical staining To compare the ability of EA0 and EA2 hydrogel in maintaining chondrocyte phenotype, the expression levels of cartilage related genes including chondrogenic markers COL2A1 and ACAN, fibrotic marker COL1A1, hypertrophic marker COL10A1, and early chondrogenic transcription factor SOX9 were quantified by RT-PCR after 7 and 14 days of culture. As was shown in Fig. 7A, the COL2A1 expression of chondrocytes in EA2 hydrogel increased significantly from day 7 to day 14 (55.2 folds); moreover, the chondrocytes in EA2 hydrogel expressed about 22.8 folds of COL2A1 compared to those in EA0 hydrogel at day 14. For ACAN expression (Fig. 7B), there was a higher
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3.7. In vivo hydrogel biocompatibility The in vivo biological response of SF/CMCS composite hydrogels was evaluated by subcutaneous implantation in a mouse model. The
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expression in EA2 hydrogel than in EA0 hydrogel at day 14 (1.7 folds). As for expression of SOX9, COL1A1, and COL10A1 (Fig. 7C, D, and E), there was no difference either between groups (EA0 vs. EA2) or with time (day 7 vs. day 14). The higher expression of chondrogenic markers COL2A1 and ACAN in EA2 group than in EA0 group, indicated that the SF/CMCS composite hydrogels with chemical-physical cross-links had a beneficial impact on chondrogenic phenotype. The immunohistochemical staining of chondrocytes-seeded SF/ CMCS composite hydrogels were shown in Fig. 8. The EA0 and EA2 hydrogels exhibited limited staining for collagen type I (fibrotic marker) at Day 7. Compared to EA0 hydrogels, EA2 hydrogels exhibited a stronger staining of collagen type II, indicating that EA2 hydrogels were more suitable in maintaining chondrocyte phenotype.
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Fig. 6. Cell viability, proliferation, and GAG production on the SF/CMCS composite hydrogels. (A) Live-dead staining (green: live, red: dead). (B) DNA content of cell-seeded hydrogels was expressed as DNA amount normalized to the dry weight of the remained gel after 7 and 14 days of culture. (C) GAG production normalized to the dry weight of the remained gel, (D) GAG production normalized to the DNA content of chondrocytes on SF/CMCS hydrogels after 7 and 14 days of culture (n = 3, *p b 0.05, **p b 0.01). (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)
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Fig. 7. Gene expression analysis of the rabbit chondrocytes on the SF/CMCS composite hydrogels. (A) COL2A1, (B) ACAN, (C) SOX9, (D) COL1A1, and (E) COL10A1 were normalized to GAPDH and expressed as relative values to day 7. (n = 3, *p b 0.05, ***p b 0.001).
macroscopic images of the explants showed that EA0 and EA1 hydrogels mostly degraded, EA2 hydrogels slightly degraded, while EA4 and EA8 hydrogels hardly degraded (Fig. 9A). This was consistent with the in vitro degradation result, and further confirmed that the SF/CMCS composite hydrogels had tunable degradation through changing ethanol treatment time. The macroscopic images of the explants in situ showed that hydrogels were integrated in the subcutaneous tissue
after 14 and 28 days of implantation (Fig. 9B and C). There was no evident sign of infections or inflammatory responses in the surrounding tissue. The hydrogels and surrounding tissues were collected and stained with H&E for qualitative histological examination. From the H&E staining images (Fig. 9D), it was observed that a thick layer of connective tissue adhered on the entire surface of the hydrogels after 14 and 28 days of implantation. The absence of edema or signs of
Fig. 8. Immunohistochemical staining of chondrocytes-seeded SF/CMCS composite hydrogels at 7 days.
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Fig. 9. In vivo degradation and biocompatibility of the SF/CMCS composite hydrogels. (A) Macroscopic images of the hydrogels after 14 days and 28 days of implantation. (B, C) Macroscopic images of the EA0 hydrogel in situ after implantation for 14 days and 28 days. (D) H&E staining of the hydrogel explants after 14 days and 28 days of implantation.
neutrophils after implantation, revealed that no acute inflammation was induced by the implanted biomaterial. These results suggested that the SF/CMCS composite hydrogels exhibited an acceptable tissue compatibility in vivo. 4. Conclusions In this study, SF/CMCS composite hydrogels were prepared by chemical cross-links with HRP/H2O2, followed by physical cross-links with ethanol treatment. Frist, FTIR, DSC, and XRD verified that SF/CMCS composite hydrogels had a tunable β-sheet structure. Next, the physicochemical properties such as morphology, equilibrium swelling, degradation, rheological properties, and mechanical properties of the hydrogels could be adjusted by varying the ethanol treatment time. Then, rabbit chondrocytes were seeded on the selected hydrogels (EA2) for cellular, biochemical, and gene expression assays, and such hydrogels maintained and even promoted chondrogenic phenotype. Last, SF/CMCS composite hydrogels exhibited an excellent tissue compatibility in vivo. Overall, our studies suggest that the chemical-physical cross-linking SF/CMCS composite hydrogels with tunable mechanical properties and biodegradability are a promising scaffolding material for cartilage TE. Declaration of Competing Interest We declare that we have no conflict of interest. Acknowledgements This work was supported by National Natural Science Foundation of China (Grant no. 81572107). References [1] D. Correa, S.A. Lietman, Articular cartilage repair: current needs, methods and research directions, Semin. Cell Dev. Biol. 62 (2017) 67–77.
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