KR102668218B1 - An injectable hydrogels with antibacterial efficiency as wound dressing for improved cutaneous wound healing - Google Patents

Abstract

The present invention relates to a hydrogel manufactured by adding spherical nanocellulose to a carboxylmethylchitosan matrix, the use of the hydrogel for antibacterial, skin regeneration and wound healing, and a hydrogel prepared by dissolving 5% (w/w) of carboxylmethylchitosan in water. It relates to a hydrogel method comprising preparing a caroboxylmethylchitosan matrix and adding 1 to 4% (w/w) of spherical nanocellulose to the caroboxylmethylchitosan matrix.

Description

Injectable hydrogel for antibacterial and wound healing and manufacturing method thereof {An injectable hydrogels with antibacterial efficiency as wound dressing for improved cutaneous wound healing}

The present invention relates to an injectable hydrogel for antibacterial and wound treatment and a method for manufacturing the same, and more specifically, to a conductive injectable hydrogel based on spherical nanocellulose derived from pine trees with excellent antibacterial efficiency as a wound dressing for improving skin wound healing. will be.

Skin is the largest organ in vertebrates and serves as the outermost barrier between the body and the surrounding environment. It consists of the epidermis, dermis and subcutaneous layers with a typical thickness of 0.5-4 mm and plays an important role in preventing damage to the body from various aspects, including chemical and mechanical factors. Loss of the structural integrity of the skin due to burns, abrasions, ulcers, and lesions causes wounds. It is difficult to achieve complete skin restoration through skin grafting due to limited donor sources and the antigenicity of donor tissues.

A variety of pharmacological, physical and phytotherapeutic approaches have been used to improve skin recovery. Tissue engineering approaches have received considerable attention for the repair and regeneration of damaged tissues, where constructs have been developed to mimic the native environment of the extracellular matrix (ECM). The developed structures must be biocompatible, biodegradable and appropriately porous to improve cell growth, differentiation, adhesion and migration. Electrospinning, solvent casting, gas foaming, melt melting, freeze-drying, and 3D printing approaches have been applied to fabricate porous scaffolds for a variety of applications. Alginate, chitosan, gelatin, collagen, poly(caprolactone), poly(lactic acid), and poly(lactic acid-co-glycolic acid) have been widely studied in tissue engineering applications due to their excellent physicochemical properties.

It is well known that the skin is sensitive to electricity and has an electrical conductivity of 2.6 - 1×10 ^-4 mS depending on the skin tissue composition. Therefore, electroactive materials with similar conductivity and adhesive potential are expected to promote rapid wound healing. Adhesion has the characteristic of allowing closed contact between the implanted material and the defective tissue without using any other substrate.

Hydrogel is a polymer structure containing a large amount of water and is used not only for medical purposes such as wound dressings and bioscaffolds, but also for refrigerant purposes such as ice packs.

[Prior patent literature]

Republic of Korea Patent Publication No. 10-2021-0147652

The present invention was created in response to the above-mentioned need, and the purpose of the present invention is to provide an injectable hydrogel for antibacterial and wound treatment.

In order to achieve the above object, the present invention provides a hydrogel prepared by adding spherical nanocellulose to a caroboxymethylchitosan matrix.

In one embodiment of the present invention, the caroboxyl methyl chitosan matrix is preferably prepared by dissolving 5% (w/w) caroboxyl methyl chitosan in water, but is not limited thereto.

In another embodiment of the present invention, the hydrogel is preferably manufactured by adding 1 to 4% (w/w) of spherical nanocellulose to the caroboxylmethylchitosan matrix, but is not limited thereto.

In a preferred embodiment of the present invention, the hydrogel is more preferably manufactured by adding 1% (w/w) of spherical nanocellulose to the caroboxylmethylchitosan matrix, but is not limited thereto.

In one embodiment of the present invention, the spherical nanocellulose is preferably derived from pine powder, but is not limited thereto.

In one embodiment of the present invention, the hydrogel preferably has an adhesive strength in the range of 11.84 to 18.32 kPa, but is not limited thereto.

In another embodiment of the present invention, the hydrogel preferably increases fibronectin and collagen 1A gene expression, but is not limited thereto.

The present invention also provides an antibacterial composition containing the hydrogel of the present invention as an active ingredient.

In one embodiment of the present invention, the composition preferably has antibacterial activity against Bacillus subtilis, but is not limited thereto.

Additionally, the present invention provides a composition for skin regeneration comprising the hydrogel of the present invention as an active ingredient.

Additionally, the present invention provides a composition for wound healing comprising the hydrogel of the present invention as an active ingredient.

In addition, the present invention prepares a caroboxyl methyl chitosan matrix by dissolving 5% (w/w) of caroboxyl methyl chitosan in water, and adds 1 to 4% (w/w) of spherical nanocellulose to the caroboxyl methyl chitosan matrix. A hydrogel method comprising the step is provided.

The present invention will be described below.

The present invention aims to evaluate the wound healing potential and antibacterial activity of a hydrogel scaffold made of water-soluble CMCS and s-NC. The developed hydrogel scaffolds were characterized by various spectroscopic techniques. Recovery strength, swelling and disintegration potential were investigated. The biocompatibility of the developed hydrogel scaffolds was investigated using co-cultures of human dermal fibroblasts (HDF), human keratinocytes (HaCaT), and human umbilical vein endothelial cells (HUVEC) via WST-1 assay in vitro.

Conductivity and antibacterial potential were also monitored. Finally, the wound healing process was evaluated using rats 14 days after scaffold implantation. We anticipate that the fabricated hydrogel scaffolds are versatile and can be explored in skin tissue engineering for rapid wound closure.

The injectable adhesive hydrogel of the present invention was developed by incorporating s-NCs into a CMCS matrix for rapid subcutaneous wound healing applications. The developed hydrogel of the present invention was interactive and a significant improvement in mechanical strength was observed. The hydrogel exhibited thixotropic behavior and ~85% recovery in viscosity was observed after deformation at high shear rate (1000 s ^-1 ). The hydrogel was easily injectable and maintained its pre-designed shape. Due to the strong interactions, the prepared hydrogels exhibit excellent adhesion to various surfaces (glass, plastic, rubber, stainless steel, and human skin). Enhanced bioactivity was observed in composite hydrogel scaffolds compared to pure polymer scaffolds.

Additionally, the biocompatibility of our developed hydrogel scaffolds was individually monitored and co-cultured with HDFs, HaCaT cells, and HUVECs. No side effects on skin cells were observed with the developed hydrogel scaffold, indicating biocompatibility. The co-culture method also improved skin cell viability, suggesting that the developed hydrogel has the potential to accelerate skin cell regeneration.

The hydrogel scaffold developed in the present invention showed higher expression of skin tissue-related genes and protein markers than the control group. The composite hydrogel showed superior antibacterial activity than the control and pure polymer hydrogels because the bacterial cell membrane was more damaged by electrostatic interaction. The developed hydrogel scaffold showed improved wound healing potential compared to the control group, which was substantially higher in the s-NC added scaffold. Histological analysis showed that the development of new tissue occurred in all groups. However, their density was high in s-NC integrated scaffolds. In addition, thicker granulation tissue was observed in the scaffold to which s-NC was added, demonstrating enhanced wound healing efficacy by stimulating the thickness of granulation tissue. Based on these results, we hypothesize that the developed hydrogel has multiple potentials and can be applied in skin tissue engineering for rapid wound closure.

Figure 1 is a schematic diagram of the prepared hydrogel and its potential application for skin wound healing;
Figure 2 is a diagram showing the characteristics of pine-derived cellulose and spherical nanocellulose (s-NC).
(a) FTIR spectra of cellulose and s-NC;
(b) AFM image of s-NC, and
(c) XRD patterns of cellulose and s-NC.
Figure 3 is a diagram showing the structural and morphological analysis of the developed hydrogel support.
(a) FTIR spectrum of the hydrogel;
(b) XRD pattern of the hydrogel support;
(c) SEM morphology of the hydrogel scaffold, and
(d) Swelling force of the developed hydrogel scaffold.
Figure 4 shows the FTIR spectrum of the hydrogel developed in the absorption region of 4000-2500 cm ^-1 .
Figure 5 is a diagram showing the mechanical strength evaluation of the hydrogel developed at 25 °C.
(a) Storage and loss coefficients of the hydrogel;
(b) the corresponding viscosity composite of the hydrogel;
(c) Analysis of the thixotropic behavior of the developed hydrogel at 25 °C, and
(d) Image of the injected hydrogel.
Figure 6 (a) Investigation of adhesion with different surfaces of developed hydrogels, (b) Schematic demonstration of adhesion measurements, (c) Adhesion shear force versus displacement of hydrogel by plastic, (d) Adhesion hydrogel with plastic substrate; and (e) illustration showing possible interactions of hydrogels with other surfaces;
Figure 7 (a) Analysis of protease enzyme degradation rate of the developed hydrogel scaffold, (b) Bioactivity of the developed hydrogel scaffold after 14 days of SBF treatment,
Figure 8 shows the electrochemical analysis of the hydrogel developed at a scan rate of 25 mV/s;
Figure 9 is a diagram showing the evaluation of the skin cell proliferation potential of the developed hydrogel scaffold, (a) HDF, (b) HaCaT, (c) HUVEC, (d) co-cultured cells, and (e) co-cultured cells. Viability of fluorescent forms.
Figure 10 is a diagram showing the evaluation of skin-related genetic markers in HDF, (a) HDF, (b) HaCaT, (c) HUVEC genetic marker, (d) fibronectin, basic cytokeratin, and PECAM, respectively, immunogenicity of HaCaT and HUVEC. fluorescence Images,
Figure 11 is a diagram showing the evaluation of the antibacterial activity of the developed hydrogel, (a) colony counting method after 24 hours of treatment and (b) optical density measurement method at different time intervals.
Figure 12 shows the evaluation of the wound healing potential of hydrogel scaffolds developed using 3-week-old ICR rats, showing (a) representative images of the used rats with regions of interest, (b) rats with implanted scaffolds; (c) Photograph of a wounded rat 7 and 14 days after transplantation, (d) Masson Trichrome staining of the remaining wound area after 7 days and 14 days of healing, (e) H&E, and (f) healed area 14 days after transplantation.

The present invention will be described in more detail below through non-limiting examples. However, the following examples are intended to illustrate the present invention, and the scope of the present invention is not to be construed as limited by the following examples.

In the present invention, high molecular weight chitosan powder, potassium hydroxide (all from Sigma Aldrich, USA), sodium chlorite (Daejoong Chemical), acetic acid, hydrochloric acid, sulfuric acid (Wako Chemicals, Korea) and sodium hydroxide (Junsei Chemicals, Japan) were used. All chemicals were used as received from the suppliers.

Example 1: Extraction of spherical nanocellulose (referred to as 's-NC') and synthesis of water-soluble carboxymethyl chitosan

s-NC extraction was performed from pine wood powder through chemical treatment, and the detailed method is as follows.

Briefly, pine wood powder (15 g) was added to a 3% (w/v) potassium hydroxide solution and heated at 90 °C for 1 h with continuous mechanical stirring to remove inorganic components from pine. The reaction medium was neutralized with 10% (v/v) hydrochloric acid solution, filtered, and washed with distilled water. The obtained material was dried at 50°C for 48 hours. Dried samples (12.9 g) were treated with sodium chlorite and acetic acid solutions at 85 °C with continuous mechanical stirring to remove lignin content from the samples. Afterwards, the reaction medium was filtered and washed with distilled water, and the collected samples were dried at 50°C for 48 hours. The dried sample (9.2 g) was further added to 17.5% (w/v) sodium hydroxide solution at room temperature and stirred for 1 hour to remove hemicellulose content. The reaction medium was neutralized with 10% (v/v) acetic acid solution, then filtered and freeze-dried using a freeze dryer (EYELA® Freeze Drying Unit 2200, Tokyo, Japan) for 36 hours. The obtained cellulose was 6.4 g.

As a follow-up operation to obtain s-NC,

The calculated amount of cellulose was treated with 1 M ammonium persulfate solution at 80 °C for 16 h. Afterwards, the reaction was quenched by adding 10 times the amount of ice-cold water, and dialyzed against distilled water using a 12-14 kDa molecular weight cellulose tube for 3 days. The dialyzed samples were lyophilized and lyophilized for further work.

Carboxymethyl chitosan (CMCS) was synthesized as follows. Briefly, a known amount of chitosan (CS) was added to a 50% sodium hydroxide solution and stored in the refrigerator for 24 h, followed by the addition of the required amount of isopropanol with continuous mechanical stirring. Afterwards, a calculated amount of monochloroacetic acid was added to the reaction mixture and stirred at 60°C for 6 hours, then filtered and washed with ethanol. The obtained sample was dissolved in distilled water and dialyzed for 3 days. The dialyzed samples were concentrated with a rotary evaporator and freeze-dried with a freeze dryer (EYELA® Freeze Drying Unit 2200, Tokyo, Japan).

Example 2: Synthesis of CMCS and composite hydrogel scaffold (CMCS/s-NC)

Hydrogels were prepared by dissolving the required amount of CMCS (5%, w/w) in distilled water. Composite hydrogels were prepared by adding different amounts of s-NC (1, 2, 4%, w/w, w.r.t. CMCS) to the CMCS matrix. The developed hydrogel was pre-cooled at -80 °C for 6 hours and then freeze-dried to prepare a porous scaffold. Pure polymer and composite hydrogel scaffolds were designed with CMCS and CMCS/s-NC-y, where y is the amount of s-NC in the CMCS matrix.

Example 3: Characterization of hydrogel scaffolds

Fourier transform infrared (FTIR) spectroscopy (Frontier, Perkin Elmer, UK) was used to monitor the interaction of the functional groups of s-NCs with the CMCS matrix at a resolution of 4 cm ^-1 in the wavelength range of 4000–400 cm ^-1 . The size and shape of s-NC were measured using a transmission electron microscope (TEM) (JEM, 2100 F, Jeol, Japan). The surface morphology of the developed hydrogel scaffold was examined by scanning electron microscopy (SEM) (S-4800, Tokyo, Japan). Structural changes in the scaffolds were evaluated with ^an The operating range (2θ) was 5-40 ^○ . The rheological properties of the developed hydrogels were evaluated with an ARES-G2 rheometer (TA Instrument, New Castle, Delaware, USA) using 6 mm parallel plates at 25 °C.

Example 4: Recovery, injectability, and adhesion of hydrogels

The recovery potential of the developed hydrogel was measured by measuring its viscosity via ARES-G2 rheometer at different shear rates (0.1, 100 and 0.1 s ^-1 ) for different time intervals (0-100, 101-200 and 201-300 s). It was evaluated by measurement. each at 25°C. The injectability properties were evaluated using a 10 mL syringe with a needle dimension of 21 G.

For adhesion measurements, 200 μL of the developed hydrogel was placed on the plastic surface to cover an area of 25 mm × 15 mm. A second plastic sheet was placed on top of the adhesive layer and left at room temperature for 15 minutes. Afterwards, uniaxial tensile measurements were performed to measure the adhesive strength of the hydrogel. All experiments were performed in triplicate (n = 3).

Example 5: Expansion force

The swelling potential of the hydrogel scaffold was measured in distilled water at room temperature. For this purpose, a known amount of scaffold was soaked in water and removed after a certain period of time. Excess water on the surface of the scaffold was removed with tissue paper, and the weight of the swollen scaffold was measured. Afterwards, the scaffold was again immersed in water and this process was repeated several times. The swelling potential of the scaffold was calculated using the following formula:

W _s and W _d are the weights of the expanded and dried scaffolds, respectively.

Example 6: Degradation The degradation efficiency of the hydrogel scaffold was analyzed with protease (1 mg/mL) enzyme at room temperature. For this purpose, the calculated weight of the scaffold was stored in the enzyme medium for a fixed time. Afterwards, the support was removed from the medium and rinsed with distilled water. After removing surface water with tissue paper, the support was dried. The dried support was weighed and this process was repeated several times. The old enzyme medium was replaced with new medium every 3 days. The degradation rate was calculated using the given formula,

W _i and W _d are the weights of the initial and decomposed scaffolds, respectively.

Example 7: Bioactivity Assay

The bioactivity of the hydrogel scaffolds was investigated in simulated body fluid (SBF) medium at room temperature. For this purpose, the fabricated scaffold was soaked in SBF solution for 14 days. During this process, the existing SBF medium was replaced with new medium after 3 days to eliminate fluctuations in cation concentration due to decomposition of the scaffold. After SBF treatment, the scaffolds were washed with water and freeze-dried before SEM examination. The medium composition of the SBF solution is shown in Table 1. The developed hydrogel scaffold was cross-linked with 5% CaCl ₂ solution before SBF treatment.

Table 1 shows the chemical composition of SBF medium.

Example 8: Conductivity and electrochemical analysis

The conductivity of the developed hydrogel was measured using a 4-probe instrument (MS Tech, Solution) connected to a Keithley 2460 source meter®. The electrochemical activity of the developed hydrogel was analyzed in the range of -0.6V to 1.2V using a Keithley 2460 source meter® at a scan rate of 25mV/s. The dimensions of the hydrogel used for electrochemical analysis were 2.2 × 0.8 × 0.4 cm ³ .

Example 9: Cell Viability

Cell viability of HDF, HaCaT and HUVEC individually or in co-culture was monitored by WST-1 assay using the developed hydrogel scaffold. For this purpose, 1 × 10 ⁴ (individual culture) and 0.3 × 10 ⁴ (co-culture) cells were cultured with or without support for 1, 3 and 5 days in a 5% CO2 incubator at 37°C. The group without scaffold treatment was considered the control group. After culturing, WST-1 dye was added to the culture medium and cultured for another 2 hours to form formazan. The amount of formazan was measured by measuring absorbance at 450 nm with a spectrophotometer (Infinite® M Nano 200 Pro, TECAN, Switzerland). All experiments were performed in triplicate (n = 3) and results are expressed as mean OD ± standard deviation (SD). Statistical significance was taken at *p < 0.05.

Example 10: Cell Morphology

The morphology of individual and co-cultured cells was visualized using an inverted fluorescence microscope (DMi8 Series, Leica Microsystems, Germany) after 3 days of treatment. For this purpose, 2.0 × 10 ⁴ (individual cells) and 0.5 × 10 ⁴ (co-culture cells) were cultured with or without supports. The group without scaffold was used as the control group. After treatment, cells were washed with PBS and fixed with 4% paraformaldehyde (PFA) (Sigma-Aldrich, USA) solution. Fixed cells were washed with PBS, treated with 0.1% Triton-X 100 for 10 minutes, and then treated with 1% Bovine Serum Albumin (BSA) (Sigma-Aldrich, USA) for 60 minutes. Afterwards, cells were stained with 200 μL of Alexa Flor 488-conjugated Phalloidin (F Actin Probe; Invitrogen, Thermo-Fischer Scientific, USA) for 20 minutes and then stained with 4,6-diamino-2-phenylindole dihydrochloride (DAPI) (Sigma-Aldrich, USA) for 5 minutes. Excess staining was removed by washing with PBS, and mounting was performed by adding 1 drop of Prolong® Antifade mounting media (Invitrogen, Thermo-Fischer Scientific, USA) and images were taken under a microscope.

Example 11: RNA extraction and real-time polymerase chain reaction (qPCR) analysis

qPCR technology was used to examine mRNA expression in HDFs, HaCaT, and HUVECs with or without scaffolds. The group that did not receive any treatment was used as the control group.

For this purpose, 1 × 10 ⁶ cells were cultured with DMEM medium in 24-well plates for 7 and 14 days, after which RNA extraction was performed using TRIzol® reagent according to the manufacturer's instructions. After monitoring the purity and concentration of the extracted RNA using a spectrometer, cDNA was synthesized using reverse transcriptase. qPCR was performed with SYBR Green Master mix with reaction conditions including 43 cycles of denaturation at 95°C for 15 s and amplification at 60°C for 1 min. All reactions were completed in triplicate (n = 3) and normalized to the glyceraldehyde 3-phosphate dehydrogenase (GAPDH) housekeeping gene. The expression of mRNA was measured using the Bio-Rad Real-Time PCR (CFX96TM Maestro Real-Time System, BioRad, USA) system. Relative mRNA expression levels in different skin cells with control and scaffold treated groups were compared in histograms. Primers used in this study are listed in Table 2.

Table 2 is a list of primer sequences used in the present invention.

Example 12: Immunocytochemical analysis of protein markers

Immunocytochemical analysis was performed to investigate the expression of various proteins, including fibronectin, basic cytokeratin, and platelet endothelial cell adhesion molecule-1 (PECAM-1 or CD 31) in HDF, HaCaT, and HUVEC, respectively.

For this purpose, 2 × 10 ⁴ cells (individual) were placed in a 24-well plate and cultured for 14 days with or without support. The group without scaffold treatment was considered the control group. Afterwards, the cells were washed with PBS and fixed with 4% PFA solution for 20 minutes at room temperature. Cells were permeabilized for 15 minutes at room temperature using 0.1% Triton Finally, cells were treated with 250 μL of mouse monoclonal antibodies against fibronectin, basic cytokeratin, and PECAM-1, and then nuclei were counterstained with 20 μL of DAPI solution for 2 minutes. Excess staining was removed by washing with PBS and mounting was performed by adding mounting media. Morphology was captured by fluorescence microscopy.

Example 13: Antibacterial activity of hydrogel support

The antibacterial activity of the hydrogel scaffold was analyzed by measuring the colony count and optical density (OD) of Bacillus subtilis (ATCC-6051). For the colony counting method, 100 μL of bacterial solution ( ¹ added. Incubated at 37°C for 24 hours. The group without hydrogel was considered the control group. Afterwards, 100 μL of the bacterial solution was transferred to an agar plate and further cultured for 24 hours to form colonies.

For OD measurements, prepared scaffolds were incorporated into bacterial medium (1 × 10 ⁶ ) and treated for 6, 24, and 48 h. Media without scaffold was used as a control. Changes in OD were measured at 600 nm using a spectrophotometer. The test was performed three times.

Example 14: In vivo wound healing assay

In vivo studies were performed to monitor the wound healing efficiency of the developed hydrogel scaffolds using 3-week-old ICR male rats (N = 6) 7 and 14 days after implantation. The experiment was divided into an untreated negative control group, a positive control group - pure polymer support (CMCS), and an experimental group - CMCS/s-NC 1.

Here, CMCS/s-NC 1 was selected as the experimental group because it showed better skin proliferation compared to other scaffolds in vitro. Each group has duplicate rats (n = 2) with two experimental sites. Briefly, the back fur of the mice was cleaned and the skin was disinfected with a 70% ethanol solution. Afterwards, a circular wound with a diameter of 1 cm was created using a sterile biopsy punch. Two layers of dorsal skin were removed and filled with scaffold. All rats were housed in a protected, soundproof room. Room temperature and relative humidity (RH) were maintained at 21 ± 1°C and 35 ± 1%, respectively. The room was maintained on an automatically controlled 12-hour light and 12-hour dark cycle. Every possible effort was made to minimize animal suffering and the number of animals needed to produce consistent scientific results. Rats were sacrificed 14 days after treatment to perform histological staining.

The specimens collected for histological analysis were fixed in 3.7% PFA solution for 2 days and then decalcified in 12% ethylene diaminetetraacetic acid (EDTA) at 4°C. The specimens were then washed and dehydrated with alcohol. Dehydrated specimens were implanted in paraffin to produce microsections. Specimen sections were stained with hematoxylin, eosin, and Masson's trichrome. All surgical protocols were approved by the Capital Medical University Animal Experiment Ethics Committee (AEEC, Institutional Animal Care and Use Committee of Capital Medical University, permit number CMUSH-IRB-KJ-YJ-2020-29). Scaffolds were sterilized by UV light treatment before implantation.

Statistical analysis of the present invention was performed by one-way analysis of variance using Origin Pro9.0 software. All results were expressed as mean ± standard deviation (SD). Statistical significance was considered at *p<0.05, **p<0.01, ***p<0.001. All comparisons were made between control and treatment groups.

The results of the above examples are described below.

Characterization of s-NCs

A schematic representation of the fabrication of the hydrogel scaffold and its potential application in wound healing is presented in Figure 1.

Fourier transform infrared spectroscopy (FTIR) spectra of cellulose and s-NCs are shown in Figure 2a. The characteristic peaks at 3325, 1640, and 1019 cm-1 are attributed to hydroxyl (-OH) groups, absorbed water, and -CO stretching groups, respectively. Cellulose and s-NC did not show a significant peak at 1500 cm ^-1 , suggesting that the lignin portion was removed from the structure by chemical treatment. The appearance of an additional absorption peak at 1731 cm ^-1 in s-NC is attributed to the carboxylation of cellulose functional groups by ammonium persulfate (APS). No other significant changes were observed in the s-NC spectra, suggesting that APS hydrolysis did not alter the structural integrity of pure cellulose. The morphology of the obtained s-NC was examined by AFM, and the image is shown in Figure 2b. The obtained s-NCs exhibited particle-like morphology with an average diameter of ~65 nm, indicating successful formation of nanoparticles by APS hydrolysis. The XRD patterns of cellulose and s-NC are shown in Figure 2c. The diffraction peaks at 11.8°, 20.0° and 21.5° are attributed to cellulose I and II structures. s-NCs exhibit sharper and more intense diffraction peaks than cellulose, indicating improved crystallinity of s-NCs due to the removal of amorphous regions.

Structural and morphological characterization of hydrogel scaffolds

The FTIR spectra of CMCS and its composite scaffold are shown in Figure 3a. The FTIR spectrum of pure CS is shown in the inset of Figure 3a. The broad characteristic peak at 3324 cm ^-1 in pure CS can be attributed to the hydrogen-bonded amine (-NH ₂ ) functional groups, which shifted to a lower wave number (3329 → 3303 cm -1) in the CMCS hydrogel (Figure 4). . In CMCS, the absorption peak at 1639 cm ^-1 is assigned to the carboxyl group overlapping with the amino group. The absorption peak at 1156 cm ^-1 is weakened in CMCS, indicating that carboxymethylation occurs at the hydroxyl group of CS. Additionally, a shift of the absorption peak (1416 → 1409 cm ^-1 ) was observed in the composite hydrogel, indicating the interaction between the integrated s-NC and CMCS. This movement of the composite hydrogel may be due to minimizing the intramolecular hydrogen bonding interactions between the polymer chains and increasing the intermolecular hydrogen bonding interactions between the integrated s-NCs and the polymer chains. These results suggest that the developed hydrogels are highly interactive and that hydrogen bonding is likely to play an important role in these changes.

Structural changes in the developed hydrogel were evaluated using an X-ray diffractometer, and the XRD pattern is shown in Figure 3b. Broad and sharp diffraction peaks at 8.9° and 20.5° are observed in CMCS, showing amorphous and crystalline regions in the structure, respectively. The composite hydrogel exhibits a diffraction pattern similar to that of the pure polymer with slight changes in peak positions. A decrease in the interplanar distance (1 → 0.93 nm) of the amorphous region was observed in the composite hydrogel, indicating the formation of an assembled structure due to the interaction between the integrated s-NCs and the polymer chains, as observed in the FTIR results. . No significant change in the crystalline interplanar distance was observed in the composite hydrogel (0.44 → 0.45 nm), suggesting that the distribution of s-NCs mainly occurred in the amorphous region of the polymer.

The surface morphology of the hydrogel scaffolds was analyzed using SEM and the images are shown in Figure 3c. The fabricated scaffolds exhibit high porosity and an interconnected morphological structure. A hierarchical structure was observed in the scaffold. The composite scaffold showed a more porous morphology than the pure polymer scaffold.

The swelling efficiency of scaffolds plays an important role in tissue engineering applications because it promotes the exchange of cellular metabolites and nutrients. The swelling potential of the developed hydrogel scaffold was investigated in aqueous conditions at room temperature, and the results are shown in Figure 3d. The composite hydrogel scaffolds had lower swelling efficiency than the pure polymer scaffolds, and this potential further decreased with increasing s-NC content. The decrease in swelling efficiency can be explained by more pronounced interactions between s-NCs and polymer chains, which results in lower availability of functional groups for water absorption. The expansion of the scaffold must be controlled under physiological conditions. Rapid expansion can weaken and degrade the scaffold. Additionally, no improvement in swelling efficiency was observed after 15 hours, indicating the saturation point of the fabricated hydrogel scaffold.

Rheological behavior and recovery analysis

Hydrogels developed for tissue engineering applications must have excellent elasticity and mechanical strength. The mechanical strength of the sample prepared under hydrogel conditions was analyzed using a rheometer, and the results are shown in Figure 5. The changes in storage modulus (G', with lines) and loss modulus (G", without lines) in the measured frequency range are shown in Figure 5a. The composite hydrogels showed higher G values than the pure polymer hydrogels; It showed high elasticity due to the large interaction between s-NC and the polymer chain.

A significant improvement in G' value (1.11 × 10 ⁵ Pa → 5.64 × 10 ⁶ Pa) was observed in the composite hydrogel due to the greater interaction between the polymer chains and s-NCs, which creates an interconnected network structure with high It leads to elasticity. The enhancement of the G' value in the high ω region of the composite hydrogel can be explained by the relaxation and formation of a more interconnected network structure, resulting in a solid-like structure and higher G' value.

Higher G" values were observed for composite hydrogels than for pure polymer hydrogels, and the values further increased with increasing s-NC content. However, the magnitude of G" was lower than G' throughout the measured ω region. , which suggests the formation of a stable elastic polymer network structure in the hydrogel.

The change in complex viscosity (η*) of the hydrogel developed in the ω region was also measured, and the results are shown in Figure 5b. In the measured area, higher η* values were observed in the composite hydrogel than in the pure polymer hydrogel, and further increased with increasing s-NC content, indicating more solid-like properties in the composite hydrogel. The η* value increases in the low ω region, while the decrease in η* value occurs in the high ω region, suggesting shear thickening and shear thinning characteristics, respectively. This is an expected property of a printing application.

The printing power was confirmed by analyzing the thixotropic behavior of the developed hydrogel. For this purpose, a series of shear rates of 0.1, 1000, and 0.1 s ^-1 (100 s each) were applied to monitor the changes in the η* value of the hydrogel. A high shear rate (1000 s ^-1 ) was used to deform the structure similar to the printing conditions. Printable hydrogels should exhibit high thixotropic behavior. The variation of η* values at different shear rates is shown in Figure 5c. The initial η* values of CMCS, CMCS/s-NC 1, CMCS/s-NC 2, and CMCS/s-NC 4 were 1314, 4335, 695, and 562 Pa, respectively 0.552 at high shear rate (1000 s ^-1 ). , decreased to 0.326, 0.443, and 0.687 Pa. This value further increased at low shear rates (0.1 s ^-1 ) and was 681, 2787, 484, and 475 Pa for CMCS, CMCS/s-NC 1, CMCS/s-NC 2, and CMCS/s-NC 4, respectively. . Resilience was 51.7, 64.2, 69.2, and 84.2% for CMCS, CMCS/s-NC 1, CMCS/s-NC 2, and CMCS/s-NC 4. The high recovery potential of the composite hydrogel is attributed to the formation of an interconnected polymer network structure with integrated s-NCs, which limits the deformation of the hydrogel.

The injection force of the developed hydrogel was tested using a syringe at room temperature, and the digital image of the injected hydrogel is shown in Figure 5d. The hydrogel was easily injected and maintained its structure. Injectable hydrogels have various advantages and can be injected into desired areas. Injectable hydrogels are considered attractive materials for skin wound regeneration because they can reduce the likelihood of gel fragmentation at the implanted site and promote sustained skin wound healing even after external mechanical deformation. It can also be applied to irregular wounds.

Adhesion of developed hydrogel

The adhesive properties of the developed hydrogel were investigated on various surfaces, and the image is shown in Figure 6a. The developed hydrogel immediately adheres to other surfaces and shows adhesive strength. The adhesive properties of the developed hydrogel can be attributed to multiple interactions between the functional groups of the hydrogel and the material surface. Adhesion is highly desirable for a variety of applications including stopping bleeding, sealing leaks, bonding tissue, and improving tissue recovery. Additionally, a lap shear test was performed to quantify the adhesion strength between the hydrophobic surface (plastic) and the developed hydrogel after 15 min of treatment by uniaxial tensile measurements. A schematic diagram of the lap shear test is shown in Figure 6b.

The developed hydrogel was sandwiched between two plastic substrates (25 mm × 15 mm) and measured. The obtained curves of adhesion strength versus displacement for plastics and different hydrogels are shown in Figure 6c. An improvement in adhesion was observed in composite hydrogels compared to pure polymer hydrogels, which is indicative of adhesion. This force was further enhanced by increasing the s-NC content in the polymer matrix, indicating improved adhesion properties. The enhanced adhesion is due to the stronger interaction between the surfaces. Various factors, including adhesion time, properties of the hydrogel, and applied substrate, widely influenced the adhesion properties.

The quantitative values of the adhesion between the developed plastic substrate and the hydrogel are shown in Figure 6d. The adhesive strengths were 5.53, 11.84, 15.13, and 18.32 kPa for CMCS, CMCS/s-NC 1, CMCS/s-NC 2, and CMCS/s-NC 4, respectively. The developed hydrogel is expected to be able to further increase adhesion on hydrophilic surfaces, including pig and human skin, due to the strong interaction between the polar group on the tissue surface and the hydrogel, and can also be applied to skin wound healing.

The possible mechanism for the adhesion of the developed hydrogel to various substrates is shown in Figure 6e. Here we describe the mechanism of interaction between glass and the human skin surface by the hydrogel surface. Hydrogels contain charged ions (NH4 ⁺ and COO ^- ) that can interact with other polar or charged functional groups through dipole-dipole or ion-dipole interactions. The glass has negatively charged silicate groups, which interact with the positively charged amino groups of the hydrogel through the interactions mentioned above. Similarly, the skin surface can interact with the hydrogel.

Degradation and bioactivity potential of hydrogel scaffolds

To monitor the degradation rate of the developed hydrogel scaffold, an in vitro degradation study was performed and the results are presented in Figure 7a. The composite scaffolds showed a greater degradation rate than the pure polymer scaffolds in the presence of proteases. This behavior can be explained by the morphological aspects of the hydrogel scaffold. The pure polymer scaffold has a relatively closed structure compared to the composite scaffold observed in SEM images, which limits the protease medium to a greater extent and causes slow degradation. Scaffolds must have an appropriate degradation rate for tissue engineering applications. High rates of decomposition are unsuitable for cellular activity, while slow decomposition may cause an inflammatory response in surrounding tissues. Decomposition of polymeric materials is greatly influenced by crystallinity, molecular weight, pH of the medium and the presence of enzymes. Biodegradation mechanisms depend on several factors, including erosion, hydrolysis, and solubilization of the material.

The bioactivity potential of the developed hydrogel scaffold was analyzed using SEM, and the results are shown in Figure 7b. The biological activity of these substances plays an important role in cell proliferation and differentiation. Higher mineralization was observed on the surface of composite scaffolds than pure polymer scaffolds, indicating improved bioactivity. The bioactive potential of the developed scaffolds was significantly influenced by the s-NC content. Biodegradable polymers with bioactive glass or bioceramics showed more significant bioactivity than the polymer scaffold alone. It was observed that pure chitosan has a low biomineralization potential and is not suitable for cell differentiation. The incorporation of nanoparticles induces the potential for biomineralization by forming an apatite layer on the support surface. The EDX spectrum of the deposited minerals is shown in the inset of the SEM image. The appearance of calcium and phosphorus peaks indicates the formation of an apatite layer on the scaffold surface. The deposition rates of the minerals formed are listed in Table 3. It was confirmed that the composite scaffold had a higher ratio of calcium and phosphorus than the pure polymer scaffold, thereby improving bioactivity.

Table 3 shows elemental analysis of minerals deposited by EDX on the hydrogel scaffold developed after 14 days of SBF medium treatment.

Conductivity and electrochemical analysis of hydrogels

The conductivity of the developed hydrogel was measured using the four-probe technique, and the results are shown in Table 4. Comparison of pure polymer hydrogels and composite hydrogels. This improvement in conductivity is due to more protonation of the amino functional groups of CMCS in the presence of s-NCs, which promotes improved ionic conductivity. The developed hydrogel exhibits ionic conductivity within the range of skin conductivity (2.6 -1 x 10 ^-4 mS/cm). Therefore, the developed hydrogel, which has conductivity and excellent adhesion, is expected to be able to significantly stimulate skin cell activity after transplantation and aid in the rapid wound healing process. Functionalization of chitosan not only improves water solubility but also improves ionic conductivity.

Sample Dimension (cm) Resistance (R) (ohms) Conductivity =L/RA (S/cm-1) CMCS 2.4 Х 0.9 Х 0.5 1629.73 6.817 Х10 ^-5 CMCS/s-NC 1 2.3 Х 1.0 Х 0.3 1599.61 6.251 Х10 ^-4 CMCS/s-NC 2 2.4 Х 0.9 Х 0.6 1572.22 7.067 Х10 ^-4 CMCS/s-NC 4 2.2 Х 0.6 Х 0.4 864.39 2.892 Х10 ^-3

Table 4 shows the conductivity analysis of the hydrogel developed using the four probe system.

The electrochemical behavior of the developed hydrogel was evaluated by cyclic voltammetry using a three-electrode system, and the results are shown in Figure 8. The pure polymer hydrogel exhibits oxidation and reduction peaks at 0.41 V and -0.015 V, which represent the protonation and deprotonation of the -NH and -COOH functional groups of the CMCS polymer, respectively. A systematic shift of the oxidation (0.41 → 0.77 V) and reduction (-0.015 → 0.051 V) peaks was observed in the composite hydrogel, suggesting that the addition of s-NCs to the polymer matrix promotes the electrochemical activity of the hydrogel. . The composite hydrogel exhibited a larger anodic current (0.41→0.76 μA) compared to the pure polymer hydrogel, demonstrating improved electrochemical performance in the presence of s-NCs. The improved electrochemical performance of the composite hydrogel can be attributed to the bridging effect of s-NC functional groups (-COOH and -OH), which promotes electron transfer. Additionally, a distinct box-shaped hysteresis was observed in the developed hydrogel, suggesting electro-oxidation of the hydrogel at the electrode surface. The hysteresis area further increases in the composite hydrogel, showing greater electro-oxidation of the hydrogel at the electrode surface.

Biocompatibility and cell morphology

The biocompatibility of the produced hydrogel scaffold was evaluated using skin cells (HDF, HaCaT cells, HUVEC) using the WST-1 assay, and the results are shown in Figure 9. The present inventors investigated the effect of the developed hydrogel scaffold on skin cells by culturing individual cells in the presence of a hydrogel scaffold and then co-culturing them. HDF viability in the presence of hydrogel scaffolds after 1, 3, and 5 days of treatment is shown in Figure 9A. The untreated group was considered the control group. No side effects were observed in the presence of the hydrogel scaffold, which showed biocompatibility. The synthetic scaffolds showed greater cell viability than the control scaffolds. The group containing the CMCS/s-NC 1 hydrogel scaffold showed improved viability compared to other scaffolds, indicating improved biocompatibility. Fibroblasts play an essential role in skin physiology and skin wound healing. Produces all ECM components such as adhesive and structural proteins, proteoglycans, and glycosaminoglycans.

The cytotoxicity of the hydrogel scaffold was analyzed using HaCaT cells, and the results are shown in Figure 9b. Medium without scaffold treatment was used as a control. The hydrogel scaffold has no toxic effects on HaCaT cells, indicating biocompatibility. The composite scaffold showed higher cell viability than the pure polymer scaffold and the control group. The improved biocompatibility of composite scaffolds is attributed to their open morphological structure and bioactive potential. HaCaT cells act as a barrier against foreign substances, reducing water and heat loss from the body.

The biocompatibility of the hydrogel scaffolds was monitored with HUVECs after different time intervals and the results are shown in Figure 9c. The untreated group was considered the control group. No cytotoxicity was observed in the scaffold treatment group, indicating biocompatibility. Higher cell viability was observed for composite scaffolds than pure polymer scaffolds and control scaffolds, suggesting improved biocompatibility. HUVECs play an important role in angiogenesis and blood vessel formation. Newly created tissue requires an active vascular network for survival.

Additionally, HDF, HaCaT cells, and HUVEC were co-cultured with the developed scaffold to generate a skin regeneration model, and the results are shown in Figure 9D. It was interesting to see that the developed scaffold promoted skin cell proliferation in all treatment groups except CMCS/s-NC 4. NC 1 treated medium shows that 1% s-NC content is optimal for improved cell activity. The improved cellular activity of the developed hydrogel scaffolds is due to its favorable topographic structure, adhesive properties, and improved bioactivity. Skin cells (containing peptide layer) interact strongly with the scaffold surface, and the open structure facilitates gas and nutrient exchange. However, a decrease in skin cell survival occurred in the CMCS/s-NC 4 treatment group compared to other treatment groups. This may be due to the higher concentration, properties, and shape of nanoparticles.

The morphology of the co-cultured cells was observed using a fluorescence microscope 3 days after treatment, and the image is shown in Figure 9e. We selected CMCS/s-NC 1 scaffold for the treatment group because of its better cellular and bioactivity potential. Scaffold-free medium was used as a control. The cells were healthy and spread over the scaffold. Cell density was higher in the scaffold treated group than in the control group, confirming excellent biocompatibility as observed in cell viability.

Expression of mRNA and protein markers

The expression of HDF-related gene markers (fibronectin and collagen 1A) in HDF was evaluated by q-PCR technique after 7 and 14 days of treatment, and the results are shown in Figure 10a. The group without scaffold treatment served as the control group. Higher expression of fibronectin and collagen 1A was observed in the scaffold-treated group compared to the control group after 7 days of treatment, which further increased after 14 days of culture, showing an accelerated effect on HDF-related gene markers. This was higher for the CMCS/s-NC1 scaffold compared to the pure polymer scaffold. The better expression of HDF-related gene markers in CMCS/s-NC1 is due to better cell activity and favorable topographic structure. Fibronectin is considered an important component of the extracellular matrix (ECM) and is required for cell adhesion to the matrix. It plays an important role in cytoskeletal modification, cell migration, and differentiation. It also accelerates embryonic development and wound healing. Collagen is the major fibrous glycoprotein of the ECM and promotes ECM assembly, ECM-cell interactions, cell migration and differentiation.

The expression of different keratinocyte-related gene markers (KRT1, KRT5, KRT10, KRT15) was also investigated in HaCaT, and the results are shown in Figure 10b. Media without scaffold treatment is used as a control. Higher expression of keratinocyte-related gene markers occurred on the scaffold compared to the control group, demonstrating keratinocyte production efficiency. KRT1/10 is considered an early genetic marker and its expression increased after 7 days of treatment and decreased after 14 days of culture. KRT5/KRT15 are important keratinocyte gene markers, and their expression occurs in late stages of differentiation. Better expression of keratinocyte-related gene makers showed that the developed scaffold could accelerate keratinocyte formation, which plays an important role in wound repair. HUVEC-related gene markers (VEGF, VE-CAD, and CD31) were further examined after 7 and 14 days of treatment, and the results are shown in Figure 10c. The group without scaffold was considered the control group. Greater expression of HUVEC-related gene markers occurred in the scaffold-treated group than in the control group, suggesting the angiogenic potential of the scaffold, which was further enhanced after 14 days of incubation. The CMCS/s-NC 1 treatment group demonstrated improved expression over the CMCS scaffold due to improved cellular activity. VEGF, VE-CAD, and CD31 play important roles in angiogenesis, maintenance, and vascular structure organization. These results indicate that the developed scaffold is efficient and can promote angiogenesis.

We further examined the expression of fibronectin, basic cytokeratin, and PECAM-1 proteins in HDF, HaCaT, and HUVEC using fluorescence microscopy after 14 days of culture, and the results are presented in Figure 10d. The group without any treatment was considered the control group. The scaffold-treated group expressed greater intensity of this protein than the control group, suggesting that the developed scaffold positively induces skin cell function. Expression of fibronectin protein occurs throughout the cytoskeleton of HDFs, whereas cytokeratin and PECAM-1 are more expressed near the nucleus.

Antibacterial power of hydrogel

The antibacterial activity of the developed hydrogel was investigated using colony counting method 24 hours after treatment with Bacillus subtilis, and the image is shown in Figure 11a. The group without hydrogel treatment was used as a control group. Compared to the control group, a significant decrease in bacterial colonies was observed in the hydrogel-treated group, indicating antibacterial activity, and no colonies were observed in the composite hydrogel-treated plate, indicating improved antibacterial activity. The antibacterial potential of the developed hydrogel is due to the presence of positively charged functional groups that electrostatically interact with and damage bacterial cell membranes. The improved antibacterial potential of composite hydrogels may be due to the wrapping of positively charged polymer chains with s-NCs, resulting in greater exposure to bacteria, resulting in more severe electrostatic interactions and consequent bacterial death.

These findings were further confirmed by measuring the optical density of the cultured bacterial solution at time intervals, and the results are shown in Figure 11b. The untreated group was considered the control group. After 6 hours of treatment, a significant decrease in OD value was observed in the hydrogel-treated group compared to the control group, indicating antibacterial activity. It further decreased with incubation time (24 hours and 48 hours), suggesting that antibacterial activity was improved. Hydrogel. The composite hydrogels have substantially lower OD values than pure polymer hydrogels, indicating improved antibacterial potential, consistent with the results of the colony counting method.

In vivo wound healing analysis

Wound healing is a complex biological process involving various steps including hemostasis, inflammation, migration, proliferation and remodeling. We evaluated the subcutaneous wound healing potential of the developed hydrogel scaffold using ICR rats based on in vitro assays. A representative image of a used rat with the desired area for implantation is shown in Figure 12A. Create a 1 cm spherical area for the wound healing process.

A representative image of the wound area where the scaffold was implanted is shown in Figure 12B. To minimize the influence of other factors such as physiological changes and number of experimental animals, two rats with two wound sites were used in each group.

Images of mice used after 7 and 14 days of treatment are shown in Figure 12C. The group implanted with the hydrogel scaffold showed more wound closure potential than the control group after 7 days of treatment, which was further enhanced after 14 days of treatment, demonstrating wound healing efficacy. Among them, mice treated with CSCM/s-NC 1 hydrogel scaffold showed greater wound closure potential compared to other mice. The improved skin regenerative efficacy of the CSCM/s-NC 1 hydrogel scaffold was attributed to better cellular activity that promoted skin cell proliferation and enhanced gene and protein expression, as observed in in vitro results.

The quantitative results of the remaining wound area after 7 and 14 days of treatment are shown in Figure 12d. The results indicate that CSCM/s-NC 1 hydrogel scaffold has better wound closure efficacy than others. The remaining wound areas were 0.05 ± 0.003, 0.03 ± 0.001, and 0.02 ± 0.001 mm ² for control, CMCS, and CMCS/s-NC 1 scaffolds, respectively, 14 days after implantation. Almost complete wound healing was observed with the CSCM/s-NC 1 hydrogel scaffold 14 days after implantation, and the wound surface was covered with newly generated epidermis.

Hematoxylin and Eosin (H&E) analysis was performed to examine newly regenerated tissue with and without scaffolds after 14 days of treatment, and the results are shown in Figure 12e. Formation of a skin layer containing blood vessels and hair follicles occurred in all groups, but its density was higher in the scaffold-treated group than in the control group. This was significantly higher in the CSCM/s-NC 1 hydrogel scaffold treatment group, showing improved regeneration potential. The higher regenerative capacity of CSCM/s-NC 1 hydrogel scaffolds is assigned to the greater expression of various factors, including VEGF, which plays an important role in angiogenesis. Enhanced expression of VEGF triggers the formation of new blood vessels and aids re-epithelialization and collagen synthesis. Fibroblasts have abundant granulation tissue and growth factors. The wound healing process can be monitored by observing the thickness of granulation tissue. The thickness of granulation tissue was 83.91, 171.31, and 225.5 μm for control, CMCS, and CMCS/s-NC 1 scaffolds, respectively. Therefore, the CMCS/s-NC 1 scaffold showed improved wound healing efficacy compared to the control group, and the CMCS scaffold showed improved wound healing efficacy by supporting the thickness of granulation tissue. Additionally, collagen deposition was examined using Masson trichrome staining 14 days after treatment. Collagen deposition occurred in all groups. However, the density was higher in the scaffold treatment group, which was significantly higher in the CMCS/s-NC 1 scaffold treatment group, suggesting better regeneration potential. A fast and effective wound healing process will significantly lower the cost of treatment and wound healing aids. These results showed that application of the developed material led to rapid skin wound healing.

Claims (12)

delete delete delete delete delete delete delete An antibacterial composition having antibacterial activity against Bacillus subtilis, comprising as an active ingredient a hydrogel prepared by adding spherical nanocellulose to a caroboxymethylchitosan matrix. The antibacterial composition according to claim 8, wherein the hydrogel is prepared by adding 1 to 4% (w/w) of spherical nanocellulose to the caroboxylmethylchitosan matrix. delete delete delete