TW201341001A - Biomaterial for wound healing and preparation thereof - Google Patents

Abstract

The present invention provides a biomaterial comprising a scaffold consisting of collagen, hyaluronic acid, and gelatin, which are cross-linked via ethyl-3-[3-dimethylaminopropyl] carbodiimide (EDC) between any two of collagen, hyaluronic acid, and gelatin. The present invention further provides a method for preparing the biomaterial and a method for enhancing wound healing with the biomaterial.

Description

Biomedical material for wound healing and method of manufacturing same

The invention relates to a biomedical material comprising a scaffold composed of collagen, hyaluronic acid and gelatin, which is composed of ethyl-3-[3-dimethylaminopropyl]carbodiimide (ethyl-3- [3-dimethylaminopropyl]carbodiimide (EDC) cross-links two components of collagen, hyaluronic acid and gelatin. The invention further relates to a method for preparing such a biomedical material and a method for enhancing wound healing using the biomedical material.

The largest organ in the vertebrate - the skin is composed of the epidermis, the complex vascular and nerve-distributed dermis, and the subcutaneous tissue composed of lipid and loose connective tissue. These three layers play an important role in protecting the body from chemical or mechanical damage (Choi et al., J Cell Sci 123, 3102-3111 (2010)). A burn wounded patient with a large amount of dermal tissue is healed by wound contracture and scar tissue. Many experimental studies have sought new ways to use modern physical and pharmacological methods or phytotherapy to improve human skin cell growth (Dainiak et al., Biomaterials 31, 67-76 (2010)). Skin substitutes are not capable of skin rejuvenation due to antigenicity or donor site limitations and have not been widely used (Bell et al., Science 211, 1052-1054 (1981); Schulz et al., Annu Rev Med 51, 231-). 244 (2000); Boyce, Burns 27, 523-533 (2001); Ma et al., Biomaterials 24, 4833-4841 (2003)). At present, the promotion of skin cell growth is an expensive procedure in all ages in the world, and has been issued in recent years. Different methods were developed for in vitro reconstruction of human skin syntheses, including dermal (or dermal mimic) epidermal junctions (Kim et al., Br J Plast Surg 52, 573-578 (1999); Kremer et Al., Br J Plast Surg 53, 459-465 (2000); Hoeller et al., Exp Dermatol 10, 264-271 (2001); Souto et al., Sao Paulo Med J 124, 71-76 (2006)), Construction is an important factor in skin tissue engineering.

Collagen is an important component and is a major part of human connective tissue, especially in soft tissues of the skin (Duan and Sheardown, J Biomed Mater Res A 75, 510-518 (2005)). Over the past few decades, collagen porous scaffolds have been widely used in tissue engineering, such as skin, cartilage, bones and nerves, as a basis for support and cell infiltration, proliferation and differentiation. However, the problem of poor mechanical strength and rapid biodegradation of untreated collagen scaffolds is the key to limiting their application. In addition to mechanical force or uncontrollable, after the thermal or biochemical treatment, the triple helix structure of collagen is easily deformed into a random coil structure. Crosslinking a collagen-based scaffold is an effective way to enhance its mechanical properties and adjust its rate of biodegradation.

Hyaluronic acid (HA) is also one of the main components of the skin and is associated with tissue repair. Hyaluronic acid is a biopolymer having repeating structural units having a molecular weight greater than 1.0 M kDa and containing more than 3,000 disaccharides. The polymer consists of repeating structural units formed by alternating residues of D-glucuronic acid (GlcUA) and N-acetyl-D-glucosamine (GlcNAc) linked by (1→3) bonds. In skin tissue, this property is critical for retaining moisture (Toole, Curr Opin Cell Biol 2, 839-844 (1990)). HA is one of the most hygroscopic molecules in nature. HA is widely distributed in the epithelium, nerves and connective tissues, and is a cell One of the important components of the extracellular matrix (ECM), which supplies cell differentiation and proliferation.

In addition, gelatin, denatured collagen allows parts of the cell to adhere to, grow and differentiate to maintain its natural form (Lee et al., Biomaterials 24, 2503-2511 (2003)). Furthermore, gelatin has functional groups such as an amine group and a carboxyl group, which is advantageous for modifying surface properties (Usta et al., Biomaterials 24, 165-172 (2003)) and contains a large amount of glycine, proline and 4-hydroxyguanamine. Acid residue. The cross-linking reaction can be used to adjust the rate of degradation and biomechanical properties (usually consistent with the properties of the tissue used for regeneration), but may reduce biocompatibility (Zeeman et al., Biomaterials 20, 921-931 (1999)).

Therefore, the cross-linking treatment of collagen/HA/gelatin scaffolds is one of the important considerations for bioporous scaffolds. There are currently two types of crosslinking methods commonly used to improve mechanical properties: physical processing and chemical techniques (Li et al., J Mater Sci Mater Med 21, 741-751 (2010)). The former includes the use of photooxidation, dehydrothermal and UV irradiation to avoid the introduction of potentially cytotoxic chemical residues and maintain good biocompatibility of collagen materials (Lee et al., Yonsei Med J 42,172- 179 (2001)), but most physical treatments do not achieve a sufficiently high degree of cross-linking to meet demand, so the use of chemical treatments is still necessary in some cases. When selecting a crosslinker for collagen/HA/gelatin porous scaffolds, heterobifunctional agents (which contain 2 different reactive groups and can directly link 2 different amino acid side chains) The degree of crosslinking can be maximized (Pieper et al., Biomaterials 21, 581-593 (2000)).

US Patent Application No. US 2004/0267362 discloses a connective tissue scaffold comprising anchored segments, bioresorbable polymeric fibers and a central segment, which may be collagen, hyaluronic acid, Made of gelatin, etc. The connective tissue scaffold facilitates replacement or enhancement of injured or torn connective tissue as an implant. However, the requirements for both biomedical materials for implanting or replacing connective tissue and biomedical materials for skin are quite different. Therefore, there is still a need for biomedical materials for dermal engineering applications.

In tissue engineering, the main goal of cell biologists is to use sophisticated culture conditions to promote cell growth. The performance of a skin mimetic depends on the cell growth after its application. Recently, there has been a lot of evidence that keratinocytes (KCs), melanocytes (MCs) and fibroblasts (FBs) in the skin affect each other's cellular functions and participate in the regulation between cells (Regnier et Al., J Invest Dermatol 109, 510-512 (1997); Berking et al., Am J Pathol 158, 943-953 (2001); Schneider et al., PLoS One 3, e1410 (2008)). For example, melanocytes are affected by external factors such as UV radiation and are also affected by intrinsic factors secreted by fibroblasts and keratinocytes (Yamaguchi et al., FASEB J 20, 1486-1488 (2006)); Certain keratinocyte-derived factors promote dendricity of isolated melanocytes (Gordon et al., J Invest Dermatol 92, 565-572 (1989)). An interesting study found that the interaction between melanocytes and the cell membrane of keratinocytes induces a transient intracellular calcium signal in keratinocytes, necessary for melanin transfer (Seiberg et al., Exp Cell Res 254, 25- 32 (2000); Joshi et al., Pigment Cell Res 20, 380-384 (2007); Yamaguchi et al., J Biol Chem 282, 27557-27561 (2007)); studies have also shown that fibroblasts can enhance keratin by recombinant collagen matrix or by the production of specific growth factors. Cell growth, growth factors produced by fibroblasts play an important role in the migration, proliferation and differentiation of keratinocytes and melanocytes (Wang et al., J Biomed Mater Res B Appl Biomater 82, 390-399 (2007)). 3-dimensional (3D) skin scaffolds provide ECM analogs as a basis for physical support and host infiltration to guide cell proliferation and migration to target functional tissues or organs (Park et al., Biomaterials 23, 1205-1212 (2002). ); Park et al., Toxicology 267, 178-181 (2010)). Ideal skin stents should have good biocompatibility, proper microstructure (eg, average pore size of 100-200 μm and porosity greater than 90%) for cell growth, controlled biodegradability, and suitable mechanical properties. . (Ma et al., Biomaterials 24, 4833-4841 (2003)). Therefore, proper skin materials still have their needs, and there is a need for better biomedical materials as bio-supporters to meet the requirements and co-culture with keratinocytes, melanocytes and fibroblasts to mimic normal human skin.

The invention provides a biomedical material comprising a scaffold consisting of collagen, hyaluronic acid and gelatin, which is composed of ethyl-3-[3-dimethylaminopropyl]carbodiimide (ethyl-3- [3-dimethylaminopropyl]carbodiimide (EDC) cross-links two components of collagen, hyaluronic acid and gelatin. EDC is a heterobifunctional, lengthless cross-linking agent that forms a bridge between two amino acids and does not itself incorporate macromolecules of collagen, hyaluronic acid and gelatin. The bracket has Wherein X-based gelatin-gelatin, gelatin-collagen or gelatin-hyaluronic acid; Y-series collagen-collagen, collagen-gelatin or collagen-hyaluronic acid; Z-system hyaluronic acid-gelatin, hyaluronic acid-collagen or hyaluronic acid-hyaluronic acid; Is 1 or an integer greater than 1. The present invention is a porous, three-dimensional (three-dimensiona) biomedical material, wherein the ratio of collagen is 30% to 45%, the ratio of hyaluronic acid is 0.1% to 5%, and the ratio of gelatin is 50% to 70%. The limitation is that the total ratio of collagen, hyaluronic acid and gelatin is 100%. In a preferred embodiment, the biomedical material has a pore size of from 10 to 500 μm. In a more preferred embodiment, the pore size is from 50 to 200 μm. The choice of crosslinker is important for the preparation of porous biomedical materials, which can result in biomedical materials of different structures. The pore size of the biomedical material can be varied according to different needs, and the thickness can also be adjusted by controlling the volume or weight/volume percentage concentration of each material. In one embodiment, the biomedical material has an average thickness of about 1 mm to mimic the true thickness of the human skin epidermis and dermis. The biomedical material of the invention has high water absorption capacity, and the expansion rate of the biomedical material is about 20 times larger than that of the dry stent. In a preferred embodiment, the expansion rate of the biomedical material is about 25 times greater than that of the dry stent. Therefore, the biomedical material has a large porous layered matrix space to increase its water content. The cross-linked porous form provides the opportunity for cells to be inoculated into the scaffold. The interconnected pores in biomedical materials provide the opportunity for cytokines released by cells (such as keratinocytes, melanocytes, and dermal fibroblasts) to interact with growth factors.

The biomedical material of the present invention promotes the secretion of collagen by fibroblasts and further reduces the neutrophil infiltration of the wound and increases the density of the epidermis in the wound. Therefore, the biomedical material of the present invention can be further used for wound healing or artificial skin.

The biomedical material of the present invention has good biocompatibility. The biomedical material can Further co-culture with fibroblasts, keratinocytes and melanocytes to form skin equivalents. Fibroblasts are inoculated prior to inoculation of keratinocytes and melanocytes, while keratinocytes and melanocytes are grown on biomedical materials with fibroblasts. This three-dimensional structure mimics the physiological environment and can be further used in skin-related experiments, including testing and Clinical application. For example, large-scale screening for compounds that interact with skin cells and the three-dimensional structure formed using the biomedical materials of the present invention illustrate their regulatory mechanisms.

The invention also provides a method for preparing the biomedical material of the invention, comprising: (a) preparing a mixture of collagen, hyaluronic acid and gelatin; (b) lyophilizing the mixture of step (a); (c) step (b) a mixture of the organic solution containing EDC; (d) removing the mixture from the organic solution containing EDC; and (e) lyophilizing the mixture to form the biomedical material. The concentration of collagen, hyaluronic acid and gelatin in the mixture of step (a) is about 0.5 to 2000 μM, 0.0025 to 1 μM and 0.1 to 40 mM, respectively. The concentration of EDC in the organic solution is about 2.5 to 1000 mM. The organic solution in the present invention includes, but is not limited to, the following solvents: alkanes, alcohols, ketones, esters, alkenones, and combinations of the foregoing.

In a preferred embodiment, the concentration of collagen is about 10 to 1000 μM, the concentration of hyaluronic acid is about 0.0125 to 0.2 μM, and the concentration of gelatin is about 0.5 to 8 mM. In a more preferred embodiment, the concentration of collagen is about 50 to 200 μM, the concentration of hyaluronic acid is about 0.025 to 0.1 μM, and the concentration of gelatin is about 1 to 4 mM.

In a preferred embodiment, the concentration of EDC in the organic solution is from about 5 to 500 mM. In a more preferred embodiment, the concentration of EDC in the organic solution is from about 10 to 250 mM. In another preferred embodiment, the organic solution is an alcohol.

The invention further provides a method for promoting wound healing comprising covering a biomedical material of the invention onto a wound. The biomedical material enhances wound healing by promoting secretion of collagen by fibroblasts, reducing neutrophil infiltration of the wound, and increasing epidermal density in the wound.

The invention is not limited to the examples mentioned below, but the following examples are merely representative of the various aspects and features of the invention.

Example 1 Preparation of biomedical materials and evaluation of their characteristics Biomedical materials for the manufacture of collagen/HA/gelatin sponge bioporous scaffolds

Collagen (Cat No. C7774, MW: 64,000), gelatin (Cat No. G9539, MW: 5,000) and N-ethyl-N'-3-[3-dimethylaminopropyl]carbodiimide (EDC) (Cat No. E1769) were purchased from Sigma-Aldrich Chemical (St. Louis, MO); HA (grade FCH-200, MW: 2-2.1 MDa) was purchased from Kibun Food Chemicals (Tokyo, Japan). Collagen/HA/gelatin were mixed into a solution at a final concentration of 93.75 μM, 0.05 μM and 2 mM, respectively, and the mixed solution was gently poured into a 6-cm culture dish and frozen at -20 ° C for 48 hours. The mixed solution was lyophilized for 24 hours to construct a collagen/HA/gelatin scaffold.

This was followed by chemical incubation at 25 ° C for 24 hours in pure ethanol containing 50 mM EDC. The reaction was stopped after 24 hours of removal of the EDC solution and washed several times with distilled water to remove any unreacted chemical (EDC). The stent was again lyophilized for 48 hours and sterilized by ethylene oxide gas. The dried collagen/HA/gelatin scaffold is made into a suitable size for subsequent use. Figure 1 is drawn based on reactions originating from established in-phase reaction paths.

Analysis of the expansion rate of biomedical materials

Several collagen/HA/gelatin scaffold samples (30 mg) were immersed in distilled water for 24 hours at 25 ° C, respectively. After removing the water, the bracket was hung until no moisture was visible and then weighed. The moisture absorption of the expansion stent is calculated by the following equation: moisture absorption = (W _{w -} W _d ) / W _d

W _w is the weight of the expansion bracket, and W _d is the weight of the dry bracket. The results were compared to the other four commercially available wound dressings.

Absorbing moisture to get important nutrients is essential for cell growth. The absorption ratio is shown in Figure 2. The water absorption capacity of the three components cross-linked with EDC is 15 to 30 times that of the dry stent. HA and gelatin are also the same; the water absorption capacity without HA is reduced to 15 times, indicating that HA is important for high water absorption properties. Factors and consistent with common sense. Although the stent containing only HA and gelatin still has a high expansion rate, collagen is an important component of the epidermis, so collagen is also used to make the stent. The concentration of the three components was lowered to evaluate the effect on the expansion ratio, and as a result, the water absorption did not change significantly, and the stent which was not subjected to the crosslinking reaction was dissolved in water and the water absorption ability could not be measured. Comparing the four commercially available skin dressings, the overall expansion rate was less than 10 times, indicating that the sponge-like scaffold retains a large porous layered matrix space to increase its water retention capacity. In general, the backbone of the three components has a large number of negatively charged carboxyl groups and is hydrophilic.

Evaluation of degradation rate of biomedical materials by lysozyme, collagenase and hyaluronic acid

The biomedical material (n=6) was accurately weighed and immersed in 1 ml of 0.1 M Tris-HCl and 0.05 M CaCl ₂ (pH 7.4) containing 10 and 20 U of collagenase I (Sigma) at 37 °C. In the solution. After a specific time, 0.2 ml of 0.25 M ethylenediaminetetraacetic acid (EDTA) was added to terminate the decomposition. The remaining scaffolds were washed three times with distilled water and lyophilized. The degradation of the biomedical material is calculated as the remaining stent weight and is expressed as a percentage of the starting weight. Hyaluronic acid and lysozyme use a similar experimental procedure. The scaffold samples were suspended in PBS (pH 7.4) containing 30 or 50 U/ml hyaluronidase, and cultured at 37 ° C for 1, 3, 5 and 7 days. The degradation of the biomedical material by lysozyme was tested by incubating the scaffold at 37 ° C in PBS containing lysozyme (10,000 and 30,000 U/ml) (pH 7.4) for up to 21 days. After the degradation is complete, the sample is removed and washed for subsequent measurements. The rate of degradation of the uncrosslinked biomedical material (n=6) was calculated by dividing the remaining weight of the biomedical material by the starting weight for comparative analysis.

The biostability of collagen/HA/gelatin was tested by in vitro degradation by lysozyme, collagenase and hyaluronan, respectively. As shown in Figure 3A, the biomedical material was not completely degraded by 30,000 U/ml lysozyme after 7 days; in Figure 3B, the biomedical material was not completely degraded by 30 and 50 U/ml hyaluronidase after 7 days. The biomedical material was completely decomposed after incubation with 20 U/ml collagenase for 3 hours (Fig. 3C).

Example 2 Cells grow in porous scaffolds Human skin primary cultures

The human keratinocyte cell line was cultured from the primary culture of the foreskin, and the material was obtained from the Zhonghe Memorial Hospital attached to Kaohsiung Medical University. Human keratinocytes were cultured in a Keratinocyte-SFM (10724; GIBCO ^TM ) medium supplemented with bovine pituitary extract (BPE, Cat No. 13028-014) and recombinant human epidermal growth factor (EGF Human Recombinant, Cat No. 10450-013). in. The culture medium and growth additives of keratinocytes include epidermal growth factor, BPE, insulin, fibroblast growth factor, and calcium ion (0.09 mM). First generation of primary human epidermal foreskin melanocytes (HEMn-MP) were purchased from Cascade Biologics ^TM, were cultured in 254 medium (Medium 254, M-254-500; Cascade Biologics TM) and adding human melanoma cells growth supplement (HMGS, Cat No.S-002-5). The 254 medium is a basal medium containing essential and non-essential amino acids, vitamins, organic compounds, trace elements, and inorganic salts. Human melanocyte growth supplement contains bovine pituitary extract, fetal bovine serum, bovine insulin, bovine transferrin, basic fibroblast growth factor, hydrocorticosterone, heparin and phorbol 12-tetradecanoic acid Ester 13-myristate 13-acetate. The primary culture of human skin fibroblasts was donated by Dr. Wu Qingying (Korean Medical University Medical Research Institute, Center for Environmental Medicine). All cell types were cultured in a humidified incubator at 37 ° C, 5% CO ₂ .

Trypan blue analysis

All cells were treated with 1 x tryptone-EDTA in PBS (BioWest) and diluted in PBS with 0.5 ml PBS, and 0.5 ml trypan blue solution (0.4% w/v) was added. The stained cell line was sampled with a Pasteur pipette and sent to the hemocytometer using capillary action. A total of at least 500 cells were counted and blue cells were counted separately.

Cell attachment rate and survival rate

The stent was sterilized with ethylene oxide, pre-wetted to remove residual ethylene oxide, and then placed in a 24-well plate. 100 μl of the cell suspension (5 × 10 ⁵ cells / 100 μl) was added to the surface of each pre-wet stent and allowed to infiltrate into the stent. The cells/scaffold constructs were then incubated for 4 hours at 37 ° C, 5% CO ₂ to attach the cells. After the cells were attached, the cell/stent construct was transferred to a new 24-well plate to remove the cells lost at the bottom of the plate, and 0.5 ml of the medium was added to the well containing the cell/scaffold construct. The medium was changed every 2 days and the plate was shaken while culturing. Cell/scaffold constructs were collected at each given time for further experimental analysis.

Analysis by 3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfonic acid phenyl)-2H-tetrazolium (MTS) To study the cell viability and proliferation rate of cells seeded on the scaffold (Han et al., Breast Cancer Res 11, R57 (2009)). MTS is reduced by dehydrogenase in living cells and the yellow formazan product is released back into the medium. The amount of formazan product is measured by the absorbance at 490 nm and is proportional to the number of viable cells in the culture. The cells in 100 μl of the medium were placed in a 20 μl CellTiter 96 AQueous One Solution (Promega, Cat No. G3582) for 3 hours according to the instruction manual, and the absorbance at 490 nm was recorded using a spectroscopic disc analyzer.

To determine the rate of cell attachment on the scaffold, the cell viability of the human fibroblasts seeded into the scaffold was first determined by testing cell viability. Approximately 75% of the cells seeded with the scaffold successfully attached to the scaffold compared to the same amount of cells directly seeded in a 24-well plate (Fig. 4A). The number of cells in which human fibroblasts survived on a spongy scaffold was measured by MTS analysis. On the 14th day after the inoculation, the cell density of the fibroblasts grown in the scaffold was greatly increased, indicating that the scaffold of the biomedical material of the present invention has the advantages of enhancing cell proliferation, differentiation, and viability.

Type of porous scaffold/biomedical material (SEM image)

The shape characteristics of the porous scaffold were observed using a scanning electron microscope image (SEM, JEOL, Tokyo, Japan). The scaffold was fixed overnight with 2.5% glutaraldehyde in 0.1 M sodium phosphate buffer (pH 7.2) and fixed in 1% osmium tetroxide. (post-fix) 1 hour, dehydrated in ethanol (30, 50, 75 and 99.5%) and subjected to critical point drying. The dried stent samples were coated with gold particles through a sputter coater at ambient temperature and microscopic images of the stent were taken at appropriate sizes (100 x and 200 x). The pore size distribution was measured by a Beckman Coulter LS32 apparatus in the range of 0.01 μm to 1000 μm. Each SEM image has a pore size of 30 pores and a total of 5 SEM images were measured and the average pore size was calculated.

The type profile of the collagen/HA/gelatin scaffold recorded by SEM is shown in Figure 4B. The scaffold has a highly porous structure interconnected, and the surface of the pore wall that has not been treated with fibroblasts is smooth and homogeneous. The SEM image of the spongy scaffold showed an open macroscopic porous structure with a pore size in the range of 132.5 ± 8.4 μm. The SEM image of the scaffold treated with fibroblasts (14 days) is shown in Figure 4B. The cell wall of the cell-treated scaffold has a rough surface and is composed of a plurality of fissures, which should be caused by degradation of fibroblasts.

3D human skin cell co-culture In vitro cell culture and fluorescence analysis

To detect the distribution of skin cells, cells were stained with PKH37 prior to inoculation into the scaffold. Cells were incubated with 5 μM PKH-67 (a green fluorescent compound, embedded in the cell membrane by selective partitioning of aliphatic reporter molecules; Sigma-Aldrich) at 25 ° C for 5 minutes with slight shaking every 30 seconds ( According to the operation manual). The medium was washed to remove the uninserted PKH-67, and the cells labeled with PKH-67 were placed on the surface of the stent at a density of 1 × 10 ⁵ /cm ² , and then the cells were collected during different cultures.

Human skin cells were visualized by fluorescence microscopy after incubation with collagen/HA/gelatin scaffolds (Fig. 5). Three skin cells in collagen eggs The white/HA/gelatin scaffold can be normally proliferated in the presence of a stent, and the scaffold material has been shown to be beneficial for cell growth. The extent of cross-linking has been confirmed to correlate with the distribution of porous structures and water retention capacity in the scaffold. The biomedical materials of the present invention have pore size and moisture absorbing ability suitable for human skin cell growth.

Immunofluorescence analysis of paraffin sections of skin mimics

A human skin model system, inoculated with 10 ⁶ 10 ⁵ keratinocytes and melanocytes to have been seeded with 5 × 10 ⁵ to give the fibers were blasts bracket 7 days. After keratinocytes and melanocytes were inoculated, they were cultured for an additional 7 days, and the medium was changed every 2 days. During the co-cultivation period, the amount of the medium and the cells were mixed in the same ratio.

The experimental procedure was slightly modified according to the published procedure (Dainiak et al., Biomaterials 31, 67-76 (2010); Wu et al., Biomaterials 31, 631-640 (2010)). The scaffold samples co-cultured with keratinocytes, melanocytes, and fibroblasts were fixed with 4% formaldehyde prepared in PBS for 24 hours at room temperature, and the samples were embedded in paraffin with a section thickness of 5 μm. The sections were dewaxed and then infiltrated with 3% H ₂ O ₂ in PBS for 15 minutes at room temperature, then blocked with fibroblasts for 1 hour and combined with anti-cytokeratin (for keratinocytes) or anti-s Primary antibody culture of -100 (for melanocytes). The sections were then washed and incubated with cy3 conjugated goat anti-rabbit antibody (Millipore) and FITC conjugated goat anti-mouse antibody for 30 minutes at room temperature with 4,6-mercapto-2-phenylindole (DAPI) ( Vector, Burlingame, CA) counterstaining, immunofluorescence images (TE300; Nikon, Japan). Sections were also stained with hematoxylin and eosin (H&E) stain to confirm cell location.

In order to construct a model that simulates the condition of real human skin, a 3D human skin cell mimic co-culture experimental procedure was established (Fig. 6A). A human skin model system, inoculated with 10 ⁶ 10 ⁵ keratinocytes and melanocytes to have been seeded with 5 × 10 ⁵ to give the fibers were blasts bracket 7 days. After the co-culture was further cultured for 7 days, the samples were vertically sectioned and stained with immunofluorescence and observed under a microscope. Figure 6B is a paraffin section of a skin mimic in a bright field; Figure 6C shows cells stained with DAPI; and in Figure 6D, the keratinocyte line is stained with anti-cytokeratin, FITC, and keratinocytes are green fluorescent; Figure 6E In the middle, the keratinocyte line was conjugated with anti-s100 protein and cy3, and it was red; Figure 6C-E was combined to obtain Fig. 6F. As shown in Fig. 6F, melanocytes and keratinocytes were distributed thereon, and DAPI stained but not Cells stained with anti-cytokeratin or anti-s-100 protein are fibroblasts.

Collagen content

To measure the total amount of collagen synthesized by the fibroblasts in the scaffold, all collagen was stained using Sirius Red dye (Direct Red; Sigma). Fibroblasts were cultured in 48-well plates or scaffolds and fibroblasts, keratinocytes, and melanocytes were co-cultured on the 2D surface or the amount of collagen secreted in the scaffold was compared. After the indicated incubation time, the medium was removed and the cells were washed twice with PBS. Add 100 μl of 0.1% Sirius Red Dye (per 50 ml of picric acid 0.05 g Sirius Red Dye) to each well and maintain at room temperature for 1 hour, then remove the unattached dye and 0.1 N at 200 μl Wash HCl five times. The attached dye (15 minutes) was extracted with 100 μl of 0.1 N NaOH and mixed well. The dye was placed in a 96-well plate and the absorbance at 540 nm was read using a microplate analyzer.

Collagen in ECM imparts appropriate mechanical force to the tissue, and collagen is also important for the production of ECM by fibroblasts. Use Sirius Red Dye The collagen secreted from the fibroblasts in the scaffold is stained. The collagen content of only the scaffold and fibroblasts cultured in the scaffold for 7 days, the collagen content of the fibroblasts inoculated on the 48-well plate for 7 days, the inoculated fibroblasts were cultured in the scaffold for 7 days, and then the keratinocytes and melanocytes were inoculated. Collagen content for 7 days of culture was compared to examine the effect of keratinocytes and melanocytes on collagen secretion by fibroblasts. In Figure 7, the amount of collagen secreted by the fibroblasts inoculated into the scaffold is about 30% greater than that of the 48-well plates. Collagen is an important component of extracellular matrix (ECM), and cell proliferation, tissue formation and tissue shape are related to collagen concentration. Therefore, the amount of collagen secretion is an important indicator for detecting the possibility of wound healing. Since the scaffold has the potential to promote collagen secretion by fibroblasts, it may be a good material for tissue engineering in wound healing.

Example 3 Wound healing rat animal model Laboratory animal preparation

Wistar rats (250-285 g) were used in all experiments. The rats were placed in a Plexiglas cage in a temperature controlled room (22 ± 1 ° C) for 12 hours light / 12 hours dark, and food and water were available at will. Rats were randomly divided into 2 groups, trauma group and treatment group. The cutting wound healing test was modified from (Huang and Yang, Int J Pharm 346, 38-46 (2008)). After the rats were anesthetized, the dorsal hair was shaved with an electric razor, and a full thickness excision of 2 cm in diameter was cut with a scalpel. After the cutting treatment, the biomedical material of the same size as the wound was rinsed with saline and immediately covered on the wound, while the wound of the wound group was not covered for comparison. After the surgery, the rats were placed in individual cages to heal the wound.

Wound size assessment

Photographs were taken with the same parameters (F7.2, 1/60) using a digital camera (Coolpix P6000, Nikon, Japan) on days 1, 2, 3, 4, 5, 7 and 10 after the injury. The area of each wound was measured using SPOT (Diagnostic Instruments, Inc., Sterling Heights, MI, USA) software. The extent of wound healing is expressed as a percentage of the wound area and is calculated as follows: (N Day Wound Area / Day 0 Wound Area) x 100%.

The wound healing efficiency of the stent was assessed in a full-thickness wound pattern. The wound area of the wound group and the treatment group decreased with time (Fig. 8), and the wound area of the treatment group was 79.7±3.4% and 73.4±3.5% on days 1, 2, 3, 4, 5 and 7 after the wound, respectively. 66.8 ± 2.2%, 60.7 ± 5.0%, 58.3 ± 6.1% and 44.9 ± 4.3%, all less than the wound area of the trauma group (97.4 ± 5.5%, 86.3 ± 2.2%, 75.3 ± 3.7%, 71.4 ± 3.8%, 67.9 ±8.1% and 62.2±9.4%). One day after wounding, the wound area of the treated group was smaller than that of the wound group. This trend was the same 10 days after the wound. The wound in the treated group was closed faster than the wound group. On the 10th day after the wound, the wound area of the stent was 24.0 ± 2.1%, and the wound area of the wound group was 41.8 ± 5.3%. This stent can significantly increase the rate of wound healing because the wound healing of the stent treatment reached more than 45% of the wound closure in the first 7 days, almost within 10 days. Up to 75% of wound closure can be achieved.

Tissue sections

Tissue sections after H&E staining (Fig. 9) showed granulation tissue in both the treatment group and the wound group, and the stent did not interfere with wound healing. The epidermis of the treatment group was denser than that of the trauma group. The stent improved the skin strength during wound healing. Compared with the trauma group, the treatment group had less neutrophil infiltration; the wound healing process Among them, neutrophils secrete substances that accelerate keratinocyte differentiation and delay wound closure, and the application of this stent can reduce neutrophil infiltration and accelerate wound closure.

This stent improves healing rate, increases epidermal density, and has less neutrophil infiltration. This evidence suggests that this stent is suitable for cutting wound healing.

The rats were sacrificed after anesthesia 10 days after the wound, and the wound skin was fixed in 4% polyoxymethylene. The skin was stained with hematoxylin-eosin for histological observation. Histological analysis images were taken using a Spot Xplorer CCD integrated camera (Diagnostic Instruments, Inc., Sterling Heights, MI, USA) and a Leica DM-6000 microscope (Leica, Wetzlar, Germany). The histological analysis was modified from Bayat et al. (2005). The skin of each sample was randomly selected from 10 regions to calculate the number of neutrophils at a magnification of 400×, and the histological examination was performed in a blind manner.

While the invention has been described and illustrated in detail, it will be understood that

It will be readily apparent to those skilled in the art that the present invention can be readily achieved, and the results and advantages mentioned, as well as those inherent in nature. The processes and methods of manufacture are representative of the preferred embodiments, which are exemplary and are not intended to limit the scope of the invention. Those skilled in the art will be aware of the modifications and other uses therein. These modifications are intended to be within the spirit of the invention and are defined in the scope of the claims.

Figure 1 illustrates the production of biomedical materials, skin culture and rat skin wound healing mode. (A), (B) schematic diagram of cross-linking of collagen, hyaluronic acid (HA) and gelatin by EDC; (C), (D) scanning electron micrograph of collagen/HA/gelatin scaffold with pore size of 132.5±8.4 Mm.

Figure 2 illustrates the expansion rate analysis (50 mM) (n = 3) of scaffolds made by cross-linking of proprotein, HA and gelatin by EDC. The scaffold without the EDC cross-linking reaction was dissolved in water (symbol: x). Compared with commercially available materials, including: Du (DuoDERM 9C52552), Hy (Hydro Coll), Te (Tegaderm M1635) and ME (MEDPOR ^® ).

Figure 3 illustrates the degradation rates of lysozyme (A), hyaluronan (B) and collagenase (C).

Figure 4 illustrates cell proliferation (n = 4) of human skin fibroblasts (FBs) seeded on scaffolds. From 1 to 14 days, the rate of proliferation can be observed by MTT assay (A); SEM images of fibroblasts seeded on scaffold for 14 days (B).

Figure 5 illustrates photographs of human keratinocytes (KCs), melanocytes (MCs), and fibroblasts (FBs) cultured in scaffolds in bright field, fluorescent, and combined images. Cell staining was performed using the fluorescent compound PKH-67 (green).

Figure 6 illustrates the preparation procedure of 3D human skin mimics (A); (B) paraffin sections of 3D human skin mimics in a microscope bright field (400 x); (CE) KCs, MCs and FBs cultured in scaffolds for 14 days Fluorescent images, stained with DAPI (blue); KCs (green) with anti-cytokeratin; MCs (red) with anti-s-100; (F) images of combined KCs, MCs and FBs, arrows KCs, MCs, and FBs are indicated in a specific color.

Figure 7 illustrates the amount of collagen secreted by FBs seeded on a culture plate or scaffold. FBs were grown on culture plates or scaffolds for the first 7 days, and then inoculated with KCs and MCs for 7 days.

Figure 8 illustrates the healing of wounds in the stent-treated group (A, top) and wound group (A, lower) at 0, 1, 2, 3, 4, 5, 7 and 10 days after wounding; evaluation of stents in full-thickness wound mode The effect of wound healing, after anesthesia, a full-thickness cut of 2 cm in diameter was cut out from male Wistar rats with a scalpel. The treatment group covered the stent on the wound immediately after cutting; the wound group was not covered for comparison. From day 1 to day 10 after trauma, the wound healing in the wound group was slower than that in the stent-treated wound, with a scale of 0.5 cm; (B) the ratio of wound contracture between the stent-treated group and the trauma group at different times. The wound area was examined on each exact day and the proportion of wound area reduction was calculated. The wound surface area was calculated by the method described in the examples. From the first day after the trauma, the wound area was reduced faster than the control group. On the 7th day, the wound area of the control group was 60%, and the stent treatment group reached this ratio on the 3rd day. There was a statistically significant difference between the wound group and the stent-treated group on the 10th day. The data were expressed as mean error of the mean (SEM). The significant difference was defined as P<0.05, and * represents significant.

Figure 9 illustrates the evaluation of skin wounds with hematoxylin and eosin (H&E) stained sections. The rats were sacrificed 10 days after the wound, the wound skin was fixed with 4% polyoxymethylene, the skin was histologically observed by H&E staining, and the skin of each sample was randomly selected from 10 regions to calculate the number of neutrophils at a magnification of 400×; The wounds after 10 days of trauma were the stent treatment group (A), the trauma group (B) and the control group (C). The wounds in the stent treatment group and the wound group all had granulation tissue. The epidermis of the treatment group was denser than the trauma group. The neutrophil infiltration of the wound was less than that of the trauma group (D). Scale bar = 200 μm; EP: epithelial cell layer, GT: granulation tissue, *: p<0.05 compared with control group, **: p<0.05 compared with trauma group and control group.

Claims (20)

A biomedical material comprising a scaffold consisting of collagen, hyaluronic acid and gelatin, wherein collagen, hyaluronic acid and gelatin are ethyl-3-[3-dimethylaminopropyl]carbodiimide (ethyl) -3-[3-dimethylaminopropyl]carbodiimide, EDC) cross-links two components of collagen, hyaluronic acid and gelatin. According to the biomedical material described in claim 1, wherein the stent has: Wherein X-based gelatin-gelatin, gelatin-collagen or gelatin-hyaluronic acid; Y-series collagen-collagen, collagen-gelatin or collagen-hyaluronic acid; Z-system hyaluronic acid-gelatin, hyaluronic acid-collagen or hyaluronic acid-hyaluronic acid; Is 1 or an integer greater than 1. According to the biomedical material of claim 1, wherein the ratio of the collagen is 30% to 45%, the ratio of the hyaluronic acid is 0.1% to 5%, and the ratio of the gelatin is 50% to 70%. The limitation is that the total ratio of collagen, hyaluronic acid and gelatin is 100%. According to the biomedical material described in claim 1 of the patent application, it is a porous, three-dimensional biomedical material. The biomedical material according to item 1 of the patent application has a pore size of 10 to 500 μm. The biomedical material according to claim 5, wherein the pore size is 50 to 200 μm. According to the biomedical material of claim 1, the expansion ratio is 20 times larger than that of the dry stent. According to the biomedical material described in claim 1 of the patent application, it is used for wound healing. According to the biomedical material described in the first aspect of the patent application, it can be further cultured with fibroblasts, keratinocytes, and melanocytes to produce a skin equivalent. A method for preparing a biomedical material according to claim 1, comprising: (a) preparing a mixture of collagen, hyaluronic acid and gelatin; (b) lyophilizing the mixture of the step (a); and (c) performing the step The mixture of (b) is cultured in an organic solution containing EDC; (d) the mixture is removed from the organic solution containing EDC; and (e) the mixture is lyophilized to form the biomedical material. The method of claim 10, wherein the concentration of the collagen is about 10 to 1000 μM, the concentration of the hyaluronic acid is about 0.0125 to 0.2 μM, and the concentration of the gelatin is about 0.5 to 8 mM. The method of claim 11, wherein the concentration of the collagen is about 50 to 200 μM, the concentration of the hyaluronic acid is about 0.025 to 0.1 μM, and the concentration of the gelatin is about 1 to 4 mM. The method of claim 10, wherein the organic solution is selected from the group consisting of an alkane, an alcohol, a ketone, an ester, and an enone. The method of claim 10, wherein the organic solution is selected from the group consisting of alcohols. The method of claim 10, wherein the concentration of the EDC in the organic solution is about 5 to 500 mM. The method of claim 10, wherein the concentration of the EDC in the organic solution is about 10 to 250 mM. A method for promoting wound healing comprising covering a wound of a biomedical material as described in claim 1 of the patent application. The method according to claim 12, which promotes wound healing by promoting secretion of collagen by fibroblasts. According to the method of claim 12, the wound healing is promoted by reducing the neutrophil infiltration of the wound. According to the method of claim 12, the wound healing is enhanced by increasing the density of the epidermis in the wound.