Oxygen-plasma treatment-induced surface engineering of biomimetic polyurethane nanofibrous scaffolds for gelatin-heparin immobilization

2.1 Materials

PU (Estane 5701) was purchased from the BF Goodrich Chemical Ltd. (Kallo, Belgium). Gelatin (M_w 40–50 kDa), acetic acid (AA, M_w 60.05 g/mol), tetrahydrofuran (THF, M_w 72.11 g/mol), dimethylformamide (DMF, M_w 73.09 g/mol) and glutaraldehyde (25%, d 1.058 g/cm³) was purchased from Merck Co. Ltd. (Darmstadt, Germany). Heparin (Grade I-A) was purchased from Sigma Co. Ltd. (MO, USA). Phosphate buffered saline (PBS, powder) was purchased from Aprin-ATD Co. Ltd. (Tehran, Iran). All reagents were used without further purification.

2.2 Preparation of PU-based nanofibrous scaffolds

PU nanofibrous scaffolds were fabricated by an electrospinning (Full Option Lab ES, Nano Azma Co., Tehran, Iran) technique. For this purpose, a solution of 7% (w/v) of PU was prepared in a solvent mixture of DMF:THF (1:1). The solution was injected using a plastic syringe and injection rate of 0.9 ml/h. The high voltage power supply (20 kV) was applied to the needle tip, which was kept at a distance of 20 cm from a rotating drum at a rotational speed of 280 rpm. The final sample was dried overnight in an oven at room temperature and kept in a desiccator until oxygen plasma surface modification.

The bilayered scaffolds of PU and gelatin-heparin were fabricated by the same technique as a control group while the gelatin-heparin solutions with a concentration of 3.5 and 2% (w/v) in deionized water, respectively, were synthesized after finishing the injection of PU. Gelatin and heparin were injected from a syringe under the following conditions:

Injection rate: 0.6 ml/h, voltage: 18 kV, rotational speed of collector: 250 rpm, distance between collector and needle: 20 cm.

All the bilayered samples were crosslinked by glutaraldehyde vapor for 12 h. Finally, the crosslinked constructs were soaked in NaBH₄ 1% for 30 min and deionized water for 2 h to remove unreacted glutaraldehyde (25). Bilayered PU/gelatin-heparin scaffolds and oxygen plasma-treated fibers after immobilization were introduced as PGHE and PGHG, respectively.

2.3 Plasma treatment and gelatin-heparin immobilization

After the fabrication of the PU-based scaffolds, oxygen plasma treatment was carried out on a plasma generator (EMITECH, K 1050X, Birmingham, UK) for 2 min with an oxygen flow rate of about 15 ml/min, pressure of about 0.13 Pa and power of 100 W. Thirty minutes after plasma treatment, all the samples were soaked in the blend of gelatin-heparin solution with concentration of 3.5 and 2% (w/v) for 1 h and dried in an oven at room temperature for 48 h.

2.4 Characterization technique

2.5 Statistical analysis

All the data were presented as the mean±standard deviation of at least five experiments. Statistical analysis was performed by one-way analysis of variance (ANOVA) and Tukey’s test, with significance reported when p≤0.05. Also for investigation of group normalizing, the Kolmogorov-Smirnov test was used.

3.1 Morphology observation

Biomimetic porous scaffolds are one of the major factors in promoting cellular proliferation and differentiation to regenerate defects. The microstructure of pores and simulation of ECM are influenced by the synthesizing conditions and finally have an effect on the performance of the prepared scaffolds (30). Electrospinning is one of the versatile methods for fabrication of nanofibrous scaffolds. Changing the synthesizing parameters in electrospinning techniques such as voltage, linear or rotational speed, distance, etc., allows the fabrication of fibrous structures with a broad range of diameters and morphologies for different tissue engineering applications (31). In this study, the PU-based nanofibers were prepared via electrospinning and modified with prolonging oxygen plasma treatment and gelatin-heparin immobilization in order to improve cell-scaffold interactions. The effects of grafting gelatin-heparin after oxygen plasma treatment were compared with bilayered electrospinning technique for covering the surface of PU fibers with gelatin-heparin. Figure 1A, B indicate the SEM micrographs of PGHE and PGHG. According to the images, both groups of fibers showed a smooth surface and bead free microstructure. The smooth surface of scaffolds can be a good sign for discouraging clot formation (32). According to the SEM images, interconnected pores have formed between the layers of fibers and have produced network microstructures with approximately 90% porosity. Increasing the diameter of fibers (Figure 1C) after immobilization of gelatin-heparin was not an unexpected result due to covering the surface of PU fibers after soaking in the gelatin-heparin solution. Besides, in PGHE scaffolds, the outer surface is composed of electrospun gelatin-heparin fibers. The presence of gelatin decreases the final diameter of fibers by increasing the charge density of the polymeric solution as Meng et al. (27) observed. As the final diameter of fibers plays a remarkable role in cellular adhesion, migration and proliferation, it can be concluded that fibers distributed within an acceptable size range to support cellular processes.

Figure 1:

Microstructural analysis of PU-based scaffolds.

SEM micrographs of the (A) PGHE and (B) PGHG biomimetic nanofibrous scaffolds, (C) distribution of the fiber diameter in bilayered and surface modified structures.

3.2 Chemical composition

In order to determine the chemical composition of the plasma-treated scaffolds, an FTIR spectrum was performed. Figure 2 demonstrates the FTIR spectrum and the chemical structure of the raw materials and gelatin-heparin immobilized on plasma-treated PU scaffolds. The presence of N-H stretching at 3306 cm⁻¹, carbonyl urethane stretching at 1731 cm⁻¹, CHN vibration at 1526 cm⁻¹, coupled C-N and C-O stretching at 1223 cm⁻¹ and finally C-O stretching at 1079 cm⁻¹ prove the characteristic peaks of PU (33). The gelatin peaks observed at 1650 and 1538 cm⁻¹ for amide I and II (34, 35). The heparin peaks at 3450 and 1650 cm⁻¹ assigned to hydroxyl stretching and bending, respectively. The characteristic peak at 1040 cm⁻¹ is related to –SO₃ stretching of heparin (36).

Figure 2:

Chemical characterization of scaffolds.

FTIR spectrum and chemical structure of PU (A), gelatin (B), heparin (C), bilayered scaffolds (D), and immobilized gelatin-heparin on plasma treated-PU nanofibers (E). Schematic of the chemical structure of raw materials (F).

Gelatin is a polyelectrolyte with ionizable groups and it was found that it has a positive charge in acidic media. In contrast, heparin is a negatively charged glycosaminoglycan (37). Herein, the mixture of gelatin and heparin interact with each other via electrostatic interactions. The absorption peak of the mentioned electrostatic interaction at 1339 cm⁻¹ is related to the sulfamide. The COS and S=O vibration are indicated at 1240 and 893 cm⁻¹. Besides, the stretching vibration of pyranose is shown at 1060 cm⁻¹ (37). Additional peaks around 1470–1570 cm⁻¹ attributed to the crosslinking reactions between the aldehyde groups on glutaraldehyde and amide groups of gelatin to form aldimine linkage in bi-layered samples (38).

Different studies (21, 39, 40, 41) indicated that plasma oxygen resulted in removing organic contaminants, the introduction of higher numbers of oxygen-containing functional groups, peroxides and free radicals on the PU fibers owing to the radical reaction between the chain backbone of the polymers and the oxygen flow. Based on Cen et al. (42) investigation, during the plasma oxygen process, the oxygen gas is ionized and an activated surface is created as a result of bombing the fibers via energetic and ionized species. Chemically reactive functional groups on the surface are attached and covalently immobilize biological macromolecules such as heparin or gelatin. Immobilization of heparin, PGE1 or albumin/heparin conjugates is an established method to improve thrombogenicity and hemostasis of polymeric surfaces (43). The FTIR spectrum after immobilization of gelatin and heparin contained all the above-mentioned peaks related to PU and gelatin-heparin interactions. But the overlap of symmetric stretching of S=O (1000–1100 cm⁻¹) in the saccharide groups of heparin and C-O-C bond of PU led to the observation of two shoulders at 1080 and 1110 cm⁻¹. The same results were obtained in the investigation by Kim et al. (40) and Kang et al. (39). According to their investigation, methyl ether groups of PU were covalently coupled with heparin.

3.3 Mechanical strength

The strength, elasticity and degradation rate of polymeric nanofibers determine their mechanical stability. In this case, the biomimetic fibers should establish a balance between biostability and degradation rate until the reconstruction process is finished (44). The stress-strain curves and ultimate tensile strength of bilayered electrospun scaffolds and plasma-treated fibers after immobilization are shown in Figure 3A, B. Accordingly, PGHG fibers showed 1.2 times higher strength compared with PGHE scaffolds. Higher fiber diameter is one of the reasons for achieving this result; also, PU is the main component of the constructs and plays an undeniable role in supplying the required flexibility (45) in the vascular graft due to thermoplastic elastomer nature which may also affect strength values. Reduction in tensile strength of PGHE fibers can originate from a crosslinking reaction by glutaraldehyde. This process decreases the final area of fibers and so reduces toughness and enhances modulus and produces stiff and brittle samples. Consequently, bilayered scaffolds with fine fibers show less resistance to periodically loads. Zhu et al. (46) observed that the increasing degree of crosslinking with glutaraldehyde increased brittleness. The balance between tensile strength to tolerate stresses due to pumping blood in the human body is the critical property for tissue regeneration that provides by plasma-treated samples after immobilization.

Figure 3:

Grafting gelatin-heparin on the modified fibers showed 1.2 times higher strength compared with produced fibers via electrospining.

(A) Stress-strain curves and (B) Ultimate tensile strength of the PGHE and PGHG scaffolds.

3.4 In-vitro PBS absorption and hydrolytic biodegradation

The hydrophilicity of the polymeric scaffolds has a strong effect on the amount of absorbed fluid. Additionally, porous samples such as electrospun scaffolds absorb greater amounts of fluids compared with compact structures with the same materials (47). Weak hydrophilicity of PU (48) resulted in the limited application in some biomedical fields. This behavior can be controlled by different strategies such as the addition of hydrophilic polymers. As can be observed from water absorption results in Figure 4A and Table 1, the capacity of absorption increases gradually from 2 to 6 h. Besides, the PGHG scaffolds absorb PBS 1.2 times higher than PGHE samples after 6 h. It shows the ability of the surface modification to alter the surface hydrophilicity. In PGHE scaffolds lack of complete covering of superhydrophobic polymer influences on the reduction of swelling ratio. Accordingly, homogeneous covering the PU fibers by gelatin-heparin after plasma treatment improved absorption capacity owing to hydrophilic functional groups such as amine, carboxylic acid and hydroxyl groups in the gelatin and heparin chemical structure. At the interface of modified-fibers and water molecules, the hydrophilic chain acts to generate high absorption capacity due to the aminolysis reaction as used by Zhu et al. (49). The results obtained illustrate that modified-polymeric samples are well able to support the diffusion of nutrients and excreting by-products in order to support cellular interactions. The interconnected microstructure is created by the electrospinning technique and the remarkable absorption capacity provided by both porous microstructure and hydrophilic materials can facilitate revascularization during tissue formation as seen in the study by Hariraksapitak and Supaphol (50).

Figure 4:

Modification of fibers and grafting hydrophilic polymers on the surface led to higher absorption capacity and increased degradation ratio.

(A) PBS absorption capacity and (B) biodegradation rate of nanofibrous scaffolds in different time intervals.

As polymeric scaffolds are not permanent, biodegradation is one of the most important properties of biomimetic scaffolds which leads to guiding cellular growth, production of extracellular matrix, and finally regeneration of defects. In this case, biocompatibility of by-products after the degradation process is one of the critical issues in tissue engineering studies (51). In-vitro hydrolytic biodegradation of PGHE and PGHG scaffolds has been shown in Figure 4B and Table 1 during 10 days immersion in the PBS solution. Results demonstrate that gelatin-heparin immobilized samples degrade faster than bilayered one as a function of time due to providing high values of hydrophilic chains and rapid swelling ratio. PU controls the biodegradation rate and resists degradation because of the superhydrophobic nature and water-resistance behavior; consequently, covering the fibers with hydrophilic functional groups enhances the biodegradation rate. However, it expects collapsing the samples to suddenly occur during the degradation process because of brittle structure that produced by glutaraldehyde interaction in PGHE constructs even though it shows the slower biodegradation rate. Moreover, covalent immobilization of gelatin-heparin to PU fibers after plasma treatment creates a more stable microstructure compared with other immobilization techniques than that investigated by Chalovich et al. (52).

3.5 In-vitro heparin-release level

Delivery of the anticoagulant small molecular weight drugs to the exact site of the defect can be an effective strategy to reduce coagulation after scaffold implantation. Hence, in this study heparin immobilized on the plasma-treated nanofibrous scaffolds owing to the presence of this material in the ECM components and take parts in angiogenic process (53). Figure 5 represented calibration and cumulative release curves of heparin after 10 days in the PBS solution. According to the results, increasing the hydrophilicity of PGHG scaffolds followed by higher absorption and biodegradation values, influence on a heparin-release level so the polymeric scaffold releases more drug compared with PGHE fibers.

Figure 5:

Cumulative heparin-release level of PGHE and PGHG scaffolds during 10 days and heparin calibration curve in different concentrations.

Uptake of release media leads to a burst release of heparin after 1 day. Then, constructs release heparin gradually with a steady gradient and is proportional to the rate of biodegradation for 10 days. The PGHG scaffold follows a constant release rate during days 4–6 and 8–10 and on other days the release rate increases gradually. This event can be owing to the stronger covalent bonding of heparin to modified surface in PGHG compared with cross-linked heparin in PGHE that leads to controllable biodegradation and affects heparin releasing behavior. It is expected that this controllable release rate at the implantation site produces a hemocompatible biodegradable implant.

3.6 Cell-scaffold interactions

The main purpose of scaffolds is to promote cellular growth to the formation of desired tissues or organs. Nanofibrous scaffolds allow cells to migrate and proliferate well, but surface modification is a proper strategy to increase the effectiveness due to the induced protein adsorption that follows cellular interaction. Moreover, modified scaffolds facilitate drug immobilization to modulate cell-scaffold interactions (53). A 48-h L929 fibroblast cell attachment to PGHE and PGHG biomimetic nanofibers is shown in Figure 6A, B by SEM micrographs. Images indicate that polymeric fibers could support cellular adhesion while cellular spreading improved after oxygen plasma treatment and immobilization of gelatin-heparin. PGHE constructs showed spindle-shaped cells while same cells spread entirely on the surface of PGHG samples. Elongated cells prove that ameliorated hydrophilicity of scaffolds can enhance cellular interactions and they will secret more ECM for the rapid regeneration process. In fact, covering the surface of the superhydrophobic polymer (PU) allows more cellular adhesion and allows connection of scaffolds and surrounding tissue. Moreover, this treatment will improve endothelial cell adhesion to vascular graft surfaces as in other investigation (54). Besides, hemocompatible PU and anticoagulant factor (heparin) will prevent clot formation during tissue reconstruction. Results of dehydrogenase activity of mitochondria after 1, 3 and 5 days in Figure 6C demonstrated that both PGHE and PGHG scaffolds are biocompatible due to the viability of more than 85% cells compared to the control group. According to the findings, PGHG scaffolds are more cytocompatible because of nearest viability level to the control group. This enhanced cellular proliferation could be owing to chemical alterations induced by plasma treatment and gelatin-heparin immobilization as seen in Pappa et al. (21) observations. Weak cellular adhesion to bilayered PGHE scaffolds and difficulty in cellular transfer leads to cellular death by-passing the time of incubation. However, PGHE samples are not cytotoxic after 7 days. In short, the higher hydrophilicity of PGHG constructs leads to better performance in the interaction with cells as noted above.

Figure 6:

Grafting hydrophylic polymers on the surface of plasma treated nanofibers improved cell-scaffold interactions.

Attachment of L929 fibroblast cells seeded on (A) PGHE and (B) PGHG nanofibrous scaffolds after 48 h, (C) The viability of L929 cells on the PGHE and PGHG scaffolds at Day 1, 3 and 5 after seeding was assessed using the MTT assay (O.D. value 570 nm).