Comparative analysis of gelatin and hydroxyethyl cellulose scaffolds crosslinked by silane coupling agent

1. Introduction

Scaffolds are one of the major components in tissue engineering, which are designed to prepare an artificial extracellular matrix (ECM) to guide three-dimensional (3D) tissue formation and support cell attachment [1, 2]. Therefore, an ideal scaffold should mimic the characteristics of the natural ECM. The native ECM is contained of a complex network of structural and extracellular macromolecules, such as enzymes, glycoproteins, and collagen, that provide structural and biochemical support to surrounding cells. These components are arrayed into a fibrous matrix. It also provides resident cells with specific ligands for cell adhesion and migration and modulates cell proliferation and function [3, 4]. Several researchers have studied to develop biomimetic scaffolds able to operate as native ECM. Hydrogels are typically composed of a natural polymer, such as collagen, gelatin, hyaluronic acid (HA), or cellulose, and have three-dimensional networks of hydrophilic polymer chains. Hydrogels can absorb and swell large quantities of water without dissolution. Furthermore, they feature unique mechanical and structural properties, which are similar to those of tissues and the ECM of the skin. The unique properties of hydrogels, make them suitable for applications as biomaterials for scaffolds in tissue engineering, drug delivery systems, therapeutics, imaging, and medical devices [5-7]. A wide assortment of scaffolds based on natural and synthetic polymers have been used for the engineering of soft tissues. Natural polymers are practically prominent to non-informational synthetic polymers because they provide structural and chemical signals (such as the Arg-Gly-Asp sequence), which promote cell attachment [8].

Among them, gelatin (G) is commercially available and a low cost biomaterial that has picked up interest in biomedical engineering, mainly due to its biodegradability. G is obtained by physical and chemical degradation or thermal denaturation of collagen, the most abundant protein in the body, present in most connective tissues like skin, bone, and tendon [9]. G can effectively increase the process of granulation and epithelialization, and different studies have demonstrated that G is beneficial during the process of tissue regeneration for the adhesion, migration, and the growth of cells [10-12]. However, the drawbacks, such as strict swelling after imbibition of water, restrict its applications. The malformation after severe swelling will make the scaffolds unable to fit the wound [13].

Cellulose is the most widespread naturally occurring polymer of glucose with various attractive advantages such as biocompatibility, non-expensive, non-toxicity, and biodegradability. However, due to the tight network between the chains of cellulose molecule, it is insoluble in water, as well as many other organic solvents. Cellulose-based derivatives have been developed to improve the solubility of cellulose, including hydroxyethyl cellulose (HEC), hydroxypropyl methylcellulose (HPMC), methylcellulose (MC), sodium carboxymethylcellulose (NaCMC), and ethyl cellulose (EC) [14]. HEC is one of the most important commercial soluble cellulose derivatives. Due to its desirable biocompatibility with non-immunogenicity and low toxicity, it can be a candidate in biomedical applications such as wound dressings and wound healing [15, 16].

Chemical and physical crosslinking methods have been used to increase G and HEC stability in aqueous media. Physical crosslinking is based on non-covalent interactions between macromolecular chains such as van der Waals forces,  hydrogen bonds, electrostatic interactions, and hydrophobic association [17]. Their main disadvantage arises from the difficulty to gain a desirable crosslinking degree. In contrast, chemical crosslinking provides mechanical integrity for a longer period, which is suitable for implantable materials in the human body. The main limitation in the use of these products is the possible presence of some unreacted cross-linker inside the scaffold with following formation of toxic products during in vivo biodegradation. Due to this reason, developing interest has been recently obtained by enzymatically [18] or naturally derived crosslinking agents, with a low toxicity [19].

 ɣ -Glycidoxypropyltrimethoxysilane (GPTMS) is a silane coupling agent, which has epoxy and methoxysilane groups. The oxirane rings on the GPTMS molecules react with the amino groups on the G chains [3], hydroxyl groups of HEC [20], and hydration of the trimethoxy groups on the GPTMS forms pendent silanol groups (Si-OH) through an acid-catalyzed reaction. Then Si-O-Si bonds were formed through condensation of two Si-OH occurring mainly during the solvent evaporation period of the membrane formation process. The Si-O-Si linkages formed from the condensation reaction provided inter-chain covalent bonds to result in a cross-linked structure[21]. GPTMS was mainly used to cross-linked chitosan films and membranes for different applications [22, 23].

In this study, G and HEC porous scaffolds cross-linked with GPTMS were fabricated and characterized, to compare their rheology behavior, porosity, swelling ratio, and degradation rates as

favorable tissue engineering scaffolds for skin regeneration.

1. Materials and methods

2.1.   Materials and scaffolds fabrication

Hydroxyethyl cellulose (average Mw = 90,000 Da) was purchased from Sigma Aldrich (USA). ɣ- glycidoxypropyltrimethoxysilane (GPTMS) and Gelatin powder were purchased from Merck (USA). Phosphate buffered saline (PBS) powder was purchased from Aprin company (Tehran, Iran).

  After measuring the weights of G and HEC powders (each polymer = 0.4 g), they were dissolved completely in 20 mL deionized water by stirring, to achieve a homogeneous solution at room temperature, respectively. Then, the GPTMS was added drop-wisely into each solution with a 10% (w/w) of polymers and further stirred (rpm: 100) for 48 h at 60˚C (Fig. 1). The obtained solutions poured into 8 cm petri-dishes, air-dried for 48 hours, and pre-frozen at -20°C and -80˚C in a deep freezer, respectively. The frozen samples were lyophilized (FD-10, Pishtaz Engineering Co. Iran) at a temperature of -58 °C and pressure 0.5 torrs for 48 hours. The lyophilized scaffolds are shown in Fig. 2.

Figure 1: The scheme of the sample preparation process.

Figure 2: freeze-dried scaffolds of G-GPTMS and HEC-GPTMS.

2.2.     Characterization of the Scaffolds

2.2.1.      Rheology Testing

The rheological properties of the samples were examined at room temperature in oscillation mode using an Anton Paar Rheometer, MRC 301, with parallel plates. In the test, 20 ml of each prepared sample was placed on the lower plate, while the distance between the plates was adjusted to be 1 mm. Measurements were made in dynamic mode to determine the viscoelastic properties of the solutions. The storage modulus (G’) and the loss modulus (G”) were plotted as a function of the angular frequency. First, the measurements were made in a variable strain state with constant frequency and temperature to obtain the highest strain of the linear viscoelastic region. The measurements were then carried out in a constant strain of 0.01% with a variable frequency of 0.1–100 s⁻¹. The complex viscosity curve as a function of angular frequency was obtained in the same way.

2.2.2.   Microstructure Analysis

The freeze-dried scaffolds affixed on carbon stubs and sputter-coated with a thin layer of gold. The surface morphology of scaffolds was examined using a scanning electron microscope (SEM) at the acceleration voltage of 20 kV. Pore dimension was quantified by analyzing the SEM images of the sample's surface using image j software. Five images for each sample type were used for the calculation of average pore sizes.

2.2.3.   Swelling ratio and degradation rate

 The swelling ratio and degradation behavior of scaffolds were measured in PBS (pH=7.4) at a constant temperature of 37˚ C, to simulate the body temperature and application of bio-medicals. The freeze-dried hydrogel (n=3) was immersed in PBS. For swelling ratio after removal from the PBS at defined intervals, they were dried on filter paper for 1 min until no dripping PBS was observed and then weighed. The content of the PBS in the swollen membranes was calculated by the following equation:

Swelling Ratio (%) = ( W_s – W_d )/ W_d × 100                                 (1)

where W_d is the weight of the freeze-dried sample and W_s is the weight of the swollen membrane.

  The degradation rate was evaluated after 1, 3, 5, and 7 days. After drying at 37^◦C in an oven, samples were weighed again and the degradation percentage was calculated as:

Degradation rate (%) = ( W₀ - W_d )/ W₀ × 100                              (2)

where W_d is the dried sample weight and W₀ is the weight of the freeze-dried sample.

1. Result and Discussion

The rheological properties of the G-GPTMS and HEC-GPTMS were investigated as function of the angular frequency, as shown in Fig. 3. The storage (or elastic) modulus G’ can be used to describe the elasticity or stored mechanical energy in response to stress, and the loss modulus G” describes the viscosity or energy dissipated in response to stress. The HEC-GPTMS spectra revealed that the loss modulus was greater than the storage modulus at all angular frequencies. The G-GPTMS spectra showed that this sample only had loss modulus. We find that G” increased as the angular frequency increased, which indicates that both samples formed a loose structure, that the samples exhibited solution and flowable behavior [24]. Both modulus of HEC-GPTMS were larger than G-GPTMS which indicates that the HEC polymer chain contributed to the strength of the sample [25, 26].

Cross-linking and sol-gel mechanisms of G-GPTMS or HEC-GPTMS (R-OH functional groups of HEC instead of G’s R-NH₂) are shown in Fig. 4. The sol-gel reaction leads to the formation of water molecules, which encircled in Fig. 4. It seems due to the presence of these water molecules, samples had solution behavior. To omit the excess water, the samples air-dried for 48 hours before freezing.

Fig. 5 illustrate the complex viscosity in both samples. Results indicated also this parameter depended on frequency, wherein HEC-GPTMS complex viscosity decreased as frequency increased and showed shear thinning behavior, which indicated the non-newtonian behavior. But in G-GPTMS viscosity increased by frequency and had a constant behavior of shear thickening.

Figure 3: Viscous and elastic modulus vs. angular frequency of solutions.

Figure 4: Cross-linking and sol-gel reaction of G-GPTMS or HEC-GPTMS.

Figure 5: Complex viscosity of solutions as a function of angular frequency

SEM analysis was performed on the surface of scaffolds to evaluate the sample morphology. Fig. 6 depicts SEM images of cross-linked scaffolds. As it can be observed, the porous scaffolds showed a foam-like morphology with interconnected pores, characterized by the presence of super-porous structures with dimensions of up to 1 micrometer, which are connected to form an open channel system. The mean pore size was found to be around 54 µm for the G-GPTMS scaffold and 120 µm for HEC-GPTMS. These porous structures can act as water channels, so that the scaffolds can take up high quantities of water, leading to a high swelling ratio [27].

Comparing with reports [16, 28], the morphological features and pore size of prepared scaffolds are suitable for skin tissue engineering.

Water content is one of the main factors of biomaterials that affecting the biocompatibility; scaffolds with a high degree of swelling have a large surface area/volume ratio, which may favor cell infiltration into the porous structure and cell attachment on the surfaces [28]. A general property of hydrogels is swelling when exposed to external solvents, which arise due to osmotic pressure, electrostatic forces, and a viscoelastic restoring force [6].

The swelling behavior of scaffolds is shown in Fig. 7 that the G-GPTMS and HEC-GPTMS freeze-dried scaffolds increased their weight immediately after the immersion in PBS with showing a high degree of water uptake after only 0.5 h incubation in PBS.

The swelling degree was then approximately stable up to 6 hours incubation time, showing values of 1730% for G-GPTMS and 3100% for HEC-GPTMS. GPTMS with longer molecular chains (epoxy) would provide the structure with more ability to expand and swell [29]. Hydrophilic groups of both polymers and GPTMS (OH, due to ring-opening of epoxy) were other reasons for enhancing this ratio. The ability of the scaffolds to absorb a large quantity of water helps to hydrates the wound and absorbed excess wound exudate [30].

The degradation profile of G-GPTMS and HEC-GPTMS scaffolds is shown in Figure 8. On fifth day, HEC-GPTMS underwent 71±13% degradation whereas G-GPTMS showed only 6.5±2% degradation. On seventh day, the percentage of degradation of HEC-GPTMS and G-GPTMS was 97±3% and 8±1%, respectively. A fast degrading material cannot support cell proliferation whereas a

Figure 6: SEM graphs from surfaces of scaffolds, a) and B) G-GPTMS, c) and d) HEC-GPTMS.

Figure 7: Swelling behavior of crosslinked porous scaffold as a function of time. Bars indicate standard deviation (n=3).

Figure 8: degradation behavior of both crosslinked scaffolds. Bars indicate standard deviation (n=3).