NHS for corneal tissue engineering applications

International Journal of Biological Macromolecules 126 (2019) 620–632

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International Journal of Biological Macromolecules journal homepage: http://www.elsevier.com/locate/ijbiomac

Preparation and in vitro characterization of cross-linked collagen–gelatin hydrogel using EDC/NHS for corneal tissue engineering applications Hamid Goodarzi a,1, Khosrow Jadidi b,1, Samiramis Pourmotabed c, Esmaeel Sharifi d,e,f,⁎, Hossein Aghamollaei g,⁎⁎ a

Department of Biomedical Engineering, Faculty of Chemical Engineering, Tarbiat Modares University, Tehran, Iran Department of Ophthalmology, Baqiyatallah University of Medical Sciences, Tehran, Iran Department of Emergency Medicine, School of Medicine, Hamadan University of Medical Sciences, Hamadan, Iran d Department of Molecular Medicine and Genetics, School of Medicine, Hamadan University of Medical Sciences, Hamadan, Iran e Research Center for Molecular Medicine, Hamadan University of Medical Sciences, Hamadan, Iran f Department of Tissue Engineering and Biomaterials, School of Advanced Medical Sciences and Technologies, Hamadan University of Medical Sciences, Hamadan, Iran. g Chemical Injuries Research Center, Systems biology and Poisonings Institute, Baqiyatallah University of Medical Sciences, Tehran, Iran b c

a r t i c l e

i n f o

Article history: Received 14 May 2018 Received in revised form 14 December 2018 Accepted 14 December 2018 Available online 15 December 2018 Keywords: Gelatin Collagen EDC/NHS Hydrogel Corneal tissue engineering

a b s t r a c t Corneal disease is considered as the second leading cause of vision loss and keratoplasty is known as an effective treatment for it. However, the tissue engineered corneal substitutes are promising tools in experimental in vivo repair of cornea. Selecting appropriate cell sources and scaffolds are two important concerns in corneal tissue engineering. The object of this study was to investigate biocompatibility and physical properties of the bioengineered cornea, fabricated from type-I collagen (COL) and gelatin (Gel). Two gelatin based hydrogels crosslinked with EDC/NHS were fabricated, and their physicochemical properties such as equilibrium water content, enzymatic degradation, mechanical properties, rheological, contact angle and optical properties as well as their ability to support human bone-marrow mesenchymal stem cells (hBM-MSCs) survival were characterized. The equilibrium water content and enzymatic degradation of these hydrogels can be easily controlled by adding COL. Our findings suggest that incorporation of COL-I increases optical properties, hydrophilicity, stiffness and Young's modulus. The viability of hBM-MSCs cultured in Gel and Gel: COL was assessed via CCK-8 assay. Also, the morphology of the hBM-MSCs on the top of Gel and Gel: COL hydrogels were characterized by phasecontrast microscopy. This biocompatible hydrogel may promise to be used as artificial corneal substitutes. © 2018 Elsevier B.V. All rights reserved.

1. Introduction Corneal tissue with a thickness of about 500 μm, contains three different cell types, three layers and has unique properties such as transparency, elasticity and clarity, with an important role in vision. Also, cornea protects the inner parts of the eye [1,2]. Keratopathy due to trauma or herpetic infection, decrease corneal transparency and are major causes of vision loss or even blindness worldwide. Each year, about 2 million people suffer from loss of vision and blindness due to corneal diseases, such as keratoconus, corneal ulceration, and trachoma [2]. It is estimated that nearly 10 million people worldwide have vision ⁎ Correspondence to: Department of Tissue Engineering and Biomaterials, School of Advanced Medical Sciences and Technologies, Hamadan University of Medical Sciences, Hamadan, Iran. ⁎⁎ Correspondence to: Chemical Injuries Research Center, Systems biology and Poisonings Institute, Baqiyatallah University of Medical Sciences, Tehran, Iran. E-mail addresses: E.sharifi@umsha.ac.ir, Esmaeel.sharifi@gmail.com (E. Sharifi), [email protected], [email protected] (H. Aghamollaei). 1 Authors contributed equally.

https://doi.org/10.1016/j.ijbiomac.2018.12.125 0141-8130/© 2018 Elsevier B.V. All rights reserved.

loss due to corneal lesion [3]. Corneal tissue engineered construct are promising tools for treatment and prevention of corneal lesions leading to vision loss and ultimately blindness [4,5]. Hydrogels are being extensively used for tissue engineering purposes [6–8]. There have been considerable attempts for preparation of biomimetic hydrogels [9] that mimic extracellular matrix (ECM) and native cell microenvironment [9,10]. It has been shown that hydrogels with natural origin (dextran, hyaluronic acid, chitosan, collagen, …) are an abundant source of biocompatible materials for tissue engineering applications [11–15]. Recent strategies for design and fabrication of the corneal construct comprise a variety of hydrogel sources [3,16–18] and manufacturing techniques, such as silk, chitosan films, gelatin hydrogels, collagen scaffolds, and electrospun collagen mat [16,19]. Collagen is the major component of ECM and used as a natural biomaterial for corneal substitution, with perfect biocompatibility, biodegradability, and conducive to active progenitor cells that promote regeneration [18,20,21]. A collagen triple helix consists of repeating residues of glycine, proline, and hydroxyproline [22]. Type I collagen (COL I) is the main collagen of body tissues and mostly found in ECM of skin, tendon, bone, ligament

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and cornea. Type-I collagen, the primary molecule of the native cornea, is a suitable biomaterial for corneal tissue engineering applications [16,19,23]. Griffith et al. [24], incorporated a collagen-chondroitin sulfate based material to develop functional human corneal equivalents. This collagen-based human corneal substitute was comparable to the human cornea regarding the stromal swelling, physiological endothelium activity, and response to chemicals. In another study, Shimmura et al. [25], blended collagen and a synthetic acrylamide-based polymer [poly (N-isopropyl acrylamide) (pNIPAAm)] to develop collagenbased composites. Pure collagen-based hydrogels are not strong enough and degrade rapidly. One of the main reasons for the weak mechanical properties of collagen hydrogels is their high water content. To increase the mechanical properties, to avoid rapid degradation during in vivo application and to suppress antigenicity of collagen-based hydrogels, different types of physical (dehydrothermal treatment, DHT) and chemical (EDC/NHS (zero-length), glutaraldehyde (non-zero-length)) crosslinking methods can be used [22,26–31]. Merrett et al. [32], used a mixture of porcine collagen and aqueous solutions of EDC/NHS to produce a cross-linked collagen tissue substitute for corneal implantation in contact lens molds. The cross-linked collagen scaffolds promote corneal cell regeneration. Mechanical properties and biodegradability of collagen hydrogels can also be improved through using natural biopolymer or synthetic polymers within the collagen network. Rafat et al. developed a hybrid polymer network scaffolds using collagen and chitosan cross-linked with EDC/NHS along with poly (ethylene glycol) dibutyraldehyde (PEG DBA). The resultant scaffolds were transparent, optimally strong and biocompatible [14]. Gelatin is formed by partial digestion/heat denaturation of the triple helical structure of collagen and

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contains identical amino acids and cell binding moieties similar to collagen [33]. Gelatin hydrogel is a good candidate for corneal tissue engineering due to its desired transparency, biocompatibility, low antigenicity, and favorable cell attachment [16,31,34,35]. Gelatin has been considered as potential scaffolds for corneal epithelium, corneal endothelium and retinal pigment epithelium as a bio-artificial corneal stroma. There are many crosslinking options to improve gelatin scaffolds' strength. Researchers have suggested that zero-length (EDC/ NHS) cross linkers may be preferable regarding both cell compatibility and biocompatibility [31]. Thus, for tissue engineering applications, adequate intermolecular crosslinks of hydrogels with appropriate biocompatible molecules is essential for developing stable materials with a demanded mechanical integrity [1,2,14,22,30]. Mesenchymal stem cells are multipotent cells that are found in many body tissues such as bone marrow, adipose tissue, heart, dental pulp, and limbal stroma of the eye. Recently, Mesenchymal stem cells are considered as a new source for corneal reconstruction [36–40]. In a study done by Liu et al. [36], mesenchymal stem directly transplanted to mouse cornea resulted in the formation of keratinocyte like cells. In another study, human umbilical cord-derived mesenchymal stem cells were used to treat thin and cloudy corneas of lumican null mouse [37]. Also, MSCs can be harvested easily and differentiated into epithelial cells. Ma et al. [41], isolated mesenchymal stem cells from human bone marrow and cultured on amnion membrane. After 7 days of injury, an amnion membrane containing cells was implanted into the rat cornea. As a result, they showed that cornea was regenerated using bone marrow-derived mesenchymal stem cells. Herein, we developed a Gel: COL hydrogel, cross-linked with EDC/NHS (Gel: COL-EDC/

Fig. 1. a) Schematic picture of the hydrogel preparation process for Gel and Gel: COL hydrogels. b) Crosslinking reaction mechanisms between collagen and gelatin by EDC/NHS.

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Fig. 2. Physio-chemical characterization of collagen and gelatin by SEM, FTIR and XRD.

NHS) for corneal tissue engineering applications. The effects of collagen on the physic-chemical and biological properties of gelatin hydrogel (structure, equilibrium water contact, enzymatic degradation, contact angle, optical properties, rheological and mechanical properties) were investigated. In this study, human bone marrow mesenchymal stem cells (hBM-MSCs) were used to assess biological properties. Incorporating COL into the Gel hydrogel promoted hBM-MSCs attachment and survival. As a result, it is assumed that contribution of collagen in gelatin hydrogel enhances its physic-chemical and biological properties and

this blend hydrogel may be favorable for being developed as corneal tissue engineered substitutes. 2. Materials and methods 2.1. Materials N-(3-Dimethylaminopropyl)-N′-ethylcarbodiamide hydrochloride (EDC; Mw 191.7; Sigma-Aldrich) and N-hydroxysuccinimide

Fig. 3. The surface morphology (SEM micrographs) and FTIR spectra patterns of freeze dried Gel and Gel: COL hydrogels.

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Table 1 The equilibrium water content, second order elastic constant, complex modulus and loss tangent of Gel and Gel: COL hydrogels. The water content between Gel and Gel: COL showed a significant difference (*p b 0.05), but there is no significant difference between blend sample with native cornea (p N 0.05). Also, There were no statistically significant differences between the complex modulus of hydrogels. Hydrogels

Equilibrium water content

Second-order elastic constant (MPa)

Complex modulus G⁎ (KPa)

tan δ

~78 [11] 55 ± 2.13 67 ± 3.13

3–13 [43] 3.92 × 10−3 ± 0.85 6.28 × 10−3 ± 1.17

– 1.32 2.23

– 0.061 0.049

Human cornea Gel Gel: COL

(NHS; MW 115.09; Sigma-Aldrich) solution were used to cross-link the hydrogels. The gelatin (type A, MW: 20000–30,000 Da) from porcine skin was purchased from Sigma-Aldrich (USA). The type I collagen was obtained from goat tendon and solved in hydrochloric acid (HCl) (6.33 mg/ml, pH 3.2). Dulbecco's modified Eagle's medium (DMEM), trypsin-EDTA, antibiotics (penicillin/streptomycin), fetal bovine serum (FBS) and phosphate buffered saline (PBS) were purchased from Gibco. Deionized water was used throughout. 2.2. Preparation of hydrogels Two types of hydrogel were fabricated for this research: gelatin (Gel) and gelatin-collagen hydrogel (Gel/COL) with the ratio of 10:1. For this purpose, gelatin was prepared in 0.5 mol/l hydrochloric acid solution at a concentration of 5% w/v. Collagen solution (6.33 mg/ml) was blended in 0.1 mol/l solution and added drop by drop at 510 μm/min speed to Gel solution using a syringe pump (FNM company, SP2000HSM model, Iran) in a stirrer at 4 °C. After that, two solutions were mixed for 2 h with a mass ratio of Gel: COL = 10:1. To reduce the biodegradation rate and enhance the mechanical properties, the prepared hydrogels were cross-linked as following: EDC and NHS were added to the Gel: COL and completely mixed at 4 °C to form a solution with a mass ratio of EDC: NHS: (Gel/COL) = 1:1:12. Crosslinking was conducted by stirring the solution for 4 h. The mixed solution was poured into a mold (5 cm in length and 1 cm in diameter), after gelation at 4 °C for 24 h. Two-step freezing was performed at −20 °C for 7 h followed by −80 °C for 24 h. To produce the desired 3D porous organization through sublimation, frozen hydrogels were moved to the freeze dryer (ALPHA1-2LD) at −57 °C and 0.05 mbar for 24 h. After freeze-drying, freeze-dried hydrogels were rinsed three times with deionized distilled water to remove the residue of EDC and NHS (Fig. 1a) [16].

Eq: ð1Þ : porosity ð%Þ ¼

ðw2 −w3 −ws Þ=ρe  100 ðw1 −w3 Þ=ρe

where W1 is the weight of bottle filled with ethanol; W2 is the weight of bottle filled with ethanol after immersion of the freezedried hydrogel; W3 is the weight of bottle filled with ethanol after removal of the freeze-dried hydrogel; Ws is the weight of the dry freeze-dried hydrogel and ρe is the density of the ethanol [42]. Hydrogel pore size was measured using Image J software on SEM images taken from several cross-sections of the hydrogels [6]. 2.4. Equilibrium water content and in-vitro biodegradation Equilibrium water content of the hydrogels was characterized by soaking them in PBS (pH = 7.4) at the physiological temperature (37 °C) of the cornea for 72 h. After being soaked in PBS to saturate, the hydrogels were blotted quickly with a filter paper to remove the absorbed water [11]. The Equilibrium water content of hydrogels (Wt) was calculated according to the following equation: Wt ð%Þ ¼ ðWet weight–Dry weightÞ=ðWet weightÞ  100% The in vitro degradation behavior of the hydrogel is a very important factor in the development of hydrogels for cornea regeneration [43]. The biodegradation of the hydrated hydrogels (Gel and Gel: COL) was determined in phosphate buffer saline (PBS) at the pH 7.4 and 37 °C. Two hydrated hydrogels were placed in vials containing 5 ml of 10 mM PBS, followed by adding 60 μl collagenase (1 and 10 U/ml) [11]. At different time intervals, each hydrogel was weighed after all surface water was carefully blotted off. The degradation percent of hydrogels was determined according to the following equation:

2.3. Physic-chemical characterization of prepared hydrogels

Degradation% ¼ Wt=W0

Phase analysis of COL, Gel and COL: Gel samples, was carried out by X-ray diffraction (XRD), Philips X'Pert MPD system (Philips, The Netherlands), using Co Kα radiation without any filter. Each run was investigated with 2θ values of 10–90 at a step size of 0.02 and a count time of 0.5 s per step. Fourier transformed infrared analysis was done for chemical characterization of fabricated hydrogels. FTIR analysis of Gel, COL and Gel: COL was carried out by Perkin- Elmer FTIR spectrometer. For this aim, 3 mg of the sample were mixed with KBr (infrared grade) in a ratio 1/100 and pressed into a pellet under vacuum. Then the pellets were investigated at 400 to 4000 cm2. The morphology and microstructure of the freeze dried COL, Gel and Gel: COL hydrogel was assessed using scanning electron microscope (SEM; XL30, Philips) operated at the acceleration voltage of 15 kV. For this reason, freeze-dried specimens of prepared hydrogels were coated with a thin layer of Gold (Au) via sputtering (EMITECH K450X, England) and investigated by scanning electron microscope. The porosity of the freeze-dried hydrogels was measured using Archimedes' principle. Briefly, freezedried hydrogel was completely immersed into a bottle filled with ethanol. The freeze-dried hydrogel with ethanol soaked into the pores was removed from the bottle. The porosity was then calculated according to the following formula:

where W0 is the initial weight of the hydrogel; Wt is the weight of the hydrogel at each time point.

Fig. 4. Enzymatic degradation profiles of Gel and Gel: COL hydrogel in collagenase enzyme (1 and 10 U/ml). The mean (n = 3) and standard deviation bars are shown.

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Fig. 5. a) Stress as a function of strain and (b) stiffness as a function of strain curves of Gel and Gel: COL hydrogels.

2.5. Mechanical characterization Mechanical properties of the hydrogels were measured using a uniaxial testing machine at 37 °C with 200 N load cell according to ASTM F2150-07 standard guide. This test was operated at a loading rate of 5 mm/min until the sample was compressed using unconfined compression to at least 80% of the original height [44]. All of the hydrogels were cut into a cylindrical shape with dimensions of 20 mm height × 10 mm diameter (Fig.1d,c c) with three replications for each group of the hydrogels. The hydrogels were soaked in a PBS for 2 h before the test. To analyze the stiffness of P hydrogels this equation was incorporated: dσ ≡ τðεÞ ¼ kn¼1 kck εk−1 dε

In the equation: stiffness = τ. stress = σ. Young's modulus = c1. strain = ε [45,46]. 2.6. Rheological characterization A MCR 301 rheometer equipped (Anton-Paar, Graz, Austria) with cone-plate geometry (20 mm diameter) and a gap of 0.8 mm was incorporated to investigate rheological properties. Hydrogel samples (1 ml,

Fig. 6. Rheological properties of Gel-based hydrogels. (a, b) Amplitude dependence and (c, d) frequency dependence of modulus of Gel and Gel: COL hydrogels, respectively.

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height 5 mm × diameter 20 mm) were prepared. Oscillatory mode in 37 °C was incorporated for rheological tests. We used amplitude and oscillatory sweep data from the linear viscoelastic area; (γ = 0.01–100, ω = 1 Hz) and (ω = 0.1–100 Hz, γ = 0.1), respectively. Storage modulus (G0 ¼ σγ 0 cosðδÞ, evaluating elasticity) and loss modulus (G″ ¼ σγ 0 sin 0

0

ðδÞ , evaluating viscosity were calculated from the linear viscoelastic region (LVR) (σ 0 is the stress, γ 0 is the strain amplitude, and δ is the phase angle between stress and strain). Complex qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 2 2 modulus (G ¼ ðG0 Þ þ ðG″ Þ ) and total viscoelastic behavior of

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2.8. Optical measurements- light transmittance The Gel and Gel: COL hydrogels were immersed in PBS for N2 h to uptake water. Then, the hydrogel samples were fixed directly into the 1.5 ml specimen chamber of the UV–visible spectrophotometer (Varian carry 50, McKinley Scientific Australia). The optical clarity and transparency of the samples were photographed. The light transmittance of the hydrogel was determined in the range from 400 nm to 800 nm [16]. 2.9. Cell evaluations

″

hydrogel samples and the loss tangent ( tanδ ¼ GG0 ; indicating the total viscoelasticity of material) were measured (tanδ b 0.1 = a strong gel or tanδ N 0.1 = weak gel) [45–48]. 2.7. Surface contact angle characterization Contact angles were measured using sessile drop method by digital microscope dino-light (Taiwan) with an injection volume of 1 μL distilled water as medium and analyzed using Dino Capture 2 software.

2.9.1. Cellular attachment on the freeze-dried hydrogels hBM-MSCs were obtained from Iran National Cell Bank and cultured in Dulbecco's modified Eagle's medium (DMEM, Gibco) supplemented with FBS (10% v/v) and streptomycin (1% v/v). Before cell seeding, the freeze-dried hydrogels were sterilized by immersion in ethanol 70%, for 15 min, then washed with PBS and treated with ultraviolet radiation for 20 min. Briefly, 1 × 104 cells were cultured on each group of freezedried hydrogels in 96-well culture plate and incubated at 37 °C, 5% CO2 with DMEM supplemented with 10% FBS. Cellular attachment on the

Fig. 7. Surface contact angle measurement of the a) Gel and b) Gel: COL hydrogels after 60 s.

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Fig. 8. The light transmittance of the a) Gel and b) Gel: COL hydrogels immersed in PBS solution and representative image of Gel and Gel: COL immersed in PBS solution.

freeze-dried hydrogels was evaluated by SEM. After 48 h, the attached cells were fixed in 2.5% (v/v) glutaraldehyde for 1 h. The freeze dried hydrogels were washed three times with PBS, dehydrated with ethanol solutions (50, 60, 70, 80, 90, and 100%), coated with gold, and investigated at an accelerating voltage of 20 kV by SEM (XL30; Philips). 2.9.2. H&E staining on the freeze-dried hydrogels 3 × 10 4 cells were seeded on top each group of freeze-dried hydrogels in 24-well culture plate (n = 3) and incubated in DMEM

supplemented with 10% fetal bovine serum (FBS) and penicillin/ streptomycin (1% v/v) in a humidified atmosphere of 5% CO 2 . 14 days post cell seeding, specimens were processed with a serial ethanol concentration from 50%, 75%, 90%, and 95% to 100% and were fixed in 10% formalin solution for 24 h and embedded in paraffin. Cross-sections of 5-μm thickness (Thermo Scientific HM315 Microtome) were obtained of the samples. Prepared samples were stained with hematoxyline-eosin (H&E) [49] for 5 min to confirm the presence of cultured cells on scaffold 14 days post cell seeding.

Fig. 9. The morphology of hBM-MSCs in the freeze dried Gel and Gel: COL hydrogels were characterized using SEM after 48 h.

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The integration of the hydrogel and cell distribution was reflected in the captured images. 2.9.3. Cellular attachment on the hydrogels by phase contrast microscopy hBM-MSCs were plated in a 24-well plate at a density of 1 × 104 cells/ml and cultured in 2D (Tissue Culture Plate: TCP) and 3D (Gel and Gel: COL hydrogels) the cells were cultured on top of the gels after for 3 and 7 days, respectively. The morphology of the 2D- and 3D-cultured hBMMSCs were observed under a phase contrast microscopy. 2.9.4. Proliferation assay on the hydrogels by cell counting Kit-8 Cell Counting Kit-8 (CCK-8) is a more sensitive assay than any other tetrazolium salts such as MTT, XTT or MTS which allows sensitive colorimetric assays for the indication of the number of viable cells in the cytotoxicity and proliferation assays. hBM-MSCs were seeded on plate (2D) and Gel and Gel: COL hydrogels (3D) in 24-well plates at 1 × 104 cells/well. Cell proliferation was investigated after 1, 3 and 7 days using Cell Counting Kit-8, according to the manufacturer's instructions. In Gel and Gel: COL hydrogels, the cells were washed twice with PBS and incubated by 1 ml serum free media with 100 μL CCK-8 (0.5 mg/ml, Sigma-Aldrich, USA) for 3 h, and then the supernatant was transferred to 96-well plate; the absorbance of the solution at 450 nm was determined. In empty 24-well plates, the cells were washed twice with PBS and incubated by 100 μL serum-free media with 10 μL CCK-8 for 3 h, and then, the absorbance of the solution at 450 nm was determined [50]. 2.10. Statistical analysis

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microscopy (SEM, XL30, Philips) as shown in Fig. 2a. Collagen sponge had a porous and interconnected structures. This consistent structure is useful for cell proliferation and migration purposes in tissue engineering. The analyses showed no remarkable differences in the microstructure of the Gel and Gel: COL hydrogels. All of the hydrogels have the porous and interconnected structure (the interconnected structure is convenient criteria for cell infiltration, proliferation, and migration into tissue engineered scaffolds) [52]. To measure the average pore sizes of the freeze-dried hydrogels, we used Image J software and SEM images [53]. Gel hydrogel presented pores with average diameters of 20–30 μm, while Gel: COL presented average diameters of 10–30 μm. An ideal tissue-engineered scaffold has N90% porosity to allow the cells to penetrate the depths of the scaffolds and to let diffusion of essential nutrients and removal of waste materials [2]. The porosity of the Gel hydrogel calculated from Eq. (1) was 98.1 ± 1.71, while, that of the Gel: COL was 85.1 ± 1.89. This indicates that the added collagen slightly decreased the porosity of the gelatin hydrogel. FT-IR spectra of hydrogels illustrate some typical spectral bands related to collagen and gelatin. FT-IR spectra (Fig 2.b) for COL scaffold showed C_O stretch coupled with COO\\ at 1635 cm−1 for amide I, N\\H bending coupled with CN stretching at 1546 cm−1 for amide II, N\\H bending coupled with CN stretching at 1237 cm−1 for amide III band, N\\H stretch coupled with hydrogen bond at 3315 cm−1 for amide A, and C\\H asymmetrical stretch at 2964 for amide B. The FTIR spectra drawn from the Gel and Gel: COL hydrogel samples showed the typical bands of the gelatin backbone, that is the Amide A, Amide I, Amide II and Amide III (Fig. 3a) [53]. Amide I band could be a promising indicator of the secondary structure of collagen and gelatin molecule [54]. By collagen addition, frequencies at 1635 cm−1 for Gel: COL sample

Each test was repeated at least three times independently. All data have been reported as mean ± standard deviations (SDs) analyzed by one-way ANOVA with Tukey's multiple evaluates using SPSS 20.0 software. P-values of b0.05 and 0.01 were considered as statistically significant. 3. Results and discussion 3.1. Physic-chemical characterization As shown in Fig. 1(a), collagen solution was first blended into the gelatin. Then, the homogeneous gelatin-collagen was stabilized by an amide type short-range EDC/NHS cross-linking agent (Gel: COL-EDC/ NHS). EDC/NHS offers short-range intermolecular cross-links that are essential to the improve scaffold's stability [14]. A simple cross-linked human collagen type I for corneal tissue application was synthesized by mixing collagen and aqueous solutions of EDC/NHS using Merrett et al. [32]. They showed that EDC/NHS improved the tensile strength and elastic modulus of the corneal substitute to what the native cornea shows. Another study has shown that incorporation of carbodiimide (EDC) into gelatin reinforced alginate hydrogels improves the stiffness [51]. A collagen-phospholipid corneal substitute was synthesized through combinations EDC/NHS cross-linked collagen and poly (ethylene glycol) diacrylate cross-linked 2-methacryloyloxyethyl phosphoryl choline (MPC). The combined hydrogel was superior to the hydrogel made of individual components, mechanically and stability wise [43]. As shown in Fig. 1(b), the EDC/NHS system facilitates the formation of short-range amide bonds by linking a carboxylic group of collagen with an amine group of gelatin. According to Rafat et al. [14], a network is constructed if the formation of bonds including “amine” or “amide” cross-links in the hydrogel is confirmed. The formation of amide bonds in our Gel: COL hydrogels is verified through infrared spectroscopy. In our study, we hypothesized that incorporating collagen type I into the gelatin network can enhance the physic-chemical and biological characterization. The surface morphology of the samples that were dried by the freeze-drying process and were detected by a scanning electron

Fig. 10. H&E staining of freeze dried Gel and Gel: COL hydrogels were characterized after 14 days. Scale bar = 50 μm.

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at 1630–1640 cm−1 region were remained and stabled. It means the secondary structure of the blend sample was unchanged. The XRD patterns of freeze-dried COL and Gel: COL are shown in Fig. 2(c, d). The integrity of the triple-helix conformation of collagen can be identified by XRD through its characteristic peaks. The peaks with 2θ = 20.5° and 2θ = 21° proved the distance triple helixes of the collagen molecule. In Gel: COL sample the secondary structure was unchanged due to the presence of 2θ = 21° in the sample [55,56]. 3.2. Equilibrium water content and in-vitro biodegradation Hydrogels can hold large quantities of water in their structure, which means that the degree of hydrogels is increasing, as it is related

to some mechanical properties and product solubility [57]. The stroma is composed of approximately 78% water, 15% collagen and 7% noncollagenous proteins, proteoglycans and salts [58]. Equilibrium water content for Gel and Gel: COL was about 55.8% and 67.2%, respectively (Table 1) and the water content between Gel and Gel: COL showed a significant difference (p b 0.05) but there is no significant difference between blend samples with native cornea (p N 0.05). The swelling of the blend sample in the PBS solution was improved through increasing collagen content. The equilibrium water content of the Gel: COL hydrogel is comparable to the human cornea (78.0 ± 3.0%) [16]. Generally, the swelling ratio is measured by evaluating the water uptake of dried hydrogels. In this case, since the hydrogels are taken from the cellcultured medium, it is appropriate to study the swelling hydrogels of

Fig. 11. Morphology of the hBM-MSCs on the top of Gel and Gel: COL hydrogels were characterized using phase contrast microscopy after 3 and 7 days. Scale bar = 100 μm.

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freshly made wet hydrogels. Also, various drying techniques can affect the hydrogel structure and negatively influence the swelling results [45]. Therefore, the high degree of the water content of Gel: COL hydrogel can be due to the electrostatic repulsive force between negativelycharged carboxylic acid of collagen and amine groups of gelatin at the physiological pH of 7.4. Also, by the addition of collagen into the gelatin solution the number of hydrophilic group of the network will be increased, and as a result, the blend sample has an appropriate ability to absorb the amount of water, so the swelling ratio of the blend sample increased. As reported in previous studies, swelling ratio of the gelatin changes with the addition of other materials. Zhao et al. [20], prepared the cross-linked collagen (COL)-citric acid (CA) films by making EDC and NHS and proved the appropriate optical performance and water content of the COL-CA films. The rise in citric acid content led to the decrease in water content. They also showed that omitting citric acid results in a 95% water content. We also showed that the collagen combination with gelatin improves the swelling ratio. The water content of the hydrogels depends on the density of polymer molecules and their crosslink density. In future studies, this ratio should be improved by changing the density of the crosslinker and the density of the polymer molecules until it reaches the normal equilibrium of the cornea. Enzymatic degradation can exactly mimic the in vivo and clinical situations [57]. Collagenase enzyme was applied to evaluate the enzymatic degradation of the hydrogels (1 and 10 U/ml in PBS). Also, In vitro degradation of the hydrogels was characterized by showing the residual mass percentage of the scaffold as a function of time (Fig. 4). Generally, with increasing collagenase concentration, the biodegradation rate of the hydrogels will be increased. Gel hydrogel showed higher biodegradation rate as compared to the blend sample. Also, Gel hydrogel completely dissolved within 48 h in 10 U/ml collagenase. However, 49% of the Gel at 7 days in 1 U/ml collagenase solution was dissolved. Additionally, Gel: COL hydrogel was entirely dissolved in 10 U/ml at 7 days. Only 15% of the Gel: COL was dissolved in the same concentration of collagenase (1 U/ml) at 7 days compared to the Gel hydrogel (49%). As shown in Fig. 4, the degradation of EDC/NHS cross-linked Gel: COL hydrogel was completed in collagenase (10 U/ml) at 7 days, indicating that once implanted in vivo, this hydrogel could be degraded or remodeled by the host tissue. The data suggested that the incorporation COL influenced the biodegradation behavior of the hydrogel. The enzymatic biodegradation data showed that the degrading time of Gel: COL hydrogel is slower than that of the Gel hydrogels at the same concentration. Biodegradation rate was generally connected with the crosslinkers and here it can be concluded that kinetics of degradation can be improved by adding collagen molecules and crosslinking agent

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(EDC/NHS). The ability to adjust the degradation rates of Gel: COL hydrogel will likely be useful for the development of tissue engineering scaffolds with controlled degradation during tissue regeneration. The Gel: COL hydrogels that degrade much slower could also be effectively incorporated in Degradation-mediated sustained and slow release of drugs from the matrix over long periods of time [53]. 3.3. Mechanical characterization The cornea has special mechanical properties and constituents that are substantial for linking morphology to mechanical behavior. The cornea is a complex anisotropic composite with nonlinear elastic and viscoelastic properties. The elastic modulus (Young's modulus, E) represents the inherent hardness and stiffness (low compliance) of the material [58]. The stress-strain curve for tissues and hydrogels in the elastic region is non-linear even at small strains. Therefore, a polynomial fit was instead of a linear fit. Also, the stiffness was illustrated as a function of strain to show the behavior of the material in a wider strain range [45]. Fig. 5(a) shows the stress-strain curves for Gel and Gel: COL hydrogels. Like most other hydrogels, The Gel and Gel: COL was initially resistant to deformation under pressure and then as the strain increased, hydrogels became stiffer. As a result, because of the permanent deformation, it ultimately leads to fracture (a phenomenon known as strain hardening) [46]. The Gel and Gel: COL hydrogels fractured at 40% at 49% strain range, respectively. Due to a higher fracture strain, the Gel: COL hydrogel was considered more elastic compared to the Gel hydrogel. The curves of stiffness as a function of increasing strain are shown in Fig. 5 (b). The stiffness was shown to be strain dependent. At low strains 0% to 25% for Gel hydrogel and 0% to 35% for Gel: COL the stiffness was quite constant. The second-order elastic constants (Young's modulus) of the hydrogels are shown in Table 1. The Young's modulus of Gel and Gel: COL hydrogels was 3.92 ± 0.85 and 6.28 ± 1.17 KPa, respectively. There is statistically significant difference (P b 0.05) between Gel and Gel: COL hydrogels, demonstrating that the compressive module was affected by adding collagen. The addition of collagen in hydrogel leads to a greater molecular weight in the cross-linked COL–Gel molecule. A complete understanding of corneal biomechanics is critical for clinical applications; so, it is of utmost importance to know how much difference between the mechanical properties of the scaffolds designed with corneal tissue. According to Elsheikh et al. [59], a study in people aged between 50 and 64, the compressive strength of normal cornea was found to be in the range of 403 to 624 KPa. Generally, Young's modulus of corneal in in vitro study was measured and the result showed that it changes from 0.1 to 57 MPa [61]. In Bektas et al. study it was 1.45 ± 0.03 and 6.53 ± 0.84 KPa for GelMA-10 and GelMA-15 hydrogels, respectively [57]. In Luo et al. in vivo and in vitro study, the gelatin (10%)/ascorbic acid (AA: 0,3,30 and 600 mg) cryogels (fabricated through carbodiimide cross-linking with cryogelation) were used as keratocyte carrier. The Young's modulus of lyophilized hydrogels G/A0, G/A3, G/A30, and G/A600 was 17.5 ± 1.1, 16.3 ± 0.8, 14 ± 0.9 and 1.1 ± 0.4 MPa, respectively [60]. Therefore, according to these studies, diverse materials based on its sources have different mechanical properties so, elastic modulus of our hydrogels was 3.92 ± 0.85 and 6.28 ± 1.17 KPa, respectively. The mechanical properties of the studied hydrogels were different from the native cornea. But calculating Young's modulus is also one of the main requirements for materials research, which should be considered in improving the mechanical properties of hydrogels in in vivo studies. 3.4. Rheological characterization

Fig. 12. The proliferation of Gel and Gel: COL hydrogels and control group (different culture time 1, 3 and 7 days). In Gel: COL hydrogel showed significantly higher than that in Gel hydrogel and control group (after 7 days) by using Cell Counting Kit-8. Results were shown as mean ± SD, n = 3. *p b 0.05, **p b 0.01.

The viscoelastic properties of the Gel and Gel: COL hydrogels were characterized using rheological measurements. Rheological evaluations were performed to evaluate the influence of the collagen content on the mechanical properties of the Gel hydrogel. The amplitude dependence of modulus is evaluated in Fig. 6(a, b). To about 10% strain, all hydrogels

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showed the linear behavior of G′. Outside the LVR, G′ decreases that are due to large deformations which indicating structure breakdown. The frequency sweep measurements were represented based on the LVR. The frequency dependence of modulus is presented in Fig. 6(c, d). G′ was higher than G″ and independent of frequency, which is typical for ideal gels [45,46]. G′ and G″ of the Gel: COL hydrogel has a more linear behavior than the Gel hydrogel. Especially, G″ of the Gel hydrogel was mostly non-linear; possibly due to the non-homogeneity of the structure. The complex modulus (G*) values of the hydrogels are showed in Table 1. The G* of the Gel: COL hydrogel was more than that of the Gel (non-significant statistically). The loss tangent (Gel: tanδ = 0.061, and Gel: COL: tanδ = 0.049) was lower than 0.1 for all hydrogels, making them labeled as a strong hydrogel. 3.5. Contact angle characterization The wettability of a liquid or solid surface depends on its chemical composition. Time dependent variations of contact angel of the hydrogels are shown in Fig. 7. Water spreading on the Gel hydrogel emerged within 60 s, and the contact angle decreased from 126.87° to 96.91°. As well as, the contact angle of Gel: COL decreased from 114.71° to 70.9°. Therefore, we found that Gel: COL is more hydrophile than Gel hydrogel. Collagen has a complex chemical structure containing both amine and carboxyl groups. This has resulted in more polar content of the Gel: COL hydrogel than that of the Gel hydrogel; so, as a result, Gel: COL hydrogel showed a lesser contact angle. Adding these hydrophilic groups increased water absorption and hydrophilicity of the hydrogels. This increase in water content increased in solid-liquid contact time. In previous works [16], Liu et al. prepared the crosslinked collagen (COL)-gelatin (Gel)-hyaluronic acid (HA) films by EDC and NHS. The test on the physical and biological properties resulted that the mass ratio of Col: Gel: HA = 6:3:1 has acceptable hydrophilicity and mechanical properties. Collagen and hyaluronic acid have an amine, hydroxyl and carboxyl groups. So, the hydrophilicity of the films increased. We also showed that by adding collagen to gelatin the hydrophilicity increased.

used human bone marrow mesenchymal stem cells (hBM-MSCs). Scaffolds' Cell attachment is crucial to precede other cellular phenomena such as spreading, migration, and differentiation. After 48 h, cell attachment on the Gel and Gel: COL freeze-dried hydrogels were investigated by SEM and is shown in Fig. 9. Cells were adhered and flattened on the surfaces of all the freeze-dried hydrogels (cytoplasmic projections of the cells are figured with white arrows). For H&E staining, three 100 × 100 μm2 distinct areas are chosen randomly and cell nuclei (the nucleus of the cells stained by hematoxylin look darker than cytoplasm which stained by eosin) were counted to compare the cell growth and proliferation of on each hydrogel. By comparing H&E staining, an increase of 3.35 folds of cells on the Gel: COL scaffold is significant. An average number of the counted nuclei is 294.3 for the Gel: COL scaffold and 87.6 for the Gel scaffold (Fig. 10).Morphology of the hBM-MSCs on the top of Gel and Gel: COL hydrogels were investigated using phase contrast microscope, as shown in Fig. 11; hBM-MSCs were successfully spread in Gel and Gel: COL hydrogels after 3 and 7 days. According to SEM micrographs and phase contrast microscope images, the cell attachment is excellent on the surface of blend samples; collagen improves the biocompatibility of pure hydrogel and consequently advances cell adhesion on the blend sample. Biocompatibility is essential for implanting biomaterials [64]. As shown in Fig. 12, at all-time points, cell viability was comparable between hydrogels and control group. The proliferation of all MSCs (culture times: 1, 3 and 7 days) in Gel: COL hydrogel was significantly different from that in Gel hydrogel. However, Cell viability was not significantly different between hydrogels and control group at day 1 but, at day 3, there was significant difference between the Gel: COL hydrogel and Gel hydrogel and control group (*p b 0.05). Moreover, hBMMSCs in Gel: COL hydrogel proliferated to the 7th day and reached 1.81 times increment, while cells cultured in the Gel increased 1.29 at the 7th day compared to the control group (*p b 0.01). The Gel: COL hydrogel showed increased viability and proliferation with time compared to the Gel hydrogel. Our result demonstrated that the presence of collagen in gelatin structure affected the cell survival and Gel: COL hydrogel is an appropriately biocompatible hydrogel for corneal tissue engineering applications.

3.6. Optical measurements- light transmittance Fig. 8 shows the light transmittance curve of the Gel and Gel: COL hydrogel when soaked in PBS. The light transmittance of the Gel and Gel: COL hydrogels, calculated at an average of 400–700 nm in the visible region [43], was about 62.53% and 80.43% respectively. The transmittance of the Gel: COL hydrogel reached its maximum and tended to be continuous (similar to a native cornea). The results also indicated that Gel: COL hydrogel had more transparency than Gel based hydrogels. Hydrogel films photographed to show their optical clarity and transparency. Gelatin is a desirable material in cornea tissue engineering because it is easy to access, relatively inexpensive, can be taken from a variety of sources and results in appropriate transparency [62]. Bektas et al. prepared methacrylate gelatin (Type A porcine skin) hydrogels (GelMA) through photo-polymerization with keratocytes to use as the corneal tissues. The transparency of the hydrogels was around 90% at 700 nm (similar to the native cornea) [57]. In Niu et al. study, transparent thin gelatin gel (TGG: 10%) scaffolds (fully flexible) were made and functionalized with heparin to support transplantation of human corneal endothelial cells (HCECs). After crosslinking, TGG hydrogels, with and without heparin, were transparent (with over 95% transmittance in the visible light range).Acellular human cornea stroma has between 88% and 90% transmittance at 400 nm–600 nm wavelengths, respectively [63]. 3.7. Cell evaluations Stem cells sources used for corneal regeneration are diverse. Mesenchymal stem cells and limbal stem cells (from the eye itself) are the two main stem cell types used for cellular therapy [40]. In this study, we

4. Conclusion The goal of this study was to investigate the effects of collagen on the physic-chemical and biological properties of gelatin hydrogel. We fabricated two type Gel based hydrogels cross-linked using EDC/NHS through adding collagen type I. We investigated their physical and mechanical properties and analyzed their potential as cell viability and attachment to the hBM-MSCs. This study shows that the biocompatible natural based Gel: COL hydrogel has a highly porous structure; with interconnected pores which makes it suitable for cell attachment and infiltration. It has appropriate hydrophilicity, water content angle, optical performance, and mechanical properties, too. Also, the Gel: COL hydrogel showed long-lasting increased viability and proliferation compared to the Gel hydrogel.

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