Novel porous scaffold and method for producing same

Abstract

The present invention relates to a porous scaffold having excellent tissue engineering properties and a manufacturing method thereof. The scaffold of the present invention can be manufactured by a simple process and exhibits high tensile strength, biocompatibility, and an extremely high cell engraftment rate, and therefore can be useful as a scaffold composition for human transplantation in various applications, including artificial ligaments and abdominal wall reinforcement supports.

Description

[Technical field]

[0001] The present invention relates to a biocompatible porous scaffold, a support composition for human implantation containing the same, and a method for producing the same. [Background technology]

[0002] Recently, with the great development of bioengineering, material engineering and surgery, tissue engineering, which aims to replace and regenerate lost body tissues, has made remarkable progress. Tissue engineering is a fusion of life science, engineering and medicine, and aims to understand the correlation between the structure and function of biological tissues, and based on this, to replace or regenerate damaged tissues and organs with normal tissues, by using artificial tissues that can be implanted in the body to maintain, improve or repair bodily functions.

[0003] Loss of body tissues is due to various causes such as degenerative diseases, trauma, surgical removal of tumors, and certain congenital malformations, but these can only be restored by regenerating the irreversibly lost tissue. In order to induce the regeneration of lost tissues by tissue engineering, it is important to first manufacture a biodegradable polymer scaffold similar to living tissue. The main requirement of the scaffold material used for the regeneration of human tissues is that it must adequately play the role of a substrate or framework so that tissue cells can adhere to the material surface to form tissues with a three-dimensional structure, and also act as an intermediate barrier between the transplanted cells and the patient's cells.

[0004] After the scaffold is implanted in the subject, it must induce the engraftment of cells necessary for tissue regeneration, initiate the formation of new tissue, and then disappear over time to fill the space left by the newly formed tissue. Therefore, the scaffold should be biodegradable without requiring surgical removal, should not induce immune rejection, inflammatory reaction, or long-term fibrous encapsulation, should not experience shrinkage of the graft volume, and should be free from serious complications such as prosthetic implants.

[0005] Therefore, in order to effectively induce tissue regeneration while naturally disappearing after playing the role of physical support for an appropriate period of time, the scaffold must have a certain level of mechanical strength and elasticity as well as biodegradability. For this reason, it is very important to select the most suitable natural or synthetic polymer and the most suitable structure.

[0006] Throughout this specification, a number of papers and patent documents are referenced and citations are displayed, the disclosures of which are incorporated herein by reference in their entirety in order to more clearly describe the state of the art to which the present invention pertains and the contents of the present invention. Summary of the Invention [Problem to be solved by the invention]

[0007] The present inventors have made extensive research efforts to develop an efficient scaffold for tissue regeneration that has sufficient physical strength and excellent biocompatibility, and can be manufactured in a relatively simple process. As a result, they have found that when a mesh-like scaffold having pores of a certain size and strands of a certain diameter is manufactured from a first polymer, and the surface of the mesh-like scaffold is coated with a second polymer different from the first polymer and has biocompatibility, it shows not only high tensile strength and biocompatibility but also a remarkably excellent cell engraftment rate, and can be used as a scaffold for human transplantation for various purposes, including artificial ligaments and supports for reinforcing the abdominal wall. In addition, they have found that when two biocompatible polymers having different structures and functions are manufactured into a three-dimensional porous structure and a two-dimensional porous structure, respectively, and then joined together, it shows significantly improved physical properties while maintaining its inherent functions, such as tissue regeneration, wound healing, and providing in vivo bonding strength, and thus completed the present invention.

[0008] It is therefore an object of the present invention to provide a porous scaffold and a method for making the same.

[0009] Another object of the present invention is to provide a support for human implantation comprising said porous scaffold.

[0010] Other objects and advantages of the present invention will become more apparent from the following detailed description of the invention, the claims and the drawings. [Means for solving the problem]

[0011] According to one aspect of the invention, the invention provides a method for producing a porous scaffold comprising the steps of: (a) 0.1-0.5 mm from the first polymer solution 2 producing a polymer mesh having pores of 0.1-0.3 mm and a strand diameter of 0.1-0.3 mm; (b) coating the surface of the resulting polymer mesh with a biocompatible second polymer solution.

[0012] The present inventors have made intensive research efforts to develop an efficient tissue regeneration scaffold that has sufficient physical strength and excellent biocompatibility, can be manufactured by a relatively simple process, and has been the subject of numerous studies. 2 They found that when a mesh having pores of 0.1-0.3 mm diameter and consisting of strands was fabricated and its surface was coated with a second polymer that is different from the first polymer and has biocompatibility, it not only exhibited high tensile strength and biocompatibility but also an extremely excellent cell engraftment rate, and could therefore be used as a scaffold for human transplantation for a variety of purposes, including artificial ligaments and supports for abdominal wall reinforcement.

[0013] As used herein, the term "scaffold" refers to a tissue engineering structure for promoting the recovery and regeneration of damaged tissue by attaching biological cells, specifically, cells derived from damaged tissue or cells involved in the recovery of damaged tissue. The term "cell attachment" refers to the direct or indirect attachment of cells to a substrate or other cells while maintaining their inherent biological activity. Specifically, the scaffold of the present invention may be a planar structure consisting of a single mesh, or may be a three-dimensional structure in which multiple meshes are laminated.

[0014] As used herein, the term "polymer" refers to a synthetic or natural polymeric compound in which monomers of the same or different types are successively linked together. Thus, polymers include homopolymers (polymers in which one type of monomer is polymerized) and hybrid polymers made by polymerization of at least two different monomers, which in turn include copolymers (polymers made from two different monomers) and polymers made from more than two different monomers.

[0015] According to a specific embodiment of the present invention, the first polymer used in the present invention is selected from the group consisting of PCL (polycaprolactone), PLLA (poly(L-lactic acid)), PGA (poly(glycolic acid)), PLGA (poly(lactic-co-glycolic acid)), LCL (poly(L-lactide-co-ε-caprolactone)), and combinations thereof.

[0016] More specifically, the first polymer is PCL (polycaprolactone).

[0017] According to the present invention, the first polymer of the present invention forms a mesh in which strands having a certain thickness are crossed at a certain interval and have pores of a certain size. The pores must have an optimal size not only for cell attachment, proliferation, and activity maintenance, and for induction of new blood vessels during tissue regeneration, but also for the mechanical strength and elasticity of the mesh itself, so that the ultimate objective of the present invention, recovery and regeneration of damaged tissue, can be efficiently achieved. Therefore, the optimal pore area is specifically 0.1-0.5 mm2. 2 More specifically, 0.1-0.4 mm 2 and more specifically 0.2-0.3 mm 2 , most specifically about 0.25 mm 2 It is.

[0018] As used herein, the term "pore area" refers to the average area of ​​repeated pores that appear due to the intersection of strands in the mesh structure of the present invention made from the first polymer, and such area refers to the area measured before coating with the second polymer solution described below.

[0019] In addition, in order to ensure the physical properties suitable for a human implant scaffold having appropriate elastic modulus and tensile modulus, the diameter of the strands forming the mesh is important in addition to the pore area mentioned above. Therefore, the suitable diameter of the strands is specifically 0.1-0.3 mm, more specifically 0.15-0.25 mm, and most specifically about 0.2 mm.

[0020] The step of producing a polymer mesh from the first polymer solution of the present invention can use various methods known in the art, including, but not limited to, 3D printing, solvent-casting particulate leaching, gas foaming, fiber meshes / fiber bonding, phase separation, melt molding, freeze drying, and electrospinning.

[0021] According to the present invention, the scaffold of the present invention is provided with biocompatibility in addition to the mechanical strength described above by coating a polymer mesh made from a first polymer solution with a biocompatible second polymer solution.

[0022] As used herein, the term "biocompatibility" refers to the property of not causing short-term or long-term side effects when administered into a living body and comes into contact with cells, tissues, or body fluids of an organ, and specifically includes not only tissue compatibility and anticoagulant compatibility, which do not cause necrosis of tissue or coagulation of blood when in contact with biological tissue or blood, but also biodegradability, which disappears after a certain period of time has passed after administration to the body.

[0023] As used herein, the term "biodegradable" refers to the property of being naturally decomposed when exposed to a physiological solution of pH 6-8, specifically, the property of being decomposed over time in vivo by body fluids, decomposing enzymes, microorganisms, etc. Biodegradable polymers that can be used in the present invention include any synthetic or natural polymers as long as they are biodegradable as described above, including, but not limited to, collagen, gelatin, chitosan, hyaluronic acid, poly(valerolactone), poly(hydroxybutyrate), poly(hydroxyvalerate), and combinations thereof.

[0024] According to a specific embodiment of the invention, the biocompatible second polymer is a natural polymer, more specifically collagen, most specifically type I collagen.

[0025] According to a more specific embodiment of the invention, the collagen solution is used in a concentration of 0.2-0.8% (v / v), even more specifically in a concentration of 0.3-0.7% (v / v), and most specifically in a concentration of 0.4-0.6% (v / v).

[0026] In the present specification, the term "coating" means to form a new layer of a certain thickness by modifying a specific substance on a target surface, and the target surface and the coating substance are modified by ionic or non-covalent bonds. The term "non-covalent bond" is a concept including not only physical bonds such as adsorption, cohesion, entanglement, and entrapment, but also bonds that occur by interactions such as hydrogen bonds and van der Waals bonds acting alone or together with the physical bonds. In the present invention, when the second polymer solution coats the polymer mesh, it may form a sealed layer completely surrounding the surface of the mesh, or it may form a partially sealed layer.

[0027] According to a particular embodiment of the invention, the method of the invention additionally comprises the step of performing a plasma surface treatment on the polymer mesh between said steps (a) and (b).

[0028] According to the present invention, when a polymer mesh is prepared using a hydrophobic polymer such as PCL (polycaprolactone) as the first polymer, a pretreatment process is carried out to impart hydrophilicity to the hydrophobic mesh, which allows a biocompatible hydrophilic second polymer to be uniformly coated. When plasma discharge is applied to the surface of a polymer material, gaseous reactive species are formed and react with the polymer surface layer, and the bonds of the constituent elements are broken by energy transfer, increasing the hydrophilicity of the surface.

[0029] Specifically, the plasma treatment can be performed under medium vacuum conditions of 1.0 to 0.1 Torr at room temperature.

[0030] Specifically, the plasma surface treatment is performed for 45-90 seconds, more specifically for 50-80 seconds, and most specifically for 50-70 seconds.

[0031] As shown in the examples described below, when plasma treatment is performed for 45 seconds or longer, the hydrophilicity of the surface is increased and a uniform collagen membrane is formed with almost no air bubbles generated on the mesh surface, but when the treatment is performed for more than 90 seconds, there is a disadvantage that the molecular weight of the first polymer decreases from the surface, weakening the mechanical strength.

[0032] According to another aspect of the invention, the present invention provides a porous scaffold comprising: (a) 0.1-0.5 mm 2 a first polymer mesh having pores and a strand diameter of 0.1-0.3 mm; and (b) a second biocompatible polymer coated on the surface of the first polymer mesh.

[0033] The first polymer and the second polymer used in the present invention have already been described above, and therefore their description will be omitted to avoid excessive duplication.

[0034] According to yet another aspect of the present invention, there is provided a support composition for human implantation comprising the porous scaffold composition.

[0035] As used herein, the term "transplantation" refers to the process of transferring living tissues, cells, or artificial supports that receive them from a donor to a recipient with the goal of maintaining the functional integrity of the transplanted tissue or cells in the recipient. Thus, the term "transplant support" refers to a physical support used in the process of transferring living tissues or cells to a recipient.

[0036] According to a specific embodiment of the present invention, the support composition of the present invention is a support for use in ligament reconstruction, craniofacial reconstruction, maxillofacial reconstruction, tissue reconstruction following removal of melanoma or head and neck cancer, chest wall reconstruction, delayed burn reconstruction, pelvic reinforcement, genitalia reinforcement or abdominal wall reinforcement, more specifically a support for use in ligament reconstruction or abdominal wall reinforcement.

[0037] According to yet another aspect of the present invention, there is provided a method for tissue reconstruction, comprising the step of implanting the above-mentioned support composition of the present invention into a living body.

[0038] According to yet another aspect of the present invention, there is provided a method for producing a dual-structure porous scaffold comprising embossing a biocompatible first polymer into a mesh shape on a surface of a support containing a biocompatible second polymer.

[0039] The present inventors have discovered that when two biocompatible polymers having different structures and functions are fabricated into a three-dimensional porous structure and a two-dimensional porous structure, respectively, and then bonded together, significantly improved physical properties are exhibited while maintaining the inherent functions of each polymer, such as tissue regeneration, wound healing, and providing bonding strength, thereby completing the present invention.

[0040] The second polymer used in the present invention has already been described above, and therefore will not be described again to avoid excessive duplication. The biocompatible second polymer of the present invention may be collagen, and in this case, the support containing the second polymer may be a collagen sponge.

[0041] In the present specification, the term "sponge" refers to a spongy porous material composed of a three-dimensional network of polymers linked by ionic or covalent bonds, with water as a dispersion medium. The collagen sponge of the present invention can be used without limitation as long as it is a spongy structure having cavities or voids in collagen, and for example, it can be produced by freeze-drying a collagen solution or dispersion, or various commercially available collagen sponges can be purchased and used.

[0042] The first polymer used in the present invention has already been described above, and therefore will not be described again to avoid excessive duplication. Specifically, the first polymer of the present invention may be PCL (polycaprolactone).

[0043] The double-structured porous scaffold of the present invention can be fabricated as a collagen sponge-PCL mesh composite by embossing a first polymer, e.g., PCL, into a mesh shape on the surface of a support, e.g., collagen sponge, containing a second polymer. As used herein, the term "emboss" refers to the process of bonding a PCL polymer to the surface of the collagen sponge so that the mesh shape is imprinted on the surface of the sponge.

[0044] According to a specific embodiment of the present invention, the embossing is performed by using a 3D printer to print the first polymer in a mesh shape on the surface of a support containing the second polymer.

[0045] According to a specific embodiment of the present invention, the mesh shape has a diameter of each strand of 0.3-0.5 mm and a spacing between the strands of 0.1-0.3 mm.

[0046] According to yet another aspect of the present invention, there is provided a dual-structure porous scaffold comprising: (a) a support containing a biocompatible second polymer; and (b) a first biocompatible polymer mesh bonded to a surface of the support;

[0047] The first polymer, the second polymer, the support, and the process of bonding the first polymer mesh to the second polymer-containing support using embossing or the like used in the present invention have already been described above, so their description will be omitted to avoid excessive duplication.

[0048] The double-structured porous scaffold of the present invention (e.g., a collagen sponge-PCL mesh composite) has superior tensile strength and bonding strength compared to typical collagen sponges used in regenerative treatments of bone tissue, skin tissue, etc., and also has biodegradable properties, providing more stable bonding function and significantly improved fixation function in the human body for the period required for wound healing and tissue regeneration. Effect of the Invention

[0049] The features and advantages of the present invention can be summarized as follows: (a) The present invention provides a porous scaffold with excellent tissue engineering properties and a method for producing the same.

[0050] (b) The scaffold of the present invention can be produced by a simple process, and has high tensile strength and biocompatibility, as well as a remarkably excellent cell engraftment rate. Therefore, the scaffold can be usefully used as a support composition for human transplantation in various applications, including artificial ligaments and supports for abdominal wall reinforcement. [Brief description of the drawings]

[0051] [Figure 1] FIG. 1 shows the results of optical microscopic observation of the polymer mesh of the present invention produced using a 3D printer.

[0052] [Diagram 2] FIG. 2 is an optical photograph showing the surface bubbles that were generated after the polymer mesh of the present invention was subjected to plasma surface treatment for various times and then coated with collagen.

[0053] [Diagram 3] FIG. 3 shows the macroscopic shape of the collagen-coated mesh.

[0054] [Figure 4] FIG. 4 is an electron micrograph showing the microstructure of the collagen-coated mesh.

[0055] [Diagram 5] FIG. 5 shows the results of an analysis of the physical strength of collagen-coated transplant mesh and acellular allogeneic dermis.

[0056] [Figure 6] Figure 6 shows the results of analysis of elements present on the surface of an uncoated collagen mesh (Figure 6A) and a mesh coated with 0.5% collagen (Figure 6B) using EDS (Energy Dispersive X-Ray Spectroscopy, EDAX, USA).

[0057] [Figure 7] FIG. 7 shows the results of observing (FIG. 7A) and quantifying (FIG. 7B) active cells after cell culture to compare the cell reactivity of meshes with and without collagen coating.

[0058] [Figure 8]To verify the biological safety of meshes with and without collagen coating, meshes were implanted into acellular allogeneic dermis and the dermis of experimental animals for 6, 12, and 20 weeks, respectively, and then tissue samples were collected and stained with Masson's trichrome (Fig. 8A). Based on this, the thickness of the membrane due to the inflammatory response (Fig. 8B) and the thickness of the implant due to biodegradation (Fig. 8C) were quantified.

[0059] [Figure 9] To verify the distribution and number of blood vessels (arterioles) inside the mesh with or without collagen coating, the mesh was implanted into acellular allogeneic dermis and animal dermis, and then immunofluorescence staining of the obtained tissue was performed (Figure 9A) and the number of blood vessels was quantified (Figure 9B).

[0060] [Figure 10] FIG. 10 is a photograph showing the macroscopic shape of a construct in which a single collagen sponge and a polymer mesh were directly printed and bonded together.

[0061] [Figure 11] FIG. 11 is an electron microscope photograph showing a microstructure in which a polymer mesh was printed and bonded onto a collagen sponge.

[0062] [Figure 12] FIG. 12 shows the results of analyzing the physical properties of a collagen sponge and a structure in which a polymer mesh is printed and bonded to the collagen sponge. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0063] The present invention will be described in more detail with reference to the following examples. It will be obvious to those skilled in the art that these examples are merely intended to more specifically illustrate the present invention, and that the scope of the present invention is not limited to these examples according to the gist of the present invention.

[0064] Working Example Example 1: Fabrication of a biodegradable polymer mesh 1-1. Preparation of polymer mesh A 3D printer (Biobots, USA) was used to fabricate the 3D polymer structure, and the 3D printing method allows the mesh size to be easily adjusted depending on conditions such as nozzle diameter, temperature, discharge pressure, and nozzle movement speed. The inventors selected a mesh shape with a diameter of 0.2 mm for each strand and a distance between strands of 1.0 mm (Figure 1) as the design that would most stably support the damaged ligament and abdominal wall, and used polycaprolactone (Sigma Aldrich, USA) as the raw polymer.

[0065] To prepare the polymer mesh, the nozzle diameter was set to 0.1-0.5 mm, the nozzle temperature to 80-90°C, the discharge pressure to 50-100 psi, and the nozzle moving speed to 2-5 mm / s. The polycaprolactone mesh prepared under these conditions was processed into circular test pieces with a diameter of 1.5 cm by punching, and then washed with 70% ethanol for about 30 minutes to remove foreign matter, and then dried at room temperature for 2 hours.

[0066] 1-2. Collagen coating of polymer mesh To impart biocompatibility to polycaprolactone meshes manufactured by 3D printing, the mesh surface was coated with collagen. In order to uniformly coat collagen, the inventors introduced a pretreatment step to impart hydrophilicity to highly hydrophobic polycaprolactone by performing surface treatment using plasma before coating. First, atelocollagen (type 1, medical device grade, Dalim Tissen, Korea) extracted from porcine dermis was dissolved in 0.5M acetic acid at 4°C for 12 hours at a concentration of 0.5% to prepare a collagen solution.

[0067] Searching for the optimal plasma treatment time In order to select the optimal plasma treatment time for the most efficient collagen coating, the washed and dried mesh was placed on a slide glass and then treated with a plasma surface treatment machine (PDC-32G Plasma Cleaner, Harrick Plasma, USA) for 0, 15, 30, 45, and 60 seconds under a medium vacuum condition of 1.0-0.1 Torr. After the surface treatment process, 250 μl of collagen solution was added per specimen, and the mesh surface was coated with collagen at 4 ° C for 30 minutes, which was observed under an optical microscope (EVOS (registered trademark) XL Core Cell Imaging System, Thermo Fisher Scientific, USA) (Figure 2). As shown in Figure 2, it was observed that the collagen-coated polycaprolactone mesh that was not plasma surface treated not only had an uneven collagen coating due to the strong hydrophobicity of the surface, but also had many bubbles on the mesh surface, and it was observed that the number of bubbles decreased as the plasma treatment time was gradually increased at 15 second intervals, and it was observed that a uniform collagen coating film was formed on the mesh surface when plasma was applied for 60 seconds.

[0068] Searching for the optimal collagen concentration The inventors then attempted to evaluate the optimal concentration of collagen to be coated on the surface, taking into consideration the physical properties and biocompatibility that the polymer mesh should have as an insert for the human body. To this end, atelocollagen was dissolved in 0.5M acetic acid at various concentrations (0.1, 0.5, 0.75, and 1.0%) for 12 hours at 4°C to prepare collagen solutions, which were then subjected to plasma surface treatment for 60 seconds, and 250 μl of each solution was placed on the mesh specimens, and coating was performed at 4°C for 30 minutes. After coating, each sample was cooled to -70°C for 12 hours, and then dried using a freeze dryer (FreeZone 12 plus, Labconco, USA) for 24 hours to create a porous surface structure of the collagen coated on the surface. Then, a neutralization process was performed to remove the acetic acid present in the form of a salt inside the freeze-dried collagen. For this purpose, the freeze-dried specimens were washed four times for 15 minutes using absolute alcohol (Ethanol Absolute, Merck KGaA, Germany), and then neutralized with acetic acid four times for 15 minutes using 0.5M NaOH (Duksan General Science, Korea) dissolved in 70% ethanol. Thereafter, to remove the residual NaOH in the specimens, they were washed four times for 15 minutes using 50% and 30% ethanol and triple distilled water. After washing, the collagen-coated meshes were cooled to -70°C for 12 hours, and then dried using a freeze dryer for 24 hours as mentioned above, and images were taken using a digital camera (EOS 500D, Canon, Japan) (Figure 3). As a result, in the groups coated with collagen at a concentration of 0.5% or more, a macroscopic shape was observed in which collagen was layered in a sponge shape on the mesh surface and blocked the pores, and this phenomenon became more severe as the collagen concentration increased, but the mesh shape was maintained at a collagen concentration of 0.1%. Next, to observe the microstructure of the collagen-coated mesh, the surface shape of the collagen-coated mesh was observed using an electron microscope (FE-SEM, MERLIN, Zeiss, Germany) (Figure 4). As a result, it was confirmed that in the case of the mesh without collagen coating, each polycaprolactone strand had a diameter of about 200 μm as originally designed. In the collagen-coated test specimens, as the collagen concentration increased from 0.1 to 0.75%, the collagen pores formed by freeze-drying decreased from about 500 μm to 20 μm, but in the test specimen coated with 1.0% collagen, the mesh surface was entirely covered with collagen and no pores could be observed. The porous structure of collagen thus formed is a useful structure for initial cell attachment and blood vessel formation inside the mesh when inserted into the human body, and the 0.5% collagen-coated mesh, which was confirmed to have pores of about 150 to 300 μm based on the results of surface observation using an electron microscope, was determined to be the most suitable mesh for biodegradable transplantation.

[0069] Example 2: Characterization of biodegradable mesh 2-1. Physical strength analysis of biodegradable mesh The tensile strength was measured to analyze the physical strength of the biodegradable transplant mesh prepared in the present invention. In order to obtain more reliable analysis results, a commercially available acellular allogeneic dermis (CG Derm, Korea) for reconstruction of soft tissues in the human body was used as a comparison group to compare the strength of the transplant mesh developed by the researchers in this study. For this purpose, each specimen was cut into a 1 cm x 5 cm rectangular shape, soaked in saline for 30 minutes, and then the tensile strength was measured by pulling the specimen at a speed of 1 mm per second using a universal test analyzer (Universal Testing Systems, Instron 3360, USA). As a result, the commercially available acellular allogeneic dermis showed a lower elasticity than the transplant mesh of the present invention until it reached a tensile rate of 50%, but at the point where it reached a tensile rate of 124%, it reached the highest tensile strength of 15.27 MPa (Figure 5). In contrast, it was confirmed that the transplant mesh of the present invention had a remarkably excellent elastic recovery force, with the elastic modulus and tensile modulus being two and five times higher than the acellular allogeneic dermis, respectively. Such high elastic recovery of the implant mesh of the present invention indicates that it has excellent properties as a body implant to provide physical reinforcement to ligaments, abdominal wall sites, and the like.

[0070] 2-2. Qualitative analysis of biodegradable mesh The elements present on the surface of the implant mesh of the present invention, depending on whether it was coated with collagen, were analyzed using EDS (Energy Dispersive X-Ray Spectroscopy, EDAX, USA). As a result, only carbon and oxygen components were detected in the polycaprolactone mesh without collagen coating, whereas nitrogen in peptides was detected in the collagen-coated test piece, confirming that nitrogen elements accounted for 12.71% of the total elements (Figure 6).

[0071] 2-3. Cellular reactivity of biodegradable mesh To evaluate the reactivity of cells with collagen-coated implant mesh in an in vitro environment, human dermal fibroblasts (Human dermal fibroblast, LONZA, USA) were cultured on the mesh surface. The circular specimens with a diameter of 1.5 cm were placed in a 24-well tissue culture plate (TCP, Corning, USA), which was then sterilized under a UV lamp for 30 minutes with 70% ethanol. Then, 50,000 fibroblasts (passage 4) were seeded on each specimen, and cells were also seeded on the TCP as a comparison group. The fibroblasts were then cultured for 7 days at 37°C under 5% carbon dioxide conditions in a medium containing DMEM (Dulbecco's Modified Eagle Medium, low glucose, Gibco, USA) mixed with 10v / v% FBS (Fetal bovine serum, Gibco, USA) and 1v / v% antibiotic (Gibco, USA). At this time, to analyze the behavior of the cells, a live and dead assay (Thermo Fisher Scientific, USA) was performed on the first and seventh days after the start of the culture to compare and analyze the survival / proliferation behavior of the cells. For this purpose, at the end of the culture, each test piece was washed three times with phosphate buffer solution (PBS, Gibco, USA), and calcein AM and EthD-1 (Ethidium homodimer-1) in the live and dead assay kit were diluted to concentrations of 2μ and 4μ, respectively, and placed in each test piece. After staining at room temperature for 30 minutes, the stained cells were observed using a confocal fluorescence microscope (LSM700, Zeiss, Germany) (Figure 7A) and quantitatively analyzed (Figure 7B). As a result, on the first day of culture, all three groups of test pieces had a cell density per unit area (1 mm 2) 20-30 highly active cells were observed, and there seemed to be no difference between the groups. However, by the seventh day of culture, the collagen-coated implant mesh group had 7 times as many cells per unit area as the non-collagen-coated group and 3 times as many as the TCP group, and clear cell reactions could be observed depending on the presence or absence of collagen. This is thought to be the result of the porous collagen structure present between the meshes providing sufficient space for cells to attach and grow. This shows that when the scaffold of the present invention is inserted into the human body after tissue incision, it is possible to efficiently induce the attachment of various cells including early fibroblasts and the formation of blood vessels inside the mesh.

[0072] Example 3: Biological safety of biodegradable mesh 3-1. Inflammatory response and biodegradation behavior of biodegradable mesh To evaluate the inflammatory response and biodegradation behavior of the implant mesh of the present invention with or without collagen coating, the mesh was implanted together with acellular allogeneic dermis (thickness: 1.5 mm, MegaDerm, L&C Bio, Korea) into the back skin of SD (Sprague Dawley) rats (6 weeks old, male N=4, Orient Bio, Korea), and the rats were euthanized at 6, 12, and 20 weeks to collect tissues, which were stained with Masson's trichrome (Sigma Aldrich, USA), and the cross-sections of the tissues were observed under an optical microscope (CX43, Olympus, Tokyo, Japan) (Figure 8A). Inflammatory response around the implant (Figure 8B) and the degree of biodegradation of the implant (Figure 8C) were also analyzed.

[0073] As shown in Figure 8A, in normal tissue, the boundaries between the epidermis, dermis, and subcutaneous tissue of the skin were clearly observed over a 20-week period, and in the groups in which mesh and acellular allogeneic dermis were inserted, it was observed that the graft was inserted under the dermal tissue without any movement. However, unlike acellular allogeneic dermis, it was confirmed that tissue filled the gaps between the porous structure formed by the mesh in all groups in which mesh was inserted, but in the acellular allogeneic dermis, a thick membrane was formed due to an excessive inflammatory response at 20 weeks, and delamination with the tissue was observed.

[0074] To analyze the inflammatory response observed above, the thickness of the membrane formed around the graft was measured (Figure 8B). As a result, it was observed that at 6 weeks after grafting, a membrane of about 250 μm was formed with the acellular allogeneic dermis, similar to that of the collagen-coated mesh, and by 12 weeks, similar values ​​were observed with membranes of 200 to 280 μm formed in all graft groups. However, at 20 weeks, it was confirmed that the acellular allogeneic dermis group had formed a thick membrane of about 340 μm, while the mesh group maintained a membrane of 250 μm, similar to that at 6 weeks, regardless of whether collagen was coated or not. It can be inferred that these results are due to an excessive inflammatory response, based on the results of Masson's Trichrome staining photographs observed at 20 weeks. Next, to compare the biodegradation behavior of the grafts, the change in thickness of each graft was measured over a 20-week period (Figure 8C). At 6 weeks, 99% of the acellular allogeneic dermis remained, which is close to the thickness of the graft when it was first inserted, while about 78% remained in the mesh group, confirming a thickness reduction of about 22%. This trend was maintained until the 12th week, with the acellular allogeneic dermis maintaining a thickness of 92% and the mesh maintaining a thickness of about 70%. However, at the 20th week, the acellular allogeneic dermis experienced a rapid decrease in thickness compared to the 12th week, remaining at only about 45% of its initial thickness, confirming rapid biodegradation at 8 weeks. This result, as shown in Figure 8B, is thought to be due to the formation of the thickest membrane around the graft due to an inflammatory response caused by the rapid biodegradation of the acellular allogeneic dermis inserted into the tissue at the 20th week.

[0075] 3-2. Vascular formation inside biodegradable mesh To evaluate the angiogenesis induction ability of the transplant mesh of the present invention with or without collagen coating, immunostaining was performed on the previously collected tissues. Cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI, Blue signal, Sigma Aldrich, USA) and vascular endothelial cells were stained with CD31 (Red signal, Thermo Fisher Scientific, Waltham, MA, USA). Then, the tissues were observed using a confocal microscope (LSM700, Carl Zeiss, Oberkochen, Germany) (Figure 9A), and the number of blood vessels (arterioles) per area was comparatively quantified (Figure 9B).

[0076] As shown in the fluorescence micrograph in Figure 9A, uneven blood vessel distribution was observed within the acellular allogeneic dermis at weeks 12 and 20, whereas the tissue with the mesh inserted had blood vessels distributed uniformly even inside the mesh, regardless of whether collagen coating was used or not. This phenomenon can be inferred from the fact that the acellular allogeneic dermis rapidly decomposed and reduced in thickness at week 20, resulting in a reduced cross-sectional area and localized blood vessel distribution.

[0077] The arterioles of SD rats are known to have a diameter of 20 to 40 μm, and the unit area (mm 2 ) was quantified, and the number of blood vessels that met the arteriole diameter criteria was quantified (Figure 9B). At 12 weeks after implantation, a similar number of blood vessels, approximately 16, was observed in the acellular allogeneic dermis and mesh, whereas the collagen-coated mesh had approximately 23 blood vessels, 40% more than the acellular allogeneic dermis. This trend was maintained up to the 20th week, confirming that the collagen coating actively induces angiogenesis inside the mesh.

[0078] Example 4: Preparation and characterization of collagen sponge-polymer mesh composites 4-1. Preparation of collagen sponge In yet another embodiment of the present invention, the present inventors prepared a collagen-containing sponge bonded with a polymer mesh by dissolving atelocollagen (type 1, medical device grade, Dalim Tissen, Korea) extracted from pig dermis in 0.5M acetic acid at a concentration of 3.0 wt%. The mixture was then placed in a brass mold, frozen in liquid nitrogen (-196°C), and freeze-dried for 24 hours as described above in Example 1. The dried collagen sponge was then subjected to dehydrothermal treatment (DHT) in an oven at 120°C for 24 hours to produce a collagen sponge (FIG. 10).

[0079] 4-2. Fabrication of collagen sponge-polymer mesh composites by 3D printing To enhance the physical properties of the collagen sponge, the sponge was fixed to a 3D printing stage, and a mesh shape with a diameter of each strand of 0.4 mm and a distance between strands of 2.0 mm was directly printed on the sponge under the printing conditions applied to prepare the polymer mesh in Example 1 to prepare a PCL-collagen conjugate (Figure 10).

[0080] Next, the surface and cross-sectional shapes of the mesh structure bonded to the collagen sponge by 3D printing were observed using an electron microscope (Figure 11). As a result, pores of 20-200 μm were formed on the surface of the collagen sponge, and cross-sectional observation confirmed that a structure had been formed in which the printed PCL and collagen were stably bonded.

[0081] 4-3. Physical strength analysis of collagen sponge-polymer mesh joint In order to compare and analyze the physical strength of the collagen sponge-polymer mesh composite, the tensile strength and the bond strength with the suture used for fixing the human body were measured. As is clear from the tensile strength measurement results in FIG. 12, the composite of the present invention to which the PCL mesh is bonded has a tensile strength about 20 times higher, a tensile modulus about 70 times higher, and an elastic modulus about 10 times higher than that of the plain collagen sponge. Next, a suture was passed through each of the collagen sponge and the collagen sponge-PCL mesh composite of the present invention to analyze the bond strength with the suture. As a result, it was observed that the strength was significantly increased from about 56.26 KPa to 496.15 KPa compared to the plain collagen sponge. From these results, it was confirmed that the composite of the present invention in which the PCL polymer is bonded to the collagen sponge provides a stable fixation function in the human body while dramatically improving the physical properties of the existing collagen sponge.

[0082] Although certain parts of the present invention have been described in detail above, it is obvious to those skilled in the art that such specific descriptions are merely preferred embodiments and therefore do not limit the scope of the present invention. Therefore, the substantial scope of the present invention is defined by the appended claims and their equivalents.

Claims

1. A method for producing a double-structured porous scaffold, comprising embossing a biocompatible polymer into a mesh shape on a surface of a collagen sponge, Here, the embossing is performed by outputting the polymer into a mesh shape on the surface of the collagen sponge using a 3D printer.

2. 2. The method of claim 1, wherein the polymer is selected from the group consisting of polycaprolactone (PCL), poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), poly(L-lactic acid-co-ε-caprolactone) (LCL), and combinations thereof.

3. The method of claim 2 , wherein the polymer is polycaprolactone (PCL).

4. The method of claim 1 , wherein the mesh configuration comprises strands each having a diameter of 0.3-0.5 mm and an inter-strand spacing of 0.1-0.3 mm.

Images