Enhanced mechanical properties of thermosensitive chitosan hydrogel by silk fibers for cartilage tissue engineering

Materials Science and Engineering C 33 (2013) 4786–4794

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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec

Enhanced mechanical properties of thermosensitive chitosan hydrogel by silk fibers for cartilage tissue engineering Fereshteh Mirahmadi a,b, Mohammad Tafazzoli-Shadpour a,⁎, Mohammad Ali Shokrgozar b,⁎⁎, Shahin Bonakdar b a b

Faculty of Biomedical Engineering, Amirkabir University of Technology, Tehran, Iran National Cell Bank of Iran, Pasteur Institute of Iran, Tehran, Iran

a r t i c l e

i n f o

Article history: Received 16 February 2013 Received in revised form 12 July 2013 Accepted 29 July 2013 Available online 6 August 2013 Keywords: Cartilage engineering Scaffold Chitosan hydrogel Silk Mechanical characterization

a b s t r a c t Articular cartilage has limited repair capability following traumatic injuries and current methods of treatment remain inefficient. Reconstructing cartilage provides a new way for cartilage repair and natural polymers are often used as scaffold because of their biocompatibility and biofunctionality. In this study, we added degummed chopped silk fibers and electrospun silk fibers to the thermosensitive chitosan/glycerophosphate hydrogels to reinforce two hydrogel constructs which were used as scaffold for hyaline cartilage regeneration. The gelation temperature and gelation time of hydrogel were analyzed by the rheometer and vial tilting method. Mechanical characterization was measured by uniaxial compression, indentation and dynamic mechanical analysis assay. Chondrocytes were then harvested from the knee joint of the New Zealand white rabbits and cultured in constructs. The cell proliferation, viability, production of glycosaminoglycans and collagen type II were assessed. The results showed that mechanical properties of the hydrogel were significantly enhanced when a hybrid with two layers of electrospun silk fibers was made. The results of GAG and collagen type II in cell-seeded scaffolds indicate support of the chondrogenic phenotype for chondrocytes with a significant increase in degummed silk fiber–hydrogel composite for GAG content and in two-layer electrospun fiber–hydrogel composite for Col II. It was concluded that these two modified scaffolds could be employed for cartilage tissue engineering. © 2013 Elsevier B.V. All rights reserved.

1. Introduction Articular cartilage injury occurs frequently as a result of sportrelated events, diseases, and traumas. The fact that articular cartilage is poorly capable of effective repair following traumatic injury has been well recognized [1]. Although various practices are employed to repair cartilage damages, current methods of treatment remain unsatisfactory and inefficient [2]. With the development of tissue engineering, reconstructing cartilage provides a new strategy for cartilage repair. Many synthetic and natural polymeric materials have been used as scaffolds for cartilage tissue regeneration. Natural polymers are often used due to their enhanced biocompatibility and biofunctional motives [3]. Such biomaterials including collagen [4], cellulose [5], chitosan [6], fibrin [7] and silk [8] maintain a differentiated cell phenotype and allow rapid cell expansion. On the other hand, in general they are not mechanically stable enough and the rate of their biodegradation is practically high [9]. To ⁎ Correspondence to: M. Tafazzoli-Shadpour, Faculty of Biomedical Engineering, Amirkabir University of Technology, 424 Hafez Ave., Tehran, Iran. Tel.: +98 21 64542385. ⁎⁎ Correspondence to: M.A. Shokrgozar, National Cell Bank of Iran (NCBI), Pasteur Institute of Iran, 69 Pasteur Ave., P.O. Box 1316943551, Tehran, Iran. Tel.: +98 21 66492595. E-mail addresses: [email protected] (M. Tafazzoli-Shadpour), [email protected] (M.A. Shokrgozar). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.07.043

enhance mechanical stability specific components including natural or synthetic materials are used to fabricate composite structures which can offer biological cues to promote tissue-specific interactions, while providing desired mechanical properties [10,11]. Natural polysaccharides and proteins which are employed as mimicry of real tissues dramatically influence cell behavior and tissue formation. The interactions between such biomaterials within the extra-cellular matrix (ECM) create macromolecular structures by interconnection [12]. Chitosan (CS) and Bombyx mori silk fibroin (SF) are two prominent natural materials for the design of tissue engineered scaffolds due to their superb biocompatibility, degradability, excellently engineered structures and mechanical characteristics. Silk fibroin has long been used in surgical sutures due to its suitable biocompatibility. It is a natural fibrous protein that has been used as a potential biomaterial for several biomedical and biotechnological applications because of its excellent biological compatibility and mechanical strength. When silk fibroin is used in regenerated form solution it shows the in vivo smallest inflammatory reactions [13–15]. Silk is also considered as a biodegradable material, although its degradation behavior is closely related to the molecular weight, structure, and the preparation method [14,16]. As an interesting biopolymer silk has been effectively used in tissue engineering of bone [17], cartilage [8,18], and artery [19]; and is used in both natural fiber and regenerated forms for tissue engineering purposes [8,14].

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Chitosan, an amino polysaccharide derived from chitin via deacetylation, has been used as a biomaterial because of its biocompatibility, biodegradability, intrinsic antibacterial activity, wound-healing properties, and low immunogenicity [20]. Chitosan shows unique tissue-adhesive properties due to its polycationic nature [21]. It is structurally similar to various glycosaminoglycans (GAGs) that exist in articular cartilage, which makes it a promising candidate for cartilage engineering [22]. Hence, CS has demonstrated proper cytocompatibility with chondrocytes [21,23] and human mesenchymal stem cells [24]. Chitosan/glycerophosphate (CS/GP) hydrogel, a novel thermosensitive hydrogel, initially reported by Chenite et al. [25], has been known to be an important applicable biological substance due to its sol–gel transition at body temperature. The CS/GP hydrogel has been thoroughly investigated for tissue reconstruction including cartilage [21,23], and bone [26,27] or for use as a drug delivery vehicle [28]. Regardless of such interesting features, usage of CS scaffolds is usually limited due to their insufficient mechanical properties. Hence, their mechanical functionality is generally improved by blending with other polymers. Combinations of chitosan with other materials such as collagen [26,29], silk [30–33], starch [34], and gelatin [35] have been studied for biomedical applications. In addition, composite biomaterials are increasingly being used in tissue engineering in order to mimic the complex properties of native ECM. There are several studies that surveyed different types of tissue engineered composites as cartilage scaffolds in order to prepare new biomaterials with enhanced properties. Bhardwaj et al. added SF to CS in order to reduce the degradation of chitosan containing scaffolds [32]. Guo et al. generated hydrogel composites of crosslinked oligo(poly(ethylene glycol)fumarate) and gelatin microparticles to fabricate a bilayer osteochondral construct and loaded TGF-b1 in the chondrogenic layer. Results showed that this composite significantly stimulated chondrogenic differentiation of MSCs [36]. The purpose of this work was to develop composite matrix of chitosan/glycerophosphate (CS/GP) hydrogel and silk fibroin fibers to obtain suitable mechanical properties while maintaining biological functionality in order to use as scaffold in cartilage engineering. The usage of SF within the structure improved mechanical characteristics of matrix while supporting chondrogenic phenotype of chondrocytes. 2. Materials and methods Here, we first characterized the concentration of GP (range: 8.3–25% w/v) required to cause gelation of chitosan, followed by fabrication of composites. Silk fibroin (SF) was produced either in the form of sheets or chopped fibers. Sheets of SF nanofibers were fabricated by electrospinning. The sheets were coated with a hydrogel solution, covered by another sheet, and then allowed for sol–gel transition to form a fiber-reinforced hydrogel sandwich structure. In the other scaffold degummed and chopped silk fibers were dispersed in hydrogel solution and allowed for sol–gel transition. The resulting materials were characterized by SEM observation, mechanical and swelling properties, and cytocompatibility; and the resultant structures were compared. To function appropriately as a scaffold for cartilage engineering, enhanced mechanical properties were aimed while proper swelling and biocompatibility were required to be maintained. 2.1. Formulation of chitosan hydrogel Chitosan chloride powder (CS, Novamatrix, Norway; Protasan® UP CL 213) was dissolved in the distilled water at 2% (w/v) chitosan, then autoclave sterilized. The chitosan solution was mixed with concentrated stock solution of filter-sterilized pre-cooled β-sodium glycerophosphate 50% (w/v) (GP, Sigma, USA), to yield a liquid CS/GP solution including different combinations of CS/GP concentration (Table 1). All procedures were performed on ice to maintain a liquid state before gelation.

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Table 1 Final concentration of CS/GP solutions. Group name

CS/GP1.67 CS/GP1.6 CS/GP1.5 CS/GP1.4 CS/GP1

Volume

Concentration %

CS

GP

CS

GP

5 4 3 2.8 2.5

1 1 1 1.2 2.5

1.67 1.6 1.5 1.4 1

8.33 10 12.5 15 25

By incubation of the CS/GP solution at 37 °C, gelation was initiated [23,28]. 2.2. Gelation time determination The test tube inverting method was used to measure the gelation time at constant temperature of 37 °C in a water bath [27,28,37]. The time measurement was initiated when samples of 1 mL in glass vials (1.2 cm inner diameter) were incubated in the bath. The fluidity of the samples was observed every 30 s by tilting the tubes. The time at which flow stopped was taken as the gelation time and the values were recorded. Since these scaffolds are meant to be used for in vivo applications, gelation time and GP concentration were used as criteria for selection of optimized combination of CS/GP. 2.3. Gelation temperature To determine gelation temperature, rheological measurement was performed with a MCR 300 rheometer equipped with a cone and plate geometry, which required about 1 mL of the solution per sample [38]. The selected CS/GP solution (CS/GP1.6) was inserted into the rheometer. To determine the gelation temperature oscillatory measurements were performed at 1 Hz, while the temperature was elevated at the rate of 1 °C/min between 10 and 60 °C. Therefore, the dynamic viscoelastic properties such as the dynamic shear elastic (storage) modulus (G′) and the viscous (loss) modulus (G″) were measured as a function of strain and temperature. The values of the strain amplitude were checked before to ensure that all measurements were carried out within the linear viscoelastic regime. To do this, the frequency sweep measurements were carried out from 0.1 to 100% strain [39]. The phase lag (tan δ = G″ / G′) was taken to determine the gelation temperature at which G′ and G″ were equivalent in value (δ = 45°) [27,40]. 2.4. Scanning electron microscopy (SEM) For SEM observation, selected hydrogel (CS/GP1.6) samples were fabricated and then frozen in −70 °C and freeze-dried for 24 h. The dried gel was cut with a sharp blade to expose internal microstructure and sputter coated with platinum–gold for SEM observation (AIS2100, South Korea). In the SEM photos, the pore size of samples was analyzed using the ImageJ image-visualization software. 2.5. Preparation of silk fibroin B. mori silkworm cocoons were obtained from Gilan University (Rasht, Iran). B. mori cocoons were boiled with 0.02 M Na2CO3 for 60 min and washed with warm distilled water several times to remove the glue-like sericin proteins from the silk fibers, and then air dried. The degummed silk fibers were dissolved in 9.3 M LiBr at 60 °C, and then were dialyzed in a cellulose dialysis membrane (12,000 MWCO, Sigma, USA) against deionized water for 3 days with changing water several times for the removal of the salt [41]. The regenerated silk fibroin sponge was obtained by lyophilization (MC4L, UNICRYO freeze dryer, −60 °C, Germany). FTIR measurements were used to make sure the regenerated silk fibroin structure is properly regenerated (graph is not

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shown). The frequencies of amide regions in regenerated silk fibroin comply with other studies [12,42]. FTIR spectroscopy of regenerated silk fibroin showed peaks in the Amide I region at 1650 cm−1, Amide II at 1537 cm−1 and Amide III at 1236 cm−1. The 10% (w/v) silk–formic acid solution was prepared for electrospinning by dissolving the silk fibroin sponge into the formic acid (98–100%, Merck) [18,30].

modulus and standard deviation were determined from the linear section of the stress–strain curve. A dynamic mechanical analyzer, DMA (Tritec 2000 DMA, England) was utilized to measure dynamic elastic and viscous moduli of samples at 37 °C. The DMA was used in the oscillatory mode at a frequency of 1 Hz [44]. 2.9. Cell characterization

2.6. Electrospinning and fiber characterization The silk–formic acid solution was placed in a 3-mL syringe (20-G). The electric field between the collection plate (cathode) and the needle tip (anode), solution flow rate, and the distance between the needle tip and the aluminum foil covered target plate were adjusted until a stable jet was obtained (Table 2). The optimized operating parameters for production of the nanofibers were as follows: voltage: 40 kV, flow rate: 0.5 mL·min−1, and distance between the needle tip and the target: 15 cm at room temperature. The entire apparatus was mounted under an insulator hood [8,25]. The morphology of electrospun fibers was observed under a scanning electron microscope. In the SEM photos, the fiber diameters of the nanofibrous membranes were analyzed using the ImageJ [43]. 2.7. Fabrication of fiber-reinforced composite Composite scaffolds were fabricated using the selected hydrogel (CS/GP1.6) in two patterns. First a hydrogel-degummed fiber scaffold was prepared by dispersing chopped degummed silk fibers in hydrogel solution in liquid state and vortexed to aid homogeneous dispersion (CS/GP-D). In the second pattern, electrospun fiber layers were utilized to form a sandwich-like composite. The nanofiber mesh was placed on the plate and the hydrogel solution was coated over the fibers, followed by placing another nanofiber layer on the top of the solution, and then allowed for sol–gel transition to form a two-layer structure (CS/GP-L). The solution acted as a “glue” to hold the fibrous layers together. The weight ratio of CS/GP to silk fiber was kept the same for the two composites. The CS/GP1.6 hydrogel was used as control to compare the results. 2.8. Scaffold characterization 2.8.1. Water uptake For determination of water uptake, 3 disks of each scaffold (8 mm diameter) were prepared and allowed to reach equilibrium at 37 °C and the swollen weight (Ws) was measured. Next, the samples were dried under ambient conditions for 12 h, followed by drying at 40 °C for 12 h, and then the dry sample weights (Wd) were recorded. The equilibrium water uptake of the scaffolds was determined by Q = (Ws − Wd) / Wd [44]. 2.8.2. Mechanical analysis Cylinder-shaped hydrated samples measuring 8 mm in diameter and 6–7 mm in height were used. Mechanical properties of the scaffolds were assessed through dynamic mechanical analysis, unconfined compression, and spherical indentation [23]. The compressive mechanical properties of scaffolds were tested using a Zwick/Roell Z050 (Germany) equipped with a 50 N load cell at ambient room temperature. The samples in the wet state were examined with crosshead speed 0.5 mm/min [33]. At least three specimens were tested for each sample. The compressive stress and strain values were graphed and the average compressive Table 2 Variable parameters of electrospinning. Factor

Factor level

Electric field (kV) Spinning distance (cm) Flow rate (mL/h)

10, 15, 20, 30, 40 7, 10, 12, 15, 20 0.25, 0.5, 1

2.9.1. Isolation of chondrocytes To assay cell-scaffold interaction regarding cartilage engineering, articular chondrocytes were isolated from knee joints of New Zealand white rabbits according to published protocols [45]. All experiments were approved by the Ethical Committee of National Cell Bank of Iran (NCBI). The cartilaginous tissue was cut off from the joint, washed several times with antibiotic containing medium and then minced to small pieces using a scalpel blade. The obtained slices were predigested in 0.25% trypsin–EDTA solution (Sigma, USA) for 1 h and then transferred to the collagenase type II solution (0.08 mg/mL, Sigma, USA) for 12 h in an incubator. The resulting cell suspension was passed through a 40 μm filter. Cells were then harvested by centrifugation at 1500 rpm for 5 min and cultured in DMEM (GIBCO, Scotland)/Ham's F12 supplemented with 10% fetal bovine serum (FBS) (Seromed, Germany) in an incubator at 37 °C, 5% CO2 [23,45]. 2.9.2. MTT assay The sterile scaffolds (3 ± 0.5 cm2/mL of culture medium) were placed in 12-well tissue culture clusters and washed three times with fresh culture medium, and incubated at 37 °C in fresh culture medium for 1, 7, and 14 days. Serum-supplemented culture medium (DMED) at the same condition was used as the control group. Cells were seeded into 96-well plates at a density of 1 × 104 cells/mL. The culture medium (100 μL) was replaced with 1, 7, and 14 day extracts after 24 h. Utilizing 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT, Sigma, USA), the viability of every group was assessed. The extracts were eliminated after another 24 h and 100 μL of a 0.5 mg/mL MTT solution was added to each well and were incubated at 37 °C for additional 4 h [26]. Then MTT solution was removed carefully and 150 μL isopropanol (Sigma, USA) was subsequently added to each well to dissolve the MTT formazan purple crystals. Absorbency of the solution was measured at 545 nm using an ELISA Reader (Stat Fax-2100, USA). The relative viability or cell growth (%) normalized by control group was calculated from the following equation: Viability ð% Þ ¼ ðaverage optical density of samplesÞ= ðaverage optical density of controlÞ100%:

2.9.3. Live/dead assay Cell viability within the constructs was observed using acridine orange–propidium iodide (AO/PI) staining. Chondrocytes were mixed with CS/GP hydrogel and poured into a 24-well cell culture plate, and then incubated at 37 °C for 30 min to allow it to gel. After 7 days, the hydrogels containing chondrocytes were incubated with the AO/PI mixture for 10 min and observed under a fluorescence microscope. Live cells were stained green (AO) while dead cells were colored red (PI) [23]. 2.9.4. Determination of glycosaminoglycan and collagen II content Chondrocytes were blended with 2 mL sterile hydrogel solution and used to fabricate each type of CS/GP, CS/GP-L and CS/GP-D scaffolds. The resulting constructs were harvested in medium and used for Col II and GAG content measurements. The proteoglycan content was determined by the amount of sulfated GAGs released into media using dimethylmethylene blue (DMMB, Sigma, USA). At each time interval 300 μL medium of each sample was aspirated and transferred to a 2-mL vial; then 1.2 mL acetone (Merck) was added and kept

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at − 20 °C for 24 h. Samples were then centrifuged at 1800 rpm for 30 min at 4 °C. The supernatant was poured out and precipitated pellet was suspended in 1 ml papain digestion solution in PBS containing 20 μg/mL papain and 5 mM L-cysteine, followed by incubation at 60 °C for 16 h. The samples were then boiled for 10–15 min. The GAG content was quantified using ELISA plate reader at 545 nm. The standard curve was generated using certain concentrations of chondroitin sulfate C (shark cartilage extract, Sigma, USA) [24,45,46]. The amount of Col II produced in the constructs was determined using collagen type II ELISA kit (MD Bioproduct, USA). Composites were broken and suspended in guanidine hydrochloride extraction buffer (4 M GuHCl/50 mM Tris, pH 8.5 Sigma Aldrich, USA). After 30 min incubation at room temperature samples were centrifuged (13,000 rpm, 15 min, 4 °C), and the pellet was used for collagen type II quantification according manufacturer's protocol [47].

study. FTIR spectra of the hydrogel showed amino peaks of chitosan at 3400–3500 cm−1 for O\H stretching and hydrogen bonding, 1660 cm−1 and 1550 cm−1 for primary and secondary amides, and the peaks at 1116 cm−1 and 1107 cm−1 which were assigned to the saccharide structure [48–50]. Gelation temperature was determined by rheology tests through measuring the storage modulus (G′) and loss modulus (G″) versus temperature as described previously. By elevation of temperature from 10 °C to 60 °C gelation temperature of the selected CS/GP solution (CS/GP1.6) increased (Fig. 2). With elevating temperature a rapid increase of the storage modulus was observed, and the crossover of the storage modulus and the loss modulus showed the phase transition of the chitosan solution to the chitosan gel close to biological temperature (36.2 °C). The higher growth rate of the storage modulus compared to the loss modulus indicated that development of the gel structure contributed to stiffening [35].

2.10. Statistical analysis

3.2. Scanning electron microscopy (SEM)

All data were reported as mean ± SD, and for each experiment at least three samples were tested. Multifactorial one-way analysis of variance (ANOVA) was performed for the comparison of groups and results were set as significant for p b 0.05.

The porous structures of the silk fiber mesh, selected hydrogel (CS/GP1.6), and composite scaffolds were studied by means of scanning electron microscopy (SEM) as shown in Fig. 3. The average diameter of the electrospun silk fibers, produced from 10% (w/v) regenerated silk fibroin in formic acid solvent, was about 510 ± 320 nm (Fig. 3a). The inner structure of the CS/GP1.6 hydrogel is shown in Fig. 3b. The pore canal in the hydrogel was uniform and coherent. The interior morphology of this hydrogel demonstrated highly porous structure, and the pores formed an interconnecting “open-cell” structure. The structures of channels which formed by pores make swelling and deswelling fast through water convection [38,51]. Silk fibers in each of the two scaffolds were properly involved with hydrogel matrix (Fig. 3c–d). The CS/GP-D composite exhibited a structure with silk fibers interspersed within the chitosan matrix. In general the pores formed by pure chitosan structure were larger than the ones formed in silk-containing scaffolds (Fig. 3b–d).

3. Results 3.1. Rheological measurement Fig. 1 shows the gelation time of CS/GP hydrogels at 37 °C. In addition to the alterations in viscosity, the opacity of the solution changed from transparent to opaque as the samples formed a gel. The CS/GP is solution at room temperature that is 24 °C; however, this solution becomes gel at 37 °C. Gelling time appeared to display an exponential decrease with increasing GP concentration (Fig. 1). No temperature induced gelation of CS/GP preparations was observable for GP concentrations below 8% up to 12 h. Higher GP concentrations led to faster gelation. Different GP concentrations [8.3–25% (w/v)], became gel around physiological temperature but in different time durations. Chitosan combined with 10–25% (w/v) GP became gel in less than 30 min whereas 8.3% (w/v) GP resulted in longer gelation time of about 300 min. Chitosan concentration also influenced the gelation time such that higher chitosan concentrations resulted in shorter gelation times when combined with the same quantity of GP (results are not shown). In view of the fact that GP, in certain concentrations above 10%, can be detrimental to cell proliferation [26], and considering gelation time as a criterion, CS/GP1.6 was selected for the subsequent steps of the

Fig. 1. Gelation time of CS/GP solutions with different GP concentrations at 37 °C.

3.3. Scaffold characterization 3.3.1. Water uptake Fig. 4 shows the water uptake (weight of water/dry sample weight) of the samples at 37 °C. All the three samples described high water uptakes. Samples with pure chitosan hydrogel showed the highest water content (28.14 g/g), while in the case of CS/GP-D and CS/GP-L; the addition of silk fibers to the hydrogel reduced the water content to 27.02 g/g and 25.5 g/g, respectively. In general, the differences among the three

Fig. 2. Loss modulus (G′) and storage modulus (G″) of CS/GP1.6 solutions as a function of temperature. The crossover points are indicative of gel formation as G′ becomes greater than G″ in a certain temperature.

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Fig. 3. SEM images of a) silk nanofibers, b) CS/GP1.6, c) CS/GP-L, and d) SC/GP-D (arrow indicates interconnectivity and * indicates degummed fiber; diameter of silk nanofibers: 510 ± 320 nm, pore size of CS/GP1.6: 12.46 ± 1.6 μm, CS/GP-L: 7.5 ± 1.5 μm, and SC/GP-D: 9.71 ± 1.62 μm).

groups were not significant (p N 0.05). The high water content of scaffolds allows for the transport of nutrients and waste through the matrix. 3.3.2. Mechanical analysis Results of compression and indentation testing on chitosan hydrogel and two other reinforced scaffolds are shown in Fig. 5. Usage of silk fibers increased the compressive modulus of chitosan hydrogel in both methods of testing with the similar trend. Determination of the linear modulus through linear stress–strain profiles in static compression test showed that silk-containing scaffolds, CS/GP-D and CS/GP-L, were respectively 1.9 and 3.1 times stiffer than pure chitosan hydrogel with

Fig. 4. Water uptake (weight of water/dry weight) of the scaffolds. Error bars correspond to mean ± SD for n = 3.

the lowest modulus. Additionally, CS/GP-L showed a higher compressive modulus than CS/GP-D in both methods. The Young's modulus of CS/GP-L was the highest and significantly differed from chitosan hydrogel (p b 0.05) in both indentation and compression tests. There was no significant difference in modulus between the CS/GP-D composite and CS/GP hydrogel. Results of compression testing on scaffolds by means of DMA are shown in Fig. 6. Results of DMA are similar to previous findings as significantly increased moduli were observed in cases of CS/GP-D and CS/GP-L compared to CS/GP hydrogel (p b 0.05). The highest dynamic compressive modulus was that of CS/GP-L.

Fig. 5. Compressive modulus of samples via static compression and indentation test (n = 3). * indicates significant difference (p b 0.05).

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Fig. 6. Compressive modulus of samples in DMA test (n = 3). * indicates significant difference (p b 0.05).

3.3.3. Cellular assay The viability of cells grown in monolayer exposed to extracts over time is shown in Fig. 7a. The figure shows results of MTT assay of the samples extract at days 1, 7, and 14. The viability of cells exposed to silk nanofibers, CS/GP-D, and CS/GP-L extracts were almost N 90%

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at different days. These results were not significantly different when compared to the control cultured group (p N 0.05). Fig. 7a indicates that all structures did not show cytotoxic effects on cells. The results show increase in cell activity in culture media containing CS/GP1.6 extract during incubation, indicating no cytotoxic effect on cell survival and growth. These results suggest that extracts of chitosan hydrogel with 10% GP (CS/GP1.6) enhanced cell proliferation compared to the control group. Results of live/dead staining showed that chondrocytes remained alive during gelation process and in the incubation time within hydrogel. Moreover, chondrocytes maintain their spherical morphology within hydrogel (Fig. 7b–c). Fig. 8a–b shows values of produced GAG and Col II contents in harvested structures. As it is observed, the values of GAG in silk fiber containing groups in comparison with chitosan hydrogel increased slightly at days 7 and 14 (Fig. 8a). Statistical analysis demonstrated a significant difference between CS/GP hydrogel and fiber reinforced groups after 21 days. The result of Col II content is shown in Fig. 8b. The mean collagen contents of constructs showed the increase in each reinforced scaffold compared to CS/GP hydrogel at day 14 while for that of CS/GP-L was significantly higher than the two others. After 21 days of culture the collagen level reached the highest content for CS/GP-L. But that of CS/GP-D was significantly lower than those of the two other scaffolds after 21 days. These results indicate support of the chondrogenic phenotype for chondrocytes as indicated by accumulation of GAG and collagen type II in cell-seeded scaffolds over time.

Fig. 7. a) Viability of electrospun silk nanofibers, CS/GP1.6, CS/GP-D, CS/GP-L extracts and control, AO/PI stained chondrocytes within the b) CS/GP1.6 and, c) CS/GP-D after 7 days — in the case of CS/GP-D cells are behind the nanofibers sheet which is observed in red shadow (alive cell: green, dead cell: red; magnification 10×).

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Fig. 8. a) GAG content, and b) Col II content in cultured scaffolds in different days (* significant difference, ** significant increase in CS/GP scaffolds in day 21 compared to day 14).

4. Discussion In this study, pure chitosan and silk fiber–chitosan composites (SC/GP-D and SC/GP-L) were fabricated for cartilage scaffolding. Gel formation was attained at physiological pH and temperature, suggesting that such materials can be used for cell encapsulation for tissue engineering application. The fabricated composite substrates contained degummed silk fibers dispersed within the chitosan matrix (SC/GP-D) and silk nanofiber sheets laminated with the chitosan hydrogel (SC/GP-L). The results of gelation time and temperature indicated that CS/GP1.6 solution was gelled around 37 °C within a relative short term as shown in Figs. 1 and 2. As shown in Fig. 1, the gelation time decreased with elevation of GP concentration at 37 °C. This is due to the increase in molecular interactions during gelation process. Chenite et al. suggested that the mechanism of sol–gel transition within the CS/GP system includes hydrophobic interaction, molecular chain movement, hydrogen bonding, and electrostatic interaction [25,52]. Cho et al. suggested that the introduction of a GP salt to chitosan aqueous solutions clearly adjusts the effective interactions responsible for sol–gel transition including: the electrostatic attraction between chitosan and GP by means of the amino and phosphate groups, hydrophobic interactions between chitosan molecules improved by the structuring action of glycerol on water, and the increase of hydrogen bonding between chitosan chains as a result of the reduction of electrostatic

repulsion due to the basic GP addition [53]. The gelation time of CS/GP1.6 at 37 °C was approximately 30 min, which might not be appropriate for direct clinical application in mini-invasive injection but could be feasible when interacted with cells and used as a scaffold containing cells. It has been demonstrated that GP, in certain concentrations above 10%, can diminish cell proliferation [26]. However, in the case of higher concentrations this effect was examined and no cytotoxic effect was observed for exposure of 30 min. The finding of cytotoxicity caused by GP has implications for use of injectable chitosan biomaterials, since it suggests that removal or inactivation of excess GP is adequate for biocompatibility [26,54]. Based on above indications CS/GP1.6 was selected for further steps of this study. The rheological analysis (Fig. 2) showed that mechanical strength of CS/GP1.6 (10% GP) close to 37 °C was greatly improved. Increasing GP concentrations led to lowering of the gelation time from about 300 min for 8.3% GP to 6 min for 25% GP. Our solution did not show any thermoreversibility of gelation. Once the gelation occurred, the hydrogels did not reverse after cooling at room temperature, in agreement with published results [25,52]. The results of water uptake of the scaffolds showed relatively high water content of the structures while there were no significant differences between samples after silk addition (Fig. 4). The interconnecting and stable internal channel system would enable fast convective water transport inwards or outwards; making fast swelling and deswelling possibly by water convection. The swelling characteristics of highly porous hydrogels in water can be modified by altering the CS/GP ratio. The high water content values might be due to the lack of function of ion in distilled water compared to real tissue fluid within cartilage [38]. In addition, both silk fibroin and chitosan are often considered as hydrophilic biomaterials in biomedical applications, exhibiting water absorption. Proper hydrophilicity of both silk fibroin and chitosan gives rise to the large wet/dry weight ratio [33]. Addition of silk fibers leads to a reduced water content from about 28 g/g (96.6%) in CS/GP to 27 g/g (96.4%), and 25.5 g/g (96.2%) for CS/GP-D and SC/GP-L, respectively (Fig. 4). Within the composite structure, this is likely due to the fact that the swelling pressure was counterbalanced by the elastic force of the silk fibers, resulting in decrease of water uptake [44]. In other words, the higher is the swelling degree the lower is the value of compressive elastic modulus. Such effect could also be contributed by the trend that after treatment with methanol, swelling of silk fibers was limited and the water content of composites was not significantly influenced [33]. From biological view the fluid content of hyaline cartilage comprises up to 80% of the wet weight of the tissue [2] and the fact that the porosity of the hydrogel contributed to the relatively high water uptake of the scaffolds may assist to mimic structure of cartilaginous ECM. The SEM images indicated different pore sizes within the composites (Fig. 3b–d), but in general internal morphology of all the three scaffolds demonstrated highly porous structures. The interconnecting and stable internal channel system enables fast convective water transport inwards or outwards, leading to fast swelling and deswelling by water convection [38]. In cartilage tissue engineering by means of different structures a wide range of pore sizes between less than 10 μm to almost 500 μm has been reported [46]. Although some studies gave preference to a small pore size of less than 50 μm, others did not find a significant difference between 20 and 83 μm pore diameters in their experiments [4]. The secretion of chondrogenic factors has been observed even in the scaffolds with the pore size smaller than 10 μm [55] comparable to our findings. The differences among the pore sizes were due to silk addition which resulted in the formation of fibers with smaller diameters and tighter network structure in porous composites. It can be speculated that slight reduction in water content in composites is due to changing the pore size and compactness of composites compared to pure chitosan hydrogel.

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Proper strength and compressive modulus is required for scaffolds in order to retain their original shapes while maintaining adequate pore space. Results of mechanical test revealed that mechanical properties were enhanced by adding silk fibers to the scaffolds. As expected, the elastic modulus of the hydrogel under the wet condition was lower than the two other scaffolds (Figs. 5, 6). Addition of silk fiber meshes to the hydrogel through lamination (SC/GP-L) significantly improved the compressive modulus under wet condition. The fibrous mesh reinforced the structure and reduced water uptake of the matrix, resulting in a laminated structure with enhanced mechanical properties. This might be due to the more compact structure of SC/GP-L compared to the others (Fig. 3c). Moreover the surface to volume ratio of nanostructure of fibrous sheets is higher which influences binding interactions between molecules and subsequently enhances elastic modulus. In the case of SC/GP-D, due to low interfacial interaction, the addition of degummed silk fibers to the hydrogel improved the modulus of the SC/GP-D less in comparison with SC/GP-L. Some studies have demonstrated noticeable effect of interfacial interactions of the matrix and additional components on viscoelastic properties of the composites [10,44,56]. Xu et al. showed that lamination increased the modulus dramatically compared to the fiber mesh or hydrogel under the wet condition [44]. Li et al. reported that the nano-hydroxyapatite/collagen/PLLA composite reinforced by chitin fibers linked by DCC not only possessed higher mechanical strength but also kept better mechanical strength during degradation than the reinforced scaffold without linking [56]. Wang et al. demonstrated that addition of collagen resulted in a stiffer material, compared to pure chitosan [26]. Various studies suggested different values of compressive modulus for pure chitosan hydrogel including 6.3 kPa [26], 1.5 kPa [21], and 1.5–5 kPa [23]. The differences among the data including results of this study may be due to the usage of different materials and structure of the designed scaffolds [33]. Although we enhanced mechanical properties of chitosan hydrogel using silk fiber, the compressive modulus of both reinforced composites were lower than the compressive modulus of articular cartilage, which is about 0.5–1 MPa [5]. However, the enhancement due to the usage of silk fibers was considerable. It should also be mentioned that such materials are not used to replace the total cartilage. They are usually employed to perform in local injuries in conjunction with cartilaginous tissue. This is the reason that many studies use hydrogel based materials. Compression modulus is a material property that is strongly dependent on solid content. Composites prepared in this study had a water content of approximately 97% while the cartilage is composed of about 70–80% tissue fluid [2,4]. It is possible to improve solid content and thus the compressive modulus. The cytocompatibility of biomaterials is an important consideration for cell encapsulation, since cell viability must be maintained during gel fabrication; and upon implantation toxicity to surrounding cells has to be low. Both silk fibers and chitosan are well known as biocompatible and biodegradable biopolymers [30–32]. They are naturally derived materials that have been widely studied due to their ability to support cell growth and to integrate with surrounding ECM [30,31]. Previous studies have demonstrated high cell viability in pure chitosan gels induced by GP after seeding chondrocytes [21,24]. Most cells must adhere to a surface to be able to grow [6]. Results of MTT assay (Fig. 7a) confirmed maintaining viability of cells as previously been suggested that CS/GP hydrogels can be used as cell carriers for chondrocytes [25,53]. Published results have confirmed biocompatibility of CS/GP hydrogels in vivo, through higher deacetylation degree of chitosan and consequently lesser inflammatory reactions [35]. In current study, the deacetylation degree of chitosan was above 84%. Results of cytotoxicity test showed that the fabricated scaffolds were not harmful to the cultured cells and not only kept the cell viability at normal level, but also enhanced cell proliferation. Results of proteoglycan and collagen type II assay over time indicated support of the chondrogenic phenotype for chondrocytes. Moreover, the scaffold should support phenotypic stability of embedded chondrocytes,

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necessary for synthesis of vital ECM proteins such as collagen type II and proteoglycans (Fig. 8). The content of proteoglycan and collagen type II accumulated in CS/GP-L construct after 21 days was higher than other scaffolds in different days, describing correlation between enhanced mechanical properties and deposition of proteoglycan and Col II [35]. These findings encourage the conception that changing structural features of a scaffold may be a viable strategy to maintain chondrogenic induction. Both composite scaffolds, CS/GP-D and CS/GP-L, were mechanically enhanced but the sandwich type (CS/GP-L) was mechanically better which can be potentially more suitable for cartilage reconstruction. However, the other reinforced scaffold might be more appropriate for in situ-forming surgery. 5. Conclusion Chitosan hydrogel composites containing two forms of silk fiber were fabricated by initiating gelation GP and temperature. This process was performed at physiological pH and temperature, such that living cells could be incorporated directly into the hydrogel matrix. The presence of silk fibers within the composite structure was associated with stiffening of the matrix. The chitosan-reinforced composite hydrogels are potentially applicable in regenerative medicine to provide structural support to the reconstructing region in load-bearing cartilage tissue. The mechanical characteristics of these composites may be further improved by the number of fiber layers in the laminate, combination of both degummed fibers and nanofiber sheets, crosslinking of hydrogels and silk fibers by nontoxic agent during gelation, content and surface treatment of fibers, control of water uptake of the hydrogel, and fiber orientation. References [1] R.P. Lanza, R. Langer, J. Vacanti, Principles of Tissue Engineering, Academic Press, New York, 1997. [2] J.S. Temenoff, A.G. Mikos, Biomaterials 21 (2000) 431–440. [3] Q.P. Pham, U. Sharma, A.G. Mikos, Tissue Eng. 12 (2006) 1197–1211. [4] T. Aigner, J. Stove, Adv. Drug Deliv. Rev. 55 (2003) 1569–1593. [5] A. Svensson, E. Nicklasson, T. Harrah, B. Panilaitis, D.L. Kaplan, M. Brittbergc, P. Gatenholm, Biomaterials 26 (2005) 419–431. [6] R.S. Tıgli, M. Gumusderelioglu, Biotechnol. Bioeng. 104 (2009) 601–610. [7] D. Eyrich, F. Brandl, B. Appel, H. Wiese, G. Maier, M. Wenzel, R. Staudenmaier, A. Goepferich, T. Blunk, Biomaterials 28 (2007) 55–65. [8] H.S. Baek, Y.H. Park, C. Seok Ki, J.-C. Park, D. Kyun Rah, Surf. Coat. Technol. 202 (2008) 5794–5797. [9] H. Tan, J. Wu, L. Lao, C. Gao, Acta Biomater. 5 (2009) 328–337. [10] Y.-K. Seo, J.-K. Park, Biotechnol. Bioprocess Eng. 15 (2010) 527–533. [11] H.-J. Yen, C.-S. Tseng, S.-h. Hsu, C.-L. Tsai, Biomed. Microdevices 11 (2009) 615–624. [12] T.-W. Chung, Y.-L. Chang, J. Mater. Sci. Mater. Med. 21 (2010) 1343–1351. [13] Y.Q. Zhang, W.L. Zhou, W.D. Shen, Y.H. Chen, X.M. Zha, K. Shirai, K. Kiguchi, J. Biotechnol. 120 (2005) 315–326. [14] C. Vepari, D.L. Kaplan, Prog. Polym. Sci. 32 (2007) 991–1007. [15] J.G. Hardy, L.M. Romer, T.R. Scheibel, Polymer 49 (2008) 4309–4327. [16] Y. Cao, B. Wang, Int. J. Mol. Sci. 10 (2009) 1514–1524. [17] D.N. Rockwood, E.S. Seok, S.-h. Park, J.A. Kluge, W. Grayson, S. Bhumiratana, R. Rajkhowa, X. Wang, S.J. Kim, G. Vunjak-novakovic, D.L. Kaplan, Acta Biomater. 7 (2011) 144–151. [18] Y. Wang, D.J. Blasioli, H.-j. Kim, H.S. Kim, D.L. Kaplan, Biomaterials 27 (2006) 4434–4442. [19] M.L. Lovett, C.M. Cannizzaro, G. Vunjak-novakovic, D.L. Kaplan, Biomaterials 29 (2008) 4650–4657. [20] I.-Y. Kim, S.-J. Seo, H.-S. Moon, M.-K. Yoo, I.-Y. Park, B.-C. Kim, C.-S. Cho, Biotechnol. Adv. 26 (2008) 1–21. [21] C.D. Hoemann, M. Hurtig, E. Rossomacha, J. Sun, A. Chevrier, M.S. Shive, M.D. Buschmann, J. Bone Joint Surg. Br. 87 (2005) 2671–2686. [22] J.K.F. Suh, H.W.T. Matthew, Biomaterials 21 (2000) 2589–2598. [23] T. Hao, N. Wen, J.K. Cao, H.B. Wang, S.H. Lü, T. Liu, Q.X. Lin, C.M. Duan, C.Y. Wang, Osteoarthr. Cartil. 18 (2010) 257–265. [24] G.R. Ragetly, D.J. Griffon, H.-B. Lee, L.P. Fredericks, W. Gordon-Evans, Y.S. Chang, Acta Biomater. 6 (2010) 1430–1436. [25] A. Chenite, C. Chaput, D. Wang, C. Combes, M.D. Buschmann, C.D. Hoemann, J.C. Leroux, B.L. Atkinson, F. Binette, A. Selmani, Biomaterials 21 (2000) 2155–2161. [26] L. Wang, J.P. Stegemann, Biomaterials 31 (2010) 3976–3985. [27] Q.S. Zhao, Q.X. Ji, K. Xing, X.Y. Li, C.S. Liu, X.G. Chen, Carbohydr. Polym. 76 (2009) 410–416. [28] S. Kim, S.K. Nishimoto, J.D. Bumgardner, W.O. Haggard, M.W. Gaber, Y. Yang, Biomaterials 31 (2010) 4157–4166.

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