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Fabrication of keratin-silica hydrogel for biomedical applications
Materials Science and Engineering C 66 (2016) 178–184
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Fabrication of keratin-silica hydrogel for biomedical applications Prachi Kakkar, Balaraman Madhan ⁎ Central Leather Research Institute, Council of Scientific and Industrial Research, Chennai, Tamil Nadu, India
a r t i c l e
i n f o
Article history: Received 28 January 2016 Received in revised form 28 March 2016 Accepted 18 April 2016 Available online 20 April 2016 Keywords: Hydrogel Textural analysis Biocompatibility Biomaterial Biomedical application Sol–gel
a b s t r a c t In the recent past, keratin has been fabricated into different forms of biomaterials like scaffold, gel, sponge, film etc. In lieu of the myriad advantages of the hydrogels for biomedical applications, a keratin-silica hydrogel was fabricated using tetraethyl orthosilicate (TEOS). Textural analysis shed light on the physical properties of the fabricated hydrogel, inturn enabling the optimization of the hydrogel. The optimized keratin-silica hydrogel was found to exhibit instant springiness, optimum hardness, with ease of spreadability. Moreover, the hydrogel showed excellent swelling with highly porous microarchitecture. MTT assay and DAPI staining revealed that keratin-silica hydrogel was biocompatible with fibroblast cells. Collectively, these properties make the fabricated keratin-silica hydrogel, a suitable dressing material for biomedical applications. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Hydrogels are one of the most established biomaterials because of their close resemblance to the in vivo histo-architecture of the natural extracellular matrix (ECM) [1,2]. Hydrogels are crosslinked 3D networks comprising of either natural or synthetic polymers. It has been observed that synthetic hydrogels have more controllable physical properties and are more reproducible, but show poor biocompatibility, lack of bioactivity and little resemblance to the natural environment [3–6]. Hence use of biopolymer such as collagen or keratin in the fabricated biomaterial enhances the biocompatibility and regeneration of normal at the site of applications [7–9]. Sol–gel technique is well-known and practiced to synthesize novel hybrid biocompatible gels for a wide range of applications [10,11]. It is interesting to note that sol–gel process is in-expensive and the gels so produced are non-toxic [12,13], involving simultaneous hydrolysis and condensation of alkoxide or salt. The sol–gel processes have been utilized to produce bioactive coatings, powders, and substrates that are suitable to be used as implants and sensors [14,15]. Advantage of employing sol–gel method to produce hybrid materials is that it combines both inorganic and organic properties. In addition, this method offers the possibility of obtaining homogeneous hybrid materials under low temperature, which further provides the scope of incorporating a variety of compounds [16,17]. Hydrogels have shown promising applications in the field of pharmacy and medicine, apart from their popular use in the non-medical sector [18]. This can partially be attributed to the absorption capacity ⁎ Corresponding author. E-mail addresses: [email protected], [email protected] (B. Madhan).
http://dx.doi.org/10.1016/j.msec.2016.04.067 0928-4931/© 2016 Elsevier B.V. All rights reserved.
and retention of mechanical strength of these hydrogels. Hydrogels are preferred because they have high water content, with soft and rubbery surface, mimicking human tissue viz., muscles, tendons and cartilage [19,20]. These characteristic features make them potential candidate for transdermal drug delivery systems. The hydrogels liquefy necrotic tissue on wound surface and prevents loss of body fluids. Moreover they are non-adherent and the dressing can be removed without trauma to the wound bed [21]. The feasibility of using natural materials including polysaccharides and proteins for the fabrication of hydrogels has been well-demonstrated. Among these materials, collagen, chitosan, keratin, gelatin, elastin, fibrin hyaluronic acid etc. exhibit prominent bioactivity in biomedical applications [22–24]. Keratins are naturally derived proteins that can be fabricated into several biomaterial forms including hydrogels [25]. Although keratins are considered to be intracellular cytoskeleton proteins belonging to the family of intermediate filaments [26], they contain the cell adhesion motifs LDV (leu-asp-val) and RGD (arg-gly-asp) [27], enabling cell attachment and inturn supporting cell growth and development [28]. Keratins in the form of coatings, fibres and films have also shown promising cytocompatibility (Reichl 2009; Rouse and Van Dyke 2010). Keratin based hydrogels have been reported to be neuro-inductive and capable of facilitating regeneration in a peripheral nerve injury [29]. Apart from this, keratin hydrogels are known to enhance wound healing by the activation of keratinocytes in the wound bed. These hydrogels also acts as a protective barrier, absorb excess wound exudates, and maintain a moist environment, which in turn helps in pain reduction [30]. In particular, silica-based hydrogels are well recognized as suitable matrices for biomaterial development. Researchers have extensively worked on sol–gel processed silica and found their potential as drug
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Fig. 1. Schematic representation of texture analysis to calculate i) Hardness (peak height), ii) Cohesiveness (B1 + B2)/(A1 + A2), iii) Adhesiveness (C), iv) Compressive strength (peak stress/peak strain), v) Resilience (A2/A1).
carriers [31–33]. It has been established that sol–gel produced pure silica are biocompatible and even enhance the formation of fibroblast or osteoblast leading to increased collagen production [12,34–36]. Our group has established the use of sol–gel processed collagen – silica composite materials for wound healing applications [37,38]. Varying the amount of silica could tune the mechanical property of the collagen scaffolds obtained [39]. Gorji, Allahgholi Ghasri, Fazaeli and Niksirat [13] used tetraethyl orthosilicate (TEOS) as a precursor for preparation of silica gels. It is a well-known fact that the hydrolysing agent in the TEOS-derived sols is water. Usually acid hydrolysis of TEOS is employed to prepare aggregative stable spinnable sols. Inorganic acids like hydrochloric acid (HCl), nitric acid (HNO3), etc, are frequently used as acid catalysts, but occasionally organic acids like acetic acid is also being used [10,40,41]. There are two different ways through which sol–gel synthesis can be carried out i.e. single-stage (acid) or two-stage (acid/acid) hydrolysis of TEOS [42–44]. In both the cases, first stage is often carried out with scarcity of water. TEOS acts as a crosslinking agent for keratin hydrogel formation. TEOS has a remarkable property to produce silicon dioxide (SiO2) when mixed with water in presence of a catalyst which can be an acid or a base as per Stober reaction which was discovered by Werner Stober. This process takes place via LaMer model or nucleation model comprising of two steps. It begins with nucleation (fast process) followed by a slow process of particle growth. In a typical process, silicon
Fig. 2. Representation of keratin-silica hydrogel withstanding tube inversion test (4:1w/w of silica:keratin).
alkoxide viz., tetraethyl orthosilicate (TEOS) is hydrolyzed by water in presence of an acidic catalyst. These reactive silanol (Si–OH) groups of the hydrolyzed silicon alkoxide go through condensation reactions which depend on the pH and temperature of the solution, to form siloxane (Si–O–Si) bonds, producing a 3D porous gel structure [45]. Earlier we have established that keratin can be effective for proliferation of fibroblast cells [7]. Moreover, different forms of keratin based materials are required as effective materials for soft tissue engineering applications such as wound healing. This paper deals with the development of a novel keratin hydrogel system by combining keratin with silica, where the sol–gel transition of silica had been advantageously used for the establishment of hydrogel. 2. Materials and methods 2.1. Materials Keratin was extracted from bovine hooves, TEOS (C8H20O4Si) was obtained from Alfa Aesar (United States), HCl was obtained from Fisher
Table 1 Gelation time of keratin-silica hydrogel at varying concentration of silica. Ratio of silica:keratin (w/w) 0.4:1 0.8:1 1.2:1 1.6:1 2:1 2.4:1 2.8:1 3.2:1 3.6:1 4:1 4.4:1 4.8:1 5.2:1 5.6:1
Gelation time No gelation No gelation 2h 1 h 30 min 1 h 10 min 1h 40 min 20 min 15 min 10 min 8 min 7 min 6 min 5 min
Fig. 3. Hardness of keratin-silica hydrogels with varying concentrations of silica.
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Fig. 4. Cohesiveness of keratin-silica hydrogels with varying concentrations of silica.
Scientific (United Kingdom). For in vitro analysis, DMEM and MTT were purchased from Sigma-Aldrich (Sigma-Aldrich Chemicals Pvt. Ltd., India), FBS from Gibco® (Invitrogen, USA) and 4′,6-diamidino-2phenylindole (DAPI) from HiMedia (HiMedia Laboratories Pvt. Ltd., India). 2.2. Methods 2.2.1. Fabrication of keratin-silica hydrogel 2.2.1.1. Preparation of keratin solution. Keratin from bovine hooves was extracted as per procedure detailed elsewhere [7]. Briefly, 10 g of defatted raw bovine hooves was subjected to reduction using 7 M Urea, 6 g sodium dodecyl sulphate and 15 ml 2-mercaptoethanol at 60 °C for 12–14 h. The supernatant was then dialyzed against water to obtain reduced keratin. Extracted lyophilized (powdered) hoof keratin was solubilized in PBS (pH 7.4) to make a stock solution with the concentration of 5 mg/ml. 2.2.1.2. Preparation of acidified TEOS. For the fabrication of hydrogels, TEOS has been used as a precursor of silica. It was acidified using 0.1 N HCl to obtain a pH close to 2. It is known that sol–gel transition of silica in TEOS can be achieved by elevating the pH from 2 to 7. For this, keratin stock solution (pH 7.4) was drop-wise added to TEOS to obtain keratinsilica hydrogels by varying silica:keratin ratio (w/w) from 0.4:1 to 5.6:1. 2.2.2. Tube inversion test For the preparation of keratin-silica hydrogels with different concentrations of TEOS, the solution mixture was poured into a glass vial. To check the formation of stable gel, test tube was inverted upside down every 1 min and time for gelation was noted down in each case.
Fig. 5. Adhesiveness of hydrogels with varying concentrations of silica.
Fig. 6. Resilience of hydrogels with varying concentrations of silica.
Tube inversion test is used to determine the effect of varied concentration of silica on gelation time of keratin. 2.2.3. Texture analysis It is necessary to evaluate the mechanical properties of hydrogel in order to determine whether the hydrogels will be able to maintain their physical structure at the site of application. Keratin-silica hydrogels with varying concentrations of TEOS were analyzed for their textural properties using Texture Analyzer Pro CT V1.4 Build 17 (Brookfield Engineering Labs, Inc). The respective samples (with the diameter of 3.3 cm), taken in triplicate under confined conditions were compressed to 50% deformation at room temperature, using a cylindrical probe with the diameter of 12.7 mm. Hardness, adhesiveness, cohesiveness, resilience and compressive strength of keratin-silica hydrogel system were computed using TexturePro CT software as shown in Fig. 1. 2.2.4. FTIR spectroscopy The optimized keratin-silica hydrogel and its constituents (i.e. keratin and TEOS) were analyzed using JASCO FT/IR-4200 A spectrometer. FTIR spectra were recorded in the range of 4000 to 500 cm−1 with the resolution of 4 cm−1 at 32 scans/sample using conventional KBr pellet method. In brief, the lyophilized (powdered) samples were mixed with KBr and formed into pellets before recording the spectrum. 2.2.5. Hydrogel morphology The surface morphology, structural integrity and interconnectivity of the pores in the optimized hydrogel were observed using Ziess SUPRA 55VP field emission scanning electron microscope (FESEM) at a magnification of 5000×. The samples for analysis were prepared by
Fig. 7. Compressive modulus of hydrogels with varying concentrations of silica.
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The final concentration of DAPI used for staining was 1 μg/ml which was adjusted using Dimethyl sulfoxide (DMSO). For this, the cellseeded control (polystyrene plate) and the hydrogel were stained with DAPI solution for 20 min in dark and then washed thrice with PBS to visualize under fluorescence microscope (Olympus CKX41). 3. Results and discussions 3.1. Fabrication of keratin-silica hydrogels
Fig. 8. FTIR spectra of keratin, TEOS and hydrogel.
sputter coating the scaffold surface with a thin layer of gold. The surface of the samples was scanned at 5 kV. 2.2.6. Swelling study The optimized keratin-silica hydrogel was freeze dried to obtain its dry weight (Wd). It was then immersed in 10 ml PBS (pH 7.4) till complete saturation. The weight of the swollen hydrogel (Ws) was measured after wiping its surface and the swelling ratio was calculated by the following equation: Swelling ð%Þ ¼ ½ðWs −Wd Þ=Wd 100
2.2.7. Biocompatibility (MTT) assay NIH 3T3 fibroblast cell lines obtained from National Centre for Cell Science (NCCS), Pune, India, were used for the cell viability test. The cells were cultured and allowed to passage for 3 times. The passaged cells were seeded optimized on keratin-silica hydrogel at a density of 1 × 105 cells/well. The hydrogel was incubated at 37 °C with 5% CO2 for 3 h. To this, DMEM containing 10% FBS, penicillin (50 unit/ml) and streptomycin (50 unit/ml) were added and incubated at 37 °C with 5% CO2. The medium was changed every 2 days and the hydrogels were observed under microscope. After 1, 3 and 5 days of culture, cell viability was measured quantitatively using MTT assay, which is a colorimetric technique for determining the number of viable cells, that relies on the conversion of MTT to MTT formazan by the enzyme mitochondrial reductase of the viable cells [46]. The cell extracts from the hydrogel were mixed with MTT solution and the absorbance at 550 nm was measured on a microplate reader (Biotek, USA). The cell-seeded keratin-silica hydrogel was also stained with DAPI to observe the attachment of NIH 3 T3 fibroblast cells with the hydrogel.
Keratin-silica hydrogels were prepared by adding different concentrations of TEOS ranging from 0.4 to 5.6 (w/w) of keratin. 5 mg/ml keratin stock solution was used for the preparation of keratin-silica hydrogels by varying the amount of keratin. Time taken for the solution to turn into gel was recorded using tube inversion method and termed as gelation time as presented in Table 1. As a representation, one of the keratin-silica hydrogel system is presented in Fig. 2. It is clearly observed that the gelation time of the hydrogels reduced with increasing silica concentration in the hydrogel. 3.2. Texture analysis of keratin-silica hydrogel Texture analysis is a fast and reproducible method, providing a more straight-forward approach to characterize the hydrogels. The parameters like hardness, adhesiveness, cohesiveness, resilience and compressive strength provide insight on the spreadability of the product, ease of product removal [19], ability of a product to regain its original shape etc. Syverud, Pettersen, Draget and Chinga-Carrasco [47] performed compression of hydrogels using texture analyzer to determine the compressive strength/young's modulus. To evaluate the mechanical properties of hydrogels, texture profile analysis (TPA) had been proposed as a suitable and reliable method by Jones, Woolfson and Brown [48]. Later, Hurler, Engesland, Poorahmary Kermany and Škalko-Basnet [49] proposed a simplified method using texture analyzer which excludes time consuming rheological characterization. Here, the keratin-silica hydrogels are represented in terms of increasing amount of silica. 3.2.1. Hardness Determining the hardness of the hydrogel is a measure of applicability of the hydrogel onto the skin. It can be defined as maximum force required to attain a given (here 50%) deformation. This is calculated by the peak of force during the first compression cycle [50]. Lower the gel hardness lower is the force required to recover the gel from the container and hence more ease of applicability. But, with decrease in hardness the retention time of the gel on the site (for example, wound) of application would reduce. Therefore, the hydrogel should possess an optimum hardness. It can be observed from Fig. 3, as the concentration
Fig. 9. Scanning electron micrographs of a) keratin-silica hydrogel with 4.4:1 w/w of silica:keratin and b) pure keratin at 5000×.
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of silica increases (from 1.2 to 5.6, w/w), hardness of the hydrogel increases from 46 ± 3 g to 184 ± 6 g. As evidenced from Fig. 3, the hardness of the hydrogels with low concentration of silica i.e., from 1.2:1 to 3.2:1 (w/w, silica:keratin) was low and it significantly increased from 4:1 (P b 0.001) onwards. Hydrogels with 5.2 and 5.6 ratio of silica:keratin (w/w) showed maximum hardness of 179 g and 184 g respectively. In terms of recovery of the hydrogels, all the prepared hydrogels showed no significant difference. 3.2.2. Cohesiveness Cohesiveness is the measure of how well a hydrogel can withstand a second deformation compared to that of the first one. In other words, it is the tendency of the hydrogel to adhere to itself. Cohesiveness is calculated using the following formula: Cohesiveness ¼
ðB1 þ B2Þ ðA1 þ A2Þ
Where, (B1 + B2) is the area under second compression cycle and (A1 + A2) is the area under first compression cycle, as shown in Fig. 1. Cohesiveness determines the reconstruction ability of the hydrogel following its application [50,51]. Higher the cohesiveness, more is the structural recovery [52,53]. Fig. 4 shows cohesiveness increased from 0.42 ± 0.04 to 1.49 ± 0.052 as the silica concentration in the keratinsilica hydrogel increased from 0.4 to 5.6 mg/ml of keratin. [19] also showed similar cohesiveness results for methacrylate based hydrogels. As evident from Fig. 4, cohesiveness of hydrogels with silica:keratin content of 1.2:1 was not significantly different from that of 5.6:1 (P N 0.05). This shows that all the hydrogels could withstand second deformation. It also means that the behaviour of the hydrogels during second compression cycle was similar to that of the first one. 3.2.3. Adhesiveness Adhesiveness represents the work required to overcome the attractive forces between the surface of the hydrogel and that of the probe. It is measured by calculating the force applied by the hydrogel to the probe when it redraws after compression. Adhesiveness was calculated as the negative force area subsequent to the first compression cycle, represented as (C) in Fig. 1. This parameter also relates to the retention time of the hydrogel onto the site of application for the desired period of time. More the retention period of the hydrogel, shorter is the treatment period resulting into improved patient compliance [50]. As evident from Fig. 5, adhesiveness of the hydrogels significantly increased from 3.2:1 to 5.6:1 (w/w) of silica:keratin. Adhesiveness of hydrogel with 4:1 ratio of silica:keratin was observed to be 0.48 ± 0.035 mJ, while that of hydrogels with 4.4:1, 4.8:1 and 5.2:1 (w/w) of silica:keratin was 0.50 ± 0.04 mJ, 0.55 ± 0.042 mJ and 0.58 ± 0.039 mJ respectively. Hence, adhesiveness of 4:1 (silica:keratin) hydrogel was not significantly different from that of 4.4:1, 4.8:1 and 5.2:1
Fig. 10. Swelling kinetics of optimized keratin-silica hydrogel.
hydrogels (P N 0.05). Whereas, 5.6:1 (silica:keratin) hydrogel showed significantly higher adhesiveness (P b 0.01). This implies that 5.6:1 (silica:keratin) hydrogel could have more retention time on the site of application. In other words, it also means that this hydrogel tends to stick to the container, which can pose difficulty while handling it. This observation is in-line with the previous result (as shown in Fig. 3), showing highest hardness among all the fabricated hydrogels, thereby reducing the ease of spreadability of the hydrogel (silica:keratin 5.6:1) on the wound site. Therefore, hydrogels with 4:1 to 5.2:1 (w/w) of silica:keratin are favourable compared to the rest. 3.2.4. Resilience Resilience is how well a substance fights to regain its original position after being compressed. It can also be considered as instant springiness. It is calculated using the following formula:
Resilience ¼
A2 A1
Where, A2 is the area during the withdrawal of first compression cycle and A2 is the area during the compression of first compression cycle [54]. It can be deduced from Fig. 6, that resilience of the hydrogels increased with increasing concentration of silica in the hydrogels. But, after a certain concentration, it again started decreasing. Resilience of 2.4:1, 2.8:1 and 3.2:1 (w/w) of silica:keratin hydrogels is similar to each other (P N 0.05). Moreover, 3.6:1 hydrogel showed a resilience of 0.68 ± 0.33, while that of 4:1 and 4.4:1 hydrogel was found to be 0.68 ± 0.35 and 0.69 ± 0.37 respectively. Therefore, the resilience of 3.6:1 (silica:keratin) hydrogel was not significantly different from that of 4:1 and 4.4:1 (P N 0.05). However, as the concentration of silica further increased from 4.8:1 to 5.6:1, the resilience dropped. This could be due to increased silica content in the hydrogels which might impart brittleness, thereby reducing the instant springiness. It is interesting to note that the resilience of 5.6:1 hydrogel was similar to that of 4.8:1 hydrogel. Therefore, based on the resilience, 3.6:1 to 4.4:1 hydrogels were found to be favourable for the current study. 3.2.5. Compressive strength Compressive strength or compressive modulus was calculated as the ratio of peak stress to the peak strain during the first compression cycle [55]. Fig. 7 shows that compressive moduli of 1.2:1, 1.6:1, 2:1, 2.4:1 and 2.8:1 silica:keratin (w/w) hydrogels was not significantly different (P N 0.05). But the compressive modulus of 3.2:1 hydrogel was found to be significantly higher than that of 1.2:1 (P b 0.001). Further, 4.4:1 hydrogel possessed appreciably high compressive modulus of 4.53 ± 0.42 kPa which was not significantly different from that of 4.8:1, 5.2:1 and 5.6:1 hydrogels.
Fig. 11. Biocompatibility assay of keratin-silica hydrogel.
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Fig. 12. Fluorescence micrographs of DAPI stained, 3T3 fibroblast cells-seeded a) polystyrene plate (control) and b) optimized keratin-silica hydrogel.
Based on the texture analysis, it can be deduced that keratin-silica hydrogel with 4.4:1 silica:keratin (w/w) could be the optimum hydrogel showing easy spreadability, good retention on the site of application, considerable springiness and appreciably high mechanical strength. This optimized hydrogel was further characterized using different analytical tools. 3.3. FTIR spectroscopy of keratin-silica hydrogel system The FTIR spectra of the keratin (K), TEOS and optimized keratinsilica hydrogel (4.4:1) are shown in Fig. 8. In keratin composite scaffold, the peaks observed near 3280 cm− 1 are characteristic of amide A, whereas amide I falls in the range of 1600–1700 cm−1, amide II near 1520 cm− 1and amide III in the range of 1220–1300 cm− 1 [56]. The peak of amide A is due to stretching vibration of N–H bonds. The peak of amide I falling at 1654 cm−1 is caused by stretching vibration of C– O bonds. N–H bending and C–H stretching vibration which is responsible for amide II band is observed at 1542 cm−1. At 1242 cm−1 the sharp peak of amide III is observed that is due to the phase combination of C–N stretching and N–H in plane bending and partial addition from C = C bending and C–C stretching vibration. In case of TEOS, asymmetric vibrations of Si–O were observed at 1085 cm−1. This peak was also clearly seen in keratin-silica hydrogel, but with lower intensity. Peaks at 945 cm−1 and 796 cm−1 indicated asymmetric vibrations of Si–OH bond and symmetric vibration of Si–O bond respectively. The hydrogel was observed to possess a shoulder band around 945 cm−1 and a low intensity band at 796 cm−1. Therefore, it could be deduced that keratin-silica hydrogel showed the characteristic properties of its constituents i.e. keratin and TEOS. In addition to this, CH3 peak at 2980 cm−1 and CH2 peak at 2930 cm−1 which are attribute to the unreacted TEOS in silica, were absent in the hydrogel. This affirms that the reaction of TEOS forming silica was complete. There was no significant appearance/disappearance of any peak in the spectrum of hydrogel as the hydrogel is most likely to be formed by the physical interactions. 3.4. Morphology of optimized keratin-silica hydrogel SEM analysis revealed that the optimized keratin-silica hydrogel possessed porous micro-architecture as shown in Fig. 9. It was seen to be a homogenous matrix with uniform distribution of its constituents. At 5000×, the prominent needle-like morphology was portrayed by silica produced by TEOS. Silica needles were completely embedded into the keratin hydrogel. 3.5. Swelling behaviour of optimized keratin-silica hydrogel Dried optimized hydrogel was immersed in PBS (pH 7.4) to understand the swelling behaviour. The samples (in triplicate) were immersed in PBS till complete saturation. When placed in water/solvent,
a polymeric hydrogel swells without dissolving in it. The solvent continues to penetrate the polymeric network of the hydrogel till the equilibrium between osmotic and elasticity forces is achieved [57]. Parameters influencing the swelling of hydrogels are network density, solvent nature and polymer solvent interaction [58]. The optimized keratin-silica hydrogel was analyzed (in triplicate) for its swelling ratio. Fig. 10 shows the kinetics of hydrogel swelling which was calculated as:
Swelling ratio ¼
Wt−W0 W0
Where, Wt refers to the weight of the hydrated hydrogel at time t, while W0 refers to weight of the dried hydrogel. The averaged swelling of the keratin-silica hydrogel was found to be 540 ± 28.97%, which is appreciably good for biomedical applications such as wound healing. The swelling kinetics of the keratin-silica hydrogel (dried form) indicated that it got saturated around 4 h.
3.6. Biocompatibility assay The NIH 3T3 cells cultured on the optimized keratin-silica hydrogels which were monitored continuously for 5 days. The untreated culture plate wells (polystyrene) comprising of cells were treated as control. The cell viability results of MTT assay (Fig. 11) performed after 1, 3 and 5 days showed no significant difference (P N 0.05) between the test sample and control, and after 7 days the cell viability of the composite scaffold was similar to that of collagen scaffold. The optimized keratin-silica hydrogel was stained with DAPI, which is a fluorescent dye and can penetrate the intact cell membrane. Fig. 12 shows the images of DAPI stained, cell-seeded optimized hydrogel and the polystyrene plate (control). It can be observed that the hydrogel (Fig. 12b) showed appreciable proliferation and infiltration of the fibroblast cells into the hydrogel matrix. There was no significant difference between the control (Fig. 12a) and the optimized hydrogel. These results corroborate with the MTT assay.
4. Conclusions Keratin from bovine hooves could be successfully fabricated into hydrogel by using TEOS as a precursor of silica. The optimized keratinsilica hydrogel showed considerable hardness, while maintaining ease of applicability. Moreover, it displayed instant springiness, high compressive modulus and biocompatibility. Together, these characteristics show that the optimized hydrogel could be used as a biomaterial for biomedical applications, especially for wound healing.
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Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Fabrication of keratin-silica hydrogel for biomedical applications Prachi Kakkar, Balaraman Madhan ⁎ Central Leather Research Institute, Council of Scientific and Industrial Research, Chennai, Tamil Nadu, India
a r t i c l e
i n f o
Article history: Received 28 January 2016 Received in revised form 28 March 2016 Accepted 18 April 2016 Available online 20 April 2016 Keywords: Hydrogel Textural analysis Biocompatibility Biomaterial Biomedical application Sol–gel
a b s t r a c t In the recent past, keratin has been fabricated into different forms of biomaterials like scaffold, gel, sponge, film etc. In lieu of the myriad advantages of the hydrogels for biomedical applications, a keratin-silica hydrogel was fabricated using tetraethyl orthosilicate (TEOS). Textural analysis shed light on the physical properties of the fabricated hydrogel, inturn enabling the optimization of the hydrogel. The optimized keratin-silica hydrogel was found to exhibit instant springiness, optimum hardness, with ease of spreadability. Moreover, the hydrogel showed excellent swelling with highly porous microarchitecture. MTT assay and DAPI staining revealed that keratin-silica hydrogel was biocompatible with fibroblast cells. Collectively, these properties make the fabricated keratin-silica hydrogel, a suitable dressing material for biomedical applications. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Hydrogels are one of the most established biomaterials because of their close resemblance to the in vivo histo-architecture of the natural extracellular matrix (ECM) [1,2]. Hydrogels are crosslinked 3D networks comprising of either natural or synthetic polymers. It has been observed that synthetic hydrogels have more controllable physical properties and are more reproducible, but show poor biocompatibility, lack of bioactivity and little resemblance to the natural environment [3–6]. Hence use of biopolymer such as collagen or keratin in the fabricated biomaterial enhances the biocompatibility and regeneration of normal at the site of applications [7–9]. Sol–gel technique is well-known and practiced to synthesize novel hybrid biocompatible gels for a wide range of applications [10,11]. It is interesting to note that sol–gel process is in-expensive and the gels so produced are non-toxic [12,13], involving simultaneous hydrolysis and condensation of alkoxide or salt. The sol–gel processes have been utilized to produce bioactive coatings, powders, and substrates that are suitable to be used as implants and sensors [14,15]. Advantage of employing sol–gel method to produce hybrid materials is that it combines both inorganic and organic properties. In addition, this method offers the possibility of obtaining homogeneous hybrid materials under low temperature, which further provides the scope of incorporating a variety of compounds [16,17]. Hydrogels have shown promising applications in the field of pharmacy and medicine, apart from their popular use in the non-medical sector [18]. This can partially be attributed to the absorption capacity ⁎ Corresponding author. E-mail addresses: [email protected], [email protected] (B. Madhan).
http://dx.doi.org/10.1016/j.msec.2016.04.067 0928-4931/© 2016 Elsevier B.V. All rights reserved.
and retention of mechanical strength of these hydrogels. Hydrogels are preferred because they have high water content, with soft and rubbery surface, mimicking human tissue viz., muscles, tendons and cartilage [19,20]. These characteristic features make them potential candidate for transdermal drug delivery systems. The hydrogels liquefy necrotic tissue on wound surface and prevents loss of body fluids. Moreover they are non-adherent and the dressing can be removed without trauma to the wound bed [21]. The feasibility of using natural materials including polysaccharides and proteins for the fabrication of hydrogels has been well-demonstrated. Among these materials, collagen, chitosan, keratin, gelatin, elastin, fibrin hyaluronic acid etc. exhibit prominent bioactivity in biomedical applications [22–24]. Keratins are naturally derived proteins that can be fabricated into several biomaterial forms including hydrogels [25]. Although keratins are considered to be intracellular cytoskeleton proteins belonging to the family of intermediate filaments [26], they contain the cell adhesion motifs LDV (leu-asp-val) and RGD (arg-gly-asp) [27], enabling cell attachment and inturn supporting cell growth and development [28]. Keratins in the form of coatings, fibres and films have also shown promising cytocompatibility (Reichl 2009; Rouse and Van Dyke 2010). Keratin based hydrogels have been reported to be neuro-inductive and capable of facilitating regeneration in a peripheral nerve injury [29]. Apart from this, keratin hydrogels are known to enhance wound healing by the activation of keratinocytes in the wound bed. These hydrogels also acts as a protective barrier, absorb excess wound exudates, and maintain a moist environment, which in turn helps in pain reduction [30]. In particular, silica-based hydrogels are well recognized as suitable matrices for biomaterial development. Researchers have extensively worked on sol–gel processed silica and found their potential as drug
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Fig. 1. Schematic representation of texture analysis to calculate i) Hardness (peak height), ii) Cohesiveness (B1 + B2)/(A1 + A2), iii) Adhesiveness (C), iv) Compressive strength (peak stress/peak strain), v) Resilience (A2/A1).
carriers [31–33]. It has been established that sol–gel produced pure silica are biocompatible and even enhance the formation of fibroblast or osteoblast leading to increased collagen production [12,34–36]. Our group has established the use of sol–gel processed collagen – silica composite materials for wound healing applications [37,38]. Varying the amount of silica could tune the mechanical property of the collagen scaffolds obtained [39]. Gorji, Allahgholi Ghasri, Fazaeli and Niksirat [13] used tetraethyl orthosilicate (TEOS) as a precursor for preparation of silica gels. It is a well-known fact that the hydrolysing agent in the TEOS-derived sols is water. Usually acid hydrolysis of TEOS is employed to prepare aggregative stable spinnable sols. Inorganic acids like hydrochloric acid (HCl), nitric acid (HNO3), etc, are frequently used as acid catalysts, but occasionally organic acids like acetic acid is also being used [10,40,41]. There are two different ways through which sol–gel synthesis can be carried out i.e. single-stage (acid) or two-stage (acid/acid) hydrolysis of TEOS [42–44]. In both the cases, first stage is often carried out with scarcity of water. TEOS acts as a crosslinking agent for keratin hydrogel formation. TEOS has a remarkable property to produce silicon dioxide (SiO2) when mixed with water in presence of a catalyst which can be an acid or a base as per Stober reaction which was discovered by Werner Stober. This process takes place via LaMer model or nucleation model comprising of two steps. It begins with nucleation (fast process) followed by a slow process of particle growth. In a typical process, silicon
Fig. 2. Representation of keratin-silica hydrogel withstanding tube inversion test (4:1w/w of silica:keratin).
alkoxide viz., tetraethyl orthosilicate (TEOS) is hydrolyzed by water in presence of an acidic catalyst. These reactive silanol (Si–OH) groups of the hydrolyzed silicon alkoxide go through condensation reactions which depend on the pH and temperature of the solution, to form siloxane (Si–O–Si) bonds, producing a 3D porous gel structure [45]. Earlier we have established that keratin can be effective for proliferation of fibroblast cells [7]. Moreover, different forms of keratin based materials are required as effective materials for soft tissue engineering applications such as wound healing. This paper deals with the development of a novel keratin hydrogel system by combining keratin with silica, where the sol–gel transition of silica had been advantageously used for the establishment of hydrogel. 2. Materials and methods 2.1. Materials Keratin was extracted from bovine hooves, TEOS (C8H20O4Si) was obtained from Alfa Aesar (United States), HCl was obtained from Fisher
Table 1 Gelation time of keratin-silica hydrogel at varying concentration of silica. Ratio of silica:keratin (w/w) 0.4:1 0.8:1 1.2:1 1.6:1 2:1 2.4:1 2.8:1 3.2:1 3.6:1 4:1 4.4:1 4.8:1 5.2:1 5.6:1
Gelation time No gelation No gelation 2h 1 h 30 min 1 h 10 min 1h 40 min 20 min 15 min 10 min 8 min 7 min 6 min 5 min
Fig. 3. Hardness of keratin-silica hydrogels with varying concentrations of silica.
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Fig. 4. Cohesiveness of keratin-silica hydrogels with varying concentrations of silica.
Scientific (United Kingdom). For in vitro analysis, DMEM and MTT were purchased from Sigma-Aldrich (Sigma-Aldrich Chemicals Pvt. Ltd., India), FBS from Gibco® (Invitrogen, USA) and 4′,6-diamidino-2phenylindole (DAPI) from HiMedia (HiMedia Laboratories Pvt. Ltd., India). 2.2. Methods 2.2.1. Fabrication of keratin-silica hydrogel 2.2.1.1. Preparation of keratin solution. Keratin from bovine hooves was extracted as per procedure detailed elsewhere [7]. Briefly, 10 g of defatted raw bovine hooves was subjected to reduction using 7 M Urea, 6 g sodium dodecyl sulphate and 15 ml 2-mercaptoethanol at 60 °C for 12–14 h. The supernatant was then dialyzed against water to obtain reduced keratin. Extracted lyophilized (powdered) hoof keratin was solubilized in PBS (pH 7.4) to make a stock solution with the concentration of 5 mg/ml. 2.2.1.2. Preparation of acidified TEOS. For the fabrication of hydrogels, TEOS has been used as a precursor of silica. It was acidified using 0.1 N HCl to obtain a pH close to 2. It is known that sol–gel transition of silica in TEOS can be achieved by elevating the pH from 2 to 7. For this, keratin stock solution (pH 7.4) was drop-wise added to TEOS to obtain keratinsilica hydrogels by varying silica:keratin ratio (w/w) from 0.4:1 to 5.6:1. 2.2.2. Tube inversion test For the preparation of keratin-silica hydrogels with different concentrations of TEOS, the solution mixture was poured into a glass vial. To check the formation of stable gel, test tube was inverted upside down every 1 min and time for gelation was noted down in each case.
Fig. 5. Adhesiveness of hydrogels with varying concentrations of silica.
Fig. 6. Resilience of hydrogels with varying concentrations of silica.
Tube inversion test is used to determine the effect of varied concentration of silica on gelation time of keratin. 2.2.3. Texture analysis It is necessary to evaluate the mechanical properties of hydrogel in order to determine whether the hydrogels will be able to maintain their physical structure at the site of application. Keratin-silica hydrogels with varying concentrations of TEOS were analyzed for their textural properties using Texture Analyzer Pro CT V1.4 Build 17 (Brookfield Engineering Labs, Inc). The respective samples (with the diameter of 3.3 cm), taken in triplicate under confined conditions were compressed to 50% deformation at room temperature, using a cylindrical probe with the diameter of 12.7 mm. Hardness, adhesiveness, cohesiveness, resilience and compressive strength of keratin-silica hydrogel system were computed using TexturePro CT software as shown in Fig. 1. 2.2.4. FTIR spectroscopy The optimized keratin-silica hydrogel and its constituents (i.e. keratin and TEOS) were analyzed using JASCO FT/IR-4200 A spectrometer. FTIR spectra were recorded in the range of 4000 to 500 cm−1 with the resolution of 4 cm−1 at 32 scans/sample using conventional KBr pellet method. In brief, the lyophilized (powdered) samples were mixed with KBr and formed into pellets before recording the spectrum. 2.2.5. Hydrogel morphology The surface morphology, structural integrity and interconnectivity of the pores in the optimized hydrogel were observed using Ziess SUPRA 55VP field emission scanning electron microscope (FESEM) at a magnification of 5000×. The samples for analysis were prepared by
Fig. 7. Compressive modulus of hydrogels with varying concentrations of silica.
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The final concentration of DAPI used for staining was 1 μg/ml which was adjusted using Dimethyl sulfoxide (DMSO). For this, the cellseeded control (polystyrene plate) and the hydrogel were stained with DAPI solution for 20 min in dark and then washed thrice with PBS to visualize under fluorescence microscope (Olympus CKX41). 3. Results and discussions 3.1. Fabrication of keratin-silica hydrogels
Fig. 8. FTIR spectra of keratin, TEOS and hydrogel.
sputter coating the scaffold surface with a thin layer of gold. The surface of the samples was scanned at 5 kV. 2.2.6. Swelling study The optimized keratin-silica hydrogel was freeze dried to obtain its dry weight (Wd). It was then immersed in 10 ml PBS (pH 7.4) till complete saturation. The weight of the swollen hydrogel (Ws) was measured after wiping its surface and the swelling ratio was calculated by the following equation: Swelling ð%Þ ¼ ½ðWs −Wd Þ=Wd 100
2.2.7. Biocompatibility (MTT) assay NIH 3T3 fibroblast cell lines obtained from National Centre for Cell Science (NCCS), Pune, India, were used for the cell viability test. The cells were cultured and allowed to passage for 3 times. The passaged cells were seeded optimized on keratin-silica hydrogel at a density of 1 × 105 cells/well. The hydrogel was incubated at 37 °C with 5% CO2 for 3 h. To this, DMEM containing 10% FBS, penicillin (50 unit/ml) and streptomycin (50 unit/ml) were added and incubated at 37 °C with 5% CO2. The medium was changed every 2 days and the hydrogels were observed under microscope. After 1, 3 and 5 days of culture, cell viability was measured quantitatively using MTT assay, which is a colorimetric technique for determining the number of viable cells, that relies on the conversion of MTT to MTT formazan by the enzyme mitochondrial reductase of the viable cells [46]. The cell extracts from the hydrogel were mixed with MTT solution and the absorbance at 550 nm was measured on a microplate reader (Biotek, USA). The cell-seeded keratin-silica hydrogel was also stained with DAPI to observe the attachment of NIH 3 T3 fibroblast cells with the hydrogel.
Keratin-silica hydrogels were prepared by adding different concentrations of TEOS ranging from 0.4 to 5.6 (w/w) of keratin. 5 mg/ml keratin stock solution was used for the preparation of keratin-silica hydrogels by varying the amount of keratin. Time taken for the solution to turn into gel was recorded using tube inversion method and termed as gelation time as presented in Table 1. As a representation, one of the keratin-silica hydrogel system is presented in Fig. 2. It is clearly observed that the gelation time of the hydrogels reduced with increasing silica concentration in the hydrogel. 3.2. Texture analysis of keratin-silica hydrogel Texture analysis is a fast and reproducible method, providing a more straight-forward approach to characterize the hydrogels. The parameters like hardness, adhesiveness, cohesiveness, resilience and compressive strength provide insight on the spreadability of the product, ease of product removal [19], ability of a product to regain its original shape etc. Syverud, Pettersen, Draget and Chinga-Carrasco [47] performed compression of hydrogels using texture analyzer to determine the compressive strength/young's modulus. To evaluate the mechanical properties of hydrogels, texture profile analysis (TPA) had been proposed as a suitable and reliable method by Jones, Woolfson and Brown [48]. Later, Hurler, Engesland, Poorahmary Kermany and Škalko-Basnet [49] proposed a simplified method using texture analyzer which excludes time consuming rheological characterization. Here, the keratin-silica hydrogels are represented in terms of increasing amount of silica. 3.2.1. Hardness Determining the hardness of the hydrogel is a measure of applicability of the hydrogel onto the skin. It can be defined as maximum force required to attain a given (here 50%) deformation. This is calculated by the peak of force during the first compression cycle [50]. Lower the gel hardness lower is the force required to recover the gel from the container and hence more ease of applicability. But, with decrease in hardness the retention time of the gel on the site (for example, wound) of application would reduce. Therefore, the hydrogel should possess an optimum hardness. It can be observed from Fig. 3, as the concentration
Fig. 9. Scanning electron micrographs of a) keratin-silica hydrogel with 4.4:1 w/w of silica:keratin and b) pure keratin at 5000×.
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of silica increases (from 1.2 to 5.6, w/w), hardness of the hydrogel increases from 46 ± 3 g to 184 ± 6 g. As evidenced from Fig. 3, the hardness of the hydrogels with low concentration of silica i.e., from 1.2:1 to 3.2:1 (w/w, silica:keratin) was low and it significantly increased from 4:1 (P b 0.001) onwards. Hydrogels with 5.2 and 5.6 ratio of silica:keratin (w/w) showed maximum hardness of 179 g and 184 g respectively. In terms of recovery of the hydrogels, all the prepared hydrogels showed no significant difference. 3.2.2. Cohesiveness Cohesiveness is the measure of how well a hydrogel can withstand a second deformation compared to that of the first one. In other words, it is the tendency of the hydrogel to adhere to itself. Cohesiveness is calculated using the following formula: Cohesiveness ¼
ðB1 þ B2Þ ðA1 þ A2Þ
Where, (B1 + B2) is the area under second compression cycle and (A1 + A2) is the area under first compression cycle, as shown in Fig. 1. Cohesiveness determines the reconstruction ability of the hydrogel following its application [50,51]. Higher the cohesiveness, more is the structural recovery [52,53]. Fig. 4 shows cohesiveness increased from 0.42 ± 0.04 to 1.49 ± 0.052 as the silica concentration in the keratinsilica hydrogel increased from 0.4 to 5.6 mg/ml of keratin. [19] also showed similar cohesiveness results for methacrylate based hydrogels. As evident from Fig. 4, cohesiveness of hydrogels with silica:keratin content of 1.2:1 was not significantly different from that of 5.6:1 (P N 0.05). This shows that all the hydrogels could withstand second deformation. It also means that the behaviour of the hydrogels during second compression cycle was similar to that of the first one. 3.2.3. Adhesiveness Adhesiveness represents the work required to overcome the attractive forces between the surface of the hydrogel and that of the probe. It is measured by calculating the force applied by the hydrogel to the probe when it redraws after compression. Adhesiveness was calculated as the negative force area subsequent to the first compression cycle, represented as (C) in Fig. 1. This parameter also relates to the retention time of the hydrogel onto the site of application for the desired period of time. More the retention period of the hydrogel, shorter is the treatment period resulting into improved patient compliance [50]. As evident from Fig. 5, adhesiveness of the hydrogels significantly increased from 3.2:1 to 5.6:1 (w/w) of silica:keratin. Adhesiveness of hydrogel with 4:1 ratio of silica:keratin was observed to be 0.48 ± 0.035 mJ, while that of hydrogels with 4.4:1, 4.8:1 and 5.2:1 (w/w) of silica:keratin was 0.50 ± 0.04 mJ, 0.55 ± 0.042 mJ and 0.58 ± 0.039 mJ respectively. Hence, adhesiveness of 4:1 (silica:keratin) hydrogel was not significantly different from that of 4.4:1, 4.8:1 and 5.2:1
Fig. 10. Swelling kinetics of optimized keratin-silica hydrogel.
hydrogels (P N 0.05). Whereas, 5.6:1 (silica:keratin) hydrogel showed significantly higher adhesiveness (P b 0.01). This implies that 5.6:1 (silica:keratin) hydrogel could have more retention time on the site of application. In other words, it also means that this hydrogel tends to stick to the container, which can pose difficulty while handling it. This observation is in-line with the previous result (as shown in Fig. 3), showing highest hardness among all the fabricated hydrogels, thereby reducing the ease of spreadability of the hydrogel (silica:keratin 5.6:1) on the wound site. Therefore, hydrogels with 4:1 to 5.2:1 (w/w) of silica:keratin are favourable compared to the rest. 3.2.4. Resilience Resilience is how well a substance fights to regain its original position after being compressed. It can also be considered as instant springiness. It is calculated using the following formula:
Resilience ¼
A2 A1
Where, A2 is the area during the withdrawal of first compression cycle and A2 is the area during the compression of first compression cycle [54]. It can be deduced from Fig. 6, that resilience of the hydrogels increased with increasing concentration of silica in the hydrogels. But, after a certain concentration, it again started decreasing. Resilience of 2.4:1, 2.8:1 and 3.2:1 (w/w) of silica:keratin hydrogels is similar to each other (P N 0.05). Moreover, 3.6:1 hydrogel showed a resilience of 0.68 ± 0.33, while that of 4:1 and 4.4:1 hydrogel was found to be 0.68 ± 0.35 and 0.69 ± 0.37 respectively. Therefore, the resilience of 3.6:1 (silica:keratin) hydrogel was not significantly different from that of 4:1 and 4.4:1 (P N 0.05). However, as the concentration of silica further increased from 4.8:1 to 5.6:1, the resilience dropped. This could be due to increased silica content in the hydrogels which might impart brittleness, thereby reducing the instant springiness. It is interesting to note that the resilience of 5.6:1 hydrogel was similar to that of 4.8:1 hydrogel. Therefore, based on the resilience, 3.6:1 to 4.4:1 hydrogels were found to be favourable for the current study. 3.2.5. Compressive strength Compressive strength or compressive modulus was calculated as the ratio of peak stress to the peak strain during the first compression cycle [55]. Fig. 7 shows that compressive moduli of 1.2:1, 1.6:1, 2:1, 2.4:1 and 2.8:1 silica:keratin (w/w) hydrogels was not significantly different (P N 0.05). But the compressive modulus of 3.2:1 hydrogel was found to be significantly higher than that of 1.2:1 (P b 0.001). Further, 4.4:1 hydrogel possessed appreciably high compressive modulus of 4.53 ± 0.42 kPa which was not significantly different from that of 4.8:1, 5.2:1 and 5.6:1 hydrogels.
Fig. 11. Biocompatibility assay of keratin-silica hydrogel.
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Fig. 12. Fluorescence micrographs of DAPI stained, 3T3 fibroblast cells-seeded a) polystyrene plate (control) and b) optimized keratin-silica hydrogel.
Based on the texture analysis, it can be deduced that keratin-silica hydrogel with 4.4:1 silica:keratin (w/w) could be the optimum hydrogel showing easy spreadability, good retention on the site of application, considerable springiness and appreciably high mechanical strength. This optimized hydrogel was further characterized using different analytical tools. 3.3. FTIR spectroscopy of keratin-silica hydrogel system The FTIR spectra of the keratin (K), TEOS and optimized keratinsilica hydrogel (4.4:1) are shown in Fig. 8. In keratin composite scaffold, the peaks observed near 3280 cm− 1 are characteristic of amide A, whereas amide I falls in the range of 1600–1700 cm−1, amide II near 1520 cm− 1and amide III in the range of 1220–1300 cm− 1 [56]. The peak of amide A is due to stretching vibration of N–H bonds. The peak of amide I falling at 1654 cm−1 is caused by stretching vibration of C– O bonds. N–H bending and C–H stretching vibration which is responsible for amide II band is observed at 1542 cm−1. At 1242 cm−1 the sharp peak of amide III is observed that is due to the phase combination of C–N stretching and N–H in plane bending and partial addition from C = C bending and C–C stretching vibration. In case of TEOS, asymmetric vibrations of Si–O were observed at 1085 cm−1. This peak was also clearly seen in keratin-silica hydrogel, but with lower intensity. Peaks at 945 cm−1 and 796 cm−1 indicated asymmetric vibrations of Si–OH bond and symmetric vibration of Si–O bond respectively. The hydrogel was observed to possess a shoulder band around 945 cm−1 and a low intensity band at 796 cm−1. Therefore, it could be deduced that keratin-silica hydrogel showed the characteristic properties of its constituents i.e. keratin and TEOS. In addition to this, CH3 peak at 2980 cm−1 and CH2 peak at 2930 cm−1 which are attribute to the unreacted TEOS in silica, were absent in the hydrogel. This affirms that the reaction of TEOS forming silica was complete. There was no significant appearance/disappearance of any peak in the spectrum of hydrogel as the hydrogel is most likely to be formed by the physical interactions. 3.4. Morphology of optimized keratin-silica hydrogel SEM analysis revealed that the optimized keratin-silica hydrogel possessed porous micro-architecture as shown in Fig. 9. It was seen to be a homogenous matrix with uniform distribution of its constituents. At 5000×, the prominent needle-like morphology was portrayed by silica produced by TEOS. Silica needles were completely embedded into the keratin hydrogel. 3.5. Swelling behaviour of optimized keratin-silica hydrogel Dried optimized hydrogel was immersed in PBS (pH 7.4) to understand the swelling behaviour. The samples (in triplicate) were immersed in PBS till complete saturation. When placed in water/solvent,
a polymeric hydrogel swells without dissolving in it. The solvent continues to penetrate the polymeric network of the hydrogel till the equilibrium between osmotic and elasticity forces is achieved [57]. Parameters influencing the swelling of hydrogels are network density, solvent nature and polymer solvent interaction [58]. The optimized keratin-silica hydrogel was analyzed (in triplicate) for its swelling ratio. Fig. 10 shows the kinetics of hydrogel swelling which was calculated as:
Swelling ratio ¼
Wt−W0 W0
Where, Wt refers to the weight of the hydrated hydrogel at time t, while W0 refers to weight of the dried hydrogel. The averaged swelling of the keratin-silica hydrogel was found to be 540 ± 28.97%, which is appreciably good for biomedical applications such as wound healing. The swelling kinetics of the keratin-silica hydrogel (dried form) indicated that it got saturated around 4 h.
3.6. Biocompatibility assay The NIH 3T3 cells cultured on the optimized keratin-silica hydrogels which were monitored continuously for 5 days. The untreated culture plate wells (polystyrene) comprising of cells were treated as control. The cell viability results of MTT assay (Fig. 11) performed after 1, 3 and 5 days showed no significant difference (P N 0.05) between the test sample and control, and after 7 days the cell viability of the composite scaffold was similar to that of collagen scaffold. The optimized keratin-silica hydrogel was stained with DAPI, which is a fluorescent dye and can penetrate the intact cell membrane. Fig. 12 shows the images of DAPI stained, cell-seeded optimized hydrogel and the polystyrene plate (control). It can be observed that the hydrogel (Fig. 12b) showed appreciable proliferation and infiltration of the fibroblast cells into the hydrogel matrix. There was no significant difference between the control (Fig. 12a) and the optimized hydrogel. These results corroborate with the MTT assay.
4. Conclusions Keratin from bovine hooves could be successfully fabricated into hydrogel by using TEOS as a precursor of silica. The optimized keratinsilica hydrogel showed considerable hardness, while maintaining ease of applicability. Moreover, it displayed instant springiness, high compressive modulus and biocompatibility. Together, these characteristics show that the optimized hydrogel could be used as a biomaterial for biomedical applications, especially for wound healing.
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