KR102364686B1 - A Novel Porous Scaffold and Methods for Preparing the Same - Google Patents

Abstract

The present invention relates to a porous scaffold having excellent tissue engineering properties and a method for manufacturing the same. The scaffold of the present invention can be manufactured by a simple process, and has high tensile strength and biocompatibility as well as a remarkably excellent cell engraftment rate. can be used

Description

A Novel Porous Scaffold and Methods for Preparing the Same

The present invention relates to a biocompatible porous scaffold, a support composition for human transplantation comprising the same, and a method for preparing the same.

Recently, as the fields of bioengineering, material engineering, and surgical operation have been greatly developed, tissue engineering for the purpose of replacing and regenerating lost body tissues is making remarkable progress. Tissue engineering is the fusion of life science, engineering, and medicine to understand the correlation between the structure and function of living tissue, and based on this, an artificial implant that can be transplanted into the body to replace or regenerate damaged tissue or organs with normal tissue. It aims to maintain, improve, or restore bodily functions through tissues.

The loss of body tissues is due to various causes such as degenerative diseases, trauma, surgical removal of tumors, and certain congenital malformations, which can be restored only through the regeneration of irreversibly lost tissues. In order to induce regeneration of lost tissue through tissue engineering, it is important to first prepare a biodegradable polymer scaffold (scaffold) similar to living tissue. The main requirement of the scaffold material used for the regeneration of human tissue is that it should sufficiently play the role of a substrate or frame so that the tissue cells can adhere to the material surface to form a tissue with a three-dimensional structure, and the transplanted cells and the patient It should also be able to act as an intermediate barrier located between cells.

After the scaffold is implanted into a subject, engraftment of cells necessary for tissue regeneration is induced and the formation of a new tissue is initiated, and the newly formed tissue should fill the space as it disappears over time. Therefore, the scaffold is preferably biodegradable that does not require surgical removal, does not undergo immunorejection, inflammatory response, or long-term fibrous encapsulation, does not undergo shrinkage of the graft volume, and is free from serious complications such as prosthetic implants. do.

Therefore, after serving as a physical support for an appropriate period, the scaffold must have a certain level of mechanical strength and elasticity along with biodegradability to efficiently induce tissue reformation while disappearing naturally. Choosing the most suitable structure is a very important issue.

Numerous papers and patent documents are referenced throughout this specification and their citations are indicated. The disclosures of the cited papers and patent documents are incorporated herein by reference in their entirety to more clearly describe the level of the technical field to which the present invention pertains and the content of the present invention.

Patent Literature 1. Korean Application No. 10-2018-7032731

The present inventors have made intensive research efforts to develop a scaffold for efficient tissue regeneration that can be manufactured in a relatively simple process while having sufficient physical strength and excellent biocompatibility. As a result, a mesh-type support composed of strands of a constant diameter while having pores of a certain size with the first polymer is manufactured, and then the surface thereof is coated with a second polymer having biocompatibility as a different polymer from the first polymer. In this case, it has been found that it can be used as a scaffold for human transplantation for various purposes, including artificial ligaments and supports for reinforcing the abdominal wall, because of its high tensile strength and biocompatibility as well as a remarkably excellent cell engraftment rate.

In addition, when two biocompatible polymers with different structures and functions are prepared into a three-dimensional porous structure and a two-dimensional porous structure, respectively, and then joined, the unique functions such as tissue regeneration, wound healing, and in vivo bonding are maintained. However, by discovering the fact that significantly improved physical properties are exhibited, the present invention has been completed.

Accordingly, it is an object of the present invention to provide a porous scaffold and a method for manufacturing the same.

Another object of the present invention is to provide a support for human transplantation comprising the porous scaffold.

Other objects and advantages of the present invention will become more apparent from the following detailed description of the invention, claims and drawings.

According to one aspect of the present invention, there is provided a method for manufacturing a porous scaffold comprising the steps of:

(a) producing a polymer mesh (mesh) having a diameter of 0.1-0.3 mm of the strands having pores of 0.1 - 0.5 mm ² from the first polymer solution;

(b) coating the surface of the resulting polymer mesh with a second polymer solution having biocompatibility.

The present inventors have made intensive research efforts to develop a scaffold for efficient tissue regeneration that can be manufactured in a relatively simple process while having sufficient physical strength and excellent biocompatibility. As a result, a mesh (mesh) made of strands having a diameter of 0.1 - 0.3 mm while having pores of 0.1 - 0.5 mm ² was prepared as the first polymer, and the surface of the second polymer was different from that of the first polymer and had biocompatibility. When coated with a polymer, it was found that it can be used as a scaffold for human transplantation for various purposes, including artificial ligaments and supports for reinforcing the abdominal wall, as it exhibits high tensile strength and biocompatibility as well as a remarkably excellent cell engraftment rate.

As used herein, the term “scaffold” refers to tissue engineering for accelerating the recovery and regeneration of damaged tissues by placing living cells, specifically, cells derived from damaged tissues or cells involved in the recovery of damaged tissues. engineering) structure. The term “cell attachment” refers to the direct or indirect adsorption of cells to a matrix or other cells while maintaining their intrinsic biological activity. Specifically, the scaffold of the present invention may have a planar structure consisting of a single mesh or a three-dimensional structure in which a plurality of meshes are stacked.

As used herein, the term “polymer” refers to a synthetic or natural high molecular compound in which the same or different types of monomers are continuously combined. Thus, polymers include homopolymers (polymers in which one type of monomer is polymerized) and interpolymers prepared by the polymerization of at least two different monomers, and interpolymers include copolymers (polymers prepared from two different monomers). polymers) and polymers prepared from more than two different monomers.

According to a specific embodiment of the present invention, the first polymer used in the present invention is PCL (polycaprolactone), PLLA (poly (L-lactic acid)), PGA (poly (glycolic acid)), PLGA (poly (lactic-co) -glycolic acid)), PLCL (poly(L-Lactide-co-ε-caprolactone)), and combinations thereof.

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

According to the present invention, the first polymer of the present invention forms a mesh having pores of a constant size while intersecting the strands at regular intervals having a constant thickness. The pores should have the most suitable size in terms of the seating, proliferation and activity of cells, induction of new blood vessels during tissue regeneration, and mechanical strength and elasticity of the mesh itself so that recovery and regeneration of damaged tissue, which is the ultimate goal of the present invention, is efficient. can be achieved with Accordingly, a suitable pore area is specifically 0.1-0.5 mm ² , more specifically 0.1-0.4 mm ² , still more specifically 0.2-0.3 mm ² , and most specifically about 0.25 mm ² .

As used herein, the term “pore area” refers to the average area of repeated pores appearing through the crossing of strands in the mesh structure of the present invention made of the first polymer, and this area is a coating using a second polymer solution to be described later. Means the area measured before performing

In addition, in order to have an appropriate elastic modulus and tensile modulus to secure physical properties suitable as a support for human transplantation, the diameter of the strand forming the mesh together with the above-described pore area is important. Accordingly, the diameter of a suitable strand is specifically 0.1 - 0.3 mm, more specifically 0.15 - 0.25 mm, and most specifically about 0.2 mm.

The step of generating the polymer mesh from the first polymer solution of the present invention may use various methods known in the art, for example, 3D printing method, salt leaching method (solvent-casting particulate leaching), salt foaming method (gas foaming) ), fiber meshes/fiber bonding, phase separation, melt moulding, freeze drying and electrospinning, including, but not limited to it is not

According to the present invention, the scaffold of the present invention imparts biocompatibility in addition to the above-described mechanical strength by coating a polymer mesh prepared with a first polymer solution with a second polymer solution having biocompatibility.

As used herein, the term “biocompatibility” refers to a property that does not cause short-term or long-term side effects when administered in vivo and in contact with cells, tissues or body fluids of organs, specifically, causes tissue necrosis by contact with living tissues or blood. This includes not only tissue compatibility and blood compatibility that does not coagulate blood, but also biodegradability that disappears after a certain period of time after in vivo administration.

As used herein, the term “biodegradability” refers to the property of being naturally decomposed when exposed to a physiological solution of pH 6-8, specifically, the passage of time by body fluids, degrading enzymes or microorganisms in a living body. It means a property that can be decomposed according to The biodegradable polymer usable in the present invention may be any synthetic or natural polymer as long as it has the above-described biodegradability, for example, collagen, gelatin, chitosan, hyaluronic acid, poly(valerolactone), poly(hydroxy butylate), poly(hydroxyvalerate) collagen, gelatin, chitosan, hyaluronic acid, and combinations thereof.

According to a specific embodiment of the present invention, the second polymer having biocompatibility is a natural polymer, more specifically, collagen, and most specifically, type I collagen.

According to a more specific embodiment of the present invention, the collagen solution is used at a concentration of 0.2 - 0.8% (v/v), more specifically, at a concentration of 0.3-0.7% (v/v), most specifically is used at a concentration of 0.4-0.6% (v/v).

As used herein, the term “coating” refers to forming a new layer of a certain thickness by modifying a specific material on the target surface, and the target surface and the coating material may be modified through ionic or non-covalent bonding. The term "non-covalent bond" refers to physical bonds such as adsorption, cohesion, entanglement and entrapment, as well as interactions such as hydrogen bonds and van der Waals bonds alone or in the above physical bonds. It is a concept that includes a bond generated by acting with a bond. In the present invention, when the second polymer solution coats the polymer mesh, a sealed layer may be formed while completely surrounding the surface of the mesh, or a partially sealed layer may be formed.

According to a specific embodiment of the present invention, the method of the present invention further comprises the step of performing plasma surface treatment on the polymer mesh between the step (a) and the step (b).

According to the present invention, when a polymer mesh is manufactured using a hydrophobic polymer such as PCL (polycaprolactone) as the first polymer, the hydrophilic second polymer having biocompatibility is homogeneously coated by going through a pretreatment process that gives hydrophilicity to the hydrophobic mesh. can do. When plasma discharge is applied to the surface of the polymer material, gas-reactive species are formed, and the hydrophilicity of the surface is increased through reaction with the polymer surface layer and cleavage of elemental bonds through energy transfer.

Specifically, plasma treatment can be performed under medium vacuum conditions of 1.0-0.1 Torr at room temperature.

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

As will be seen in Examples to be described later, when plasma treatment is performed for 45 seconds or more, the hydrophilicity of the surface is increased, so that a uniform collagen film is formed without almost generating bubbles on the surface of the mesh. There is a disadvantage in that the molecular weight is reduced and the mechanical strength is weakened.

According to another aspect of the present invention, there is provided a porous scaffold comprising:

(a) a first polymer mesh having pores of 0.1 - 0.5 mm ² and a strand diameter of 0.1 - 0.3 mm; and

(b) a second polymer having a biocompatibility coated on the surface of the first polymer mesh.

Since the first polymer and the second polymer used in the present invention have already been described above, their description is omitted to avoid excessive overlap.

According to another aspect of the present invention, there is provided a support composition for human transplantation comprising the porous scaffold composition.

As used herein, the term “transplantation” refers to a process of delivering living tissue, cells, or an artificial scaffold for receiving the transplanted tissue or cells from a donor to a recipient for the purpose of maintaining the functional integrity of the transplanted tissue or cells. Accordingly, the term “support for transplantation” refers to a physical support used in the process of delivering living tissue or cells to a recipient.

According to a specific embodiment of the present invention, the scaffold composition of the present invention can be used for ligament reconstruction, craniofacial reconstruction, maxillofacial reconstruction, tissue reconstruction after removal of melanoma or head and neck cancer, chest wall reconstruction, delayed burn reconstruction, pelvic reinforcement, genital reinforcement, or It is a support used for reinforcing the abdominal wall, and more specifically, it is a support used for ligament reconstruction or reinforcement of the abdominal wall.

According to another aspect of the present invention, the present invention provides a tissue reconstruction method comprising the step of implanting the above-described scaffold composition of the present invention in vivo.

According to another aspect of the present invention, the present invention provides a dual structure comprising the step of embossing a first polymer having biocompatibility into a mesh shape on the surface of a support containing a second polymer having biocompatibility. A method for manufacturing a porous scaffold is provided.

The present inventors have prepared two biocompatible polymers having different structures and functions into a three-dimensional porous structure and a two-dimensional porous structure, respectively, and then, when they are joined, the unique functions of each polymer, such as tissue regeneration, wound healing, and provision of bonding strength By discovering the fact that remarkably improved physical properties are exhibited while maintaining the , the present invention has been completed.

Since the second polymer used in the present invention has already been described above, the description thereof is omitted to avoid excessive overlap. The second polymer of the present invention having biocompatibility may be collagen, and in this case, the support containing the second polymer may be a collagen sponge.

As used herein, the term “sponge” refers to a spongy porous material composed of a three-dimensional network connected by ionic or covalent bonds of polymers and using water as a dispersion medium. The collagen sponge of the present invention may be used without limitation as long as it is a spongy structure having voids or pores in collagen, for example, it may be prepared by freeze-drying a collagen solution or dispersion, or various commercially available ready-made collagen sponges may be purchased and used. there is.

Since the first polymer used in the present invention has also been described above, the description thereof is omitted to avoid excessive duplication. Specifically, the first polymer of the present invention may be PCL (polycaprolactone).

The dual structure porous scaffold of the present invention can be fabricated as a conjugate of a collagen sponge-PCL mesh by embossing a first polymer, for example PCL, into a mesh form on a support containing a second polymer, for example, a collagen sponge surface. there is. As used herein, the term “embossing” refers to a process of bonding PCL polymer to the sponge surface so that a mesh shape is engraved on the surface of the collagen sponge in a protruding form.

According to a specific embodiment of the present invention, the embossing may be performed by outputting the first polymer in a mesh shape using a 3D printer on the surface of the support containing the second polymer.

According to a specific embodiment of the present invention, in the mesh form, the diameter of each strand is 0.3-0.5 mm and the spacing between the strands is 0.1-0.3 mm.

According to another aspect of the present invention, there is provided a bi-structured porous scaffold comprising:

(a) a support containing a second polymer having biocompatibility; and

(b) a first polymer mesh bonded to the surface of the support and having biocompatibility.

the first polymer used in the present invention; a second polymer; support; And since the process of bonding the first polymer mesh to the second polymer-containing support using embossing or the like has already been described above, the description thereof will be omitted to avoid excessive overlap.

The dual structure porous scaffold of the present invention (for example, a collagen sponge-PCL mesh conjugate) has superior tensile and bonding strength and biodegradable properties compared to general collagen sponges used for regenerative treatment of bone and skin tissues, etc. It provides a more stable bonding function and a significantly improved fixation function in the human body for the period required for wound healing and tissue regeneration.

The features and advantages of the present invention are summarized as follows:

(a) The present invention provides a porous scaffold having excellent tissue engineering properties and a method for manufacturing the same.

(b) The scaffold of the present invention can be manufactured by a simple process and exhibits high tensile strength and biocompatibility as well as a remarkably excellent cell engraftment rate. can be usefully used as

1 shows the results of observing the polymer mesh of the present invention manufactured using a 3d printer with an optical microscope.
2 is an optical photograph showing the surface bubbles generated after the polymer mesh of the present invention is subjected to plasma surface treatment for various times and then coated with collagen.
3 is a diagram showing the macroscopic shape of the collagen-coated mesh.
4 is an electron micrograph showing the microscopic shape of the collagen-coated mesh.
5 is a diagram showing the results of analyzing the physical strength of the collagen-coated graft mesh and acellular allogeneic dermis.
6 is an analysis of the elements present on the surface of the collagen uncoated mesh (FIG. 6a) and the 0.5% collagen-coated mesh (FIG. 6b) using EDS (Energy Dispersive X-Ray Spectroscopy, EDAX, USA). Each result is shown.
7 shows the results of observing active cells after cell culture ( FIG. 7a ) and quantifying them ( FIG. 7b ), respectively, in order to compare the cell reactivity of the mesh according to the collagen coating.
Figure 8 shows the results of staining with Mason's trichrome after transplanting the mesh into the acellular allogeneic dermis and the dermis of experimental animals for 6, 12, and 20 weeks, respectively, in order to verify the biological safety of the mesh according to whether or not collagen is coated (Fig. 8a), and the results of quantification of the thickness of the film according to the inflammatory reaction (FIG. 8B) and the thickness of the implant according to the biodegradation (FIG. 8C) are shown, respectively.
9 is a result of immunofluorescence staining on the tissue obtained after transplanting the mesh into the acellular allogeneic dermis and the animal dermis in order to verify the distribution and number of blood vessels (arterioles) inside the mesh depending on whether or not collagen is coated (Fig. 9a) and the result of quantifying the number of blood vessels (FIG. 9b) are respectively shown.
10 is a photograph showing the macroscopic shape of a structure in which a single collagen sponge and a polymer mesh are directly printed and bonded.
11 is an electron micrograph showing a microstructure bonded to a polymer mesh printed on a collagen sponge.
12 is a diagram showing the results of analyzing the physical properties of a collagen sponge and a structure bonded thereto by printing a polymer mesh.

Hereinafter, the present invention will be described in more detail through examples. These examples are only for illustrating the present invention in more detail, and it will be apparent to those of ordinary skill in the art that the scope of the present invention is not limited by these examples according to the gist of the present invention. .

Example

Example 1: Preparation of biodegradable polymer mesh

1-1. Fabrication of polymer mesh

A 3D printer (Biobots, USA) was used to manufacture the 3D polymer structure, and the 3D printing technique can easily adjust the mesh size according to conditions such as nozzle diameter, temperature, discharge pressure, and nozzle movement speed. The present inventors selected a mesh shape with a diameter of each strand of 0.2 mm and a spacing of 1.0 mm between the strands as a design that can most stably support the damaged ligament and abdominal wall (Fig. 1), and as a raw material polymer, polycaprolactone (Polycaprolactone, sigma aldrich, USA) was used.

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

1-2. Collagen coating on polymer mesh

In order to provide biocompatibility of the polycaprolactone mesh prepared by 3D printing, the surface of the mesh was coated with collagen. The present inventors introduced a pretreatment process that imparts hydrophilicity to polycaprolactone with strong hydrophobicity before coating for a homogeneous coating of collagen by performing surface treatment using plasma. First, a collagen solution was prepared by dissolving atelo collagen (type 1, medical device grade, Dalim Thyssen, Korea) extracted from pig dermis in 0.5 M acetic acid at a concentration of 0.5% at 4°C for 12 hours.

Searching for optimal plasma processing time

In order to select the optimal plasma treatment time for the most efficient collagen coating, the mesh after washing and drying is placed on a slide glass and then subjected to a plasma surface treatment machine (PDC) for 0, 15, 30, 45, 60 seconds under medium vacuum conditions of 1.0-0.1 Torr. -32G Plasma Cleaner, Harrick Plasma, USA) was treated with polycaprolactone mesh. After the surface treatment process, 250 μl of collagen solution per specimen was put and collagen was coated on the mesh surface at 4° C. for 30 minutes, and this was observed with an optical microscope (EVOS®XL Core Cell Imaging System, Thermo Fisher scientific, USA) (FIG. 2). ). As shown in Fig. 2, it was observed that the collagen-coated polycaprolactone mesh without plasma surface treatment not only had an uneven collagen coating due to the strong hydrophobicity of the surface, but also generated many bubbles on the surface of the mesh, As the plasma treatment time was gradually increased at 15-second intervals, a tendency to decrease bubbles could be observed, and it could be observed that a uniform collagen coating film was formed on the mesh surface when plasma was applied for 60 seconds.

Exploration of optimal collagen concentration

Then, the present inventors tried 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 a human body insert. For this, a collagen solution was prepared by dissolving atelo collagen in various concentrations (0.1, 0.5, 0.75, and 1.0%) in 0.5 M acetic acid at 4° C. for 12 hours, and after plasma surface treatment for 60 seconds, it was applied to the mesh specimen. 250 μl of each was added and coating was carried out at 4° C. for 30 minutes. After the coating was completed, 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 collagen coated on the surface. . Thereafter, a neutralization operation was performed to remove the acetic acid present in the form of a salt in the freeze-dried collagen. For this, the freeze-dried specimen was washed 4 times for 15 minutes using anhydrous alcohol (Ethanol absolute, Merck KGaA, Germany), and 0.5 M NaOH (Duksan General Science, Korea) was dissolved in 70% ethanol. After that, acetic acid was neutralized 4 times for 15 minutes. Then, in order to remove the residual amount of NaOH present in the specimen, it was sequentially washed 4 times for 15 minutes using 50% and 30% ethanol and tertiary distilled water. After the washed collagen-coated mesh was cooled to -70°C for 12 hours, and dried using a freeze dryer for 24 hours as mentioned above, a digital camera (EOS 500D, Canon, Japan) was used. to obtain an image (FIG. 3). As a result, in the group coated with collagen at a concentration of 0.5% or more, a macroscopic shape blocking pores was observed while collagen was laminated on the mesh surface in the form of a sponge. This phenomenon worsened as the concentration of collagen increased, but 0.1% concentration of collagen, the shape of the mesh was completely preserved.

Next, the surface shape of the collagen-coated mesh was observed using an electron microscope (FE-SEM, MERLIN, Zeiss, Germany) to observe the microscopic shape of the collagen-coated mesh (FIG. 4). As a result, in the case of the mesh not coated with collagen, it was confirmed that each polycaprolactone strand had a diameter of about 200 μm as originally designed. In the collagen-coated specimens, as the collagen concentration increased to 0.1 - 0.75%, the collagen pores formed by freeze-drying decreased from about 500 μm to 20 μm. The stoma could not be observed. The porous structure of collagen thus formed is a useful structure for initial cell adhesion and blood vessel formation into the mesh when inserted into the human body. Judging from the results of surface observation using an electron microscope, 0.5 confirmed to have pores of about 150 to 300 μm % collagen-coated mesh was judged to be most suitable as a biodegradable mesh for transplantation.

Example 2: Characterization of biodegradable meshes

2-1. Physical strength analysis of biodegradable meshes

The tensile strength was measured to analyze the physical strength of the biodegradable graft mesh fabricated in the present invention. In order to secure a more reliable analysis result, ready-made acellular allogeneic dermis (CG Derm, Korea) was set as a comparison group for soft tissue reconstruction of the human body, and the strength was compared with the mesh for transplantation developed by this research team. For this, each specimen is processed into a 1 cm x 5 cm rectangle, soaked in physiological saline for 30 minutes, and then the tensile strength is measured by pulling the specimen at a speed of 1 mm per second using a universal testing system (Universal Testing Systems, Instron 3360, USA). did As a result, the ready-made acellular allogeneic dermis showed lower elasticity than the mesh for transplantation of the present invention until it showed a tensile rate of 50%, but showed the highest tensile strength of 15.27 MPa at the point of showing a tensile rate of 124%. (Fig. 5). On the other hand, the elastic modulus and tensile modulus of the mesh for transplantation of the present invention were 2 times and 5 times higher than those of acellular allogeneic dermis, respectively, and it was confirmed that the elastic restoring force was remarkably excellent. This high elastic restoring force of the mesh for implantation of the present invention shows that it has very excellent properties as a human body insert for providing physical reinforcement to the ligaments or abdominal wall region.

2-2. Qualitative analysis of biodegradable meshes

Elements present on the surface were analyzed using EDS (Energy Dispersive X-Ray Spectroscopy, EDAX, USA) depending on whether collagen was coated on the mesh for transplantation of the present invention. As a result, in the polycaprolactone mesh not coated with collagen, only carbon and oxygen components were detected, whereas in the specimen coated with collagen, nitrogen in the peptide was detected, so that 12.71% of the nitrogen element among the total element ratio was detected. It could be confirmed that there is (FIG. 6).

2-3. Cellular Reactivity of Biodegradable Meshes

Human dermal fibroblasts (LONZA, USA) were cultured on the mesh surface to evaluate the reactivity between the collagen-coated mesh for transplantation and cells in an in vitro environment. After placing the previously prepared circular specimens having a diameter of 1.5 cm on a 24-well tissue culture plate (TCP, Corning, USA), 70% ethanol was added and sterilization was performed for 30 minutes under a UV lamp. After that, fibroblasts (passage number 4) were seeded 50,000 in each specimen, and after seeding TCP as a control group, each fibroblast was incubated with DMEM (Dulbecco's Modified Eagle Medium, low glucose, Gibco, USA) for 7 days. Cells were cultured at 37°C under 5% carbon dioxide conditions using a medium mixed with 10 v/v% FBS (Fetal bovine serum, Gibco, USA) and 1 v/v% antibiotic (Gibco, USA). . At this time, for cell behavior analysis, live and dead assays (Thermo Fisher scientific, USA) were performed on the 1st and 7th days after the start of culture, respectively, to compare and analyze the survival/proliferation behavior of cells. To this end, at the end of the culture, each specimen was washed 3 times with a phosphate buffer solution (PBS, Gibco, USA), and 2 μ each of calcein AM and EthD-1 (Ethidium homodimer-1) in the live and dead assay kit were added. And after dilution to a concentration of 4 μ and added to each specimen, after staining at room temperature for 30 minutes, the stained cells were observed using a confocal fluorescence microscope (LSM700, Zeiss, Germany) (FIG. 7a) and quantitatively analyzed (FIG. 7b). As a result, on the first day of culture, 20-30 cells with high activity per unit area (1 mm ² ) were observed in all three groups of specimens, and there was no difference between the groups. However, by the 7th day of culture, the number of cells in the collagen-coated mesh group for transplantation per unit area was 7 times higher than that of the collagen-uncoated group and 3 times higher than that of TCP. Obvious cellular response results could be observed. This appears to be the result obtained because the porous collagen structure existing between the meshes provides enough space for cells to attach and proliferate. Accordingly, it can be seen that when the scaffold of the present invention is inserted into the human body after tissue dissection, the adhesion of various cells including the initial fibroblasts and the formation of blood vessels into the mesh can be efficiently induced.

Example 3: Biological safety of biodegradable mesh

3-1. Inflammation and biodegradation behavior of biodegradable mesh

In order to evaluate the inflammatory response and biodegradation behavior according to whether collagen is coated on the mesh for transplantation of the present invention, acellular allograft on the dorsal skin of SD (Sprague Dawley) rats (6 weeks old, male N=4, Orient Bio, Korea) After transplanting the mesh together with the dermis (thickness: 1.5mm, MegaDerm, L&C Bio, Korea) and euthanizing the rats at 6 weeks, 12 weeks, and 20 weeks, the tissues were collected and stained with Mason's trichrome (sigma aldrich, USA). The tissue section was observed with an optical microscope (CX43, Olympus, Tokyo, Japan) (FIG. 8a). In addition, the inflammatory response (Fig. 8b) around the graft and the degree of biodegradation of the graft (Fig. 8c) were analyzed.

As shown in Figure 8a, in normal tissue, the epidermis, dermis, and subcutaneous tissue of the skin were all clearly observed at the interface for 20 weeks, and in the group in which the mesh and acellular allogeneic dermis were inserted, the implants moved under the dermal tissue. It could be observed that they were inserted without However, unlike the acellular allogeneic dermis, it was confirmed that the tissue was filled between the porous structures formed by the mesh in all groups in which the mesh was inserted. , detachment from the tissue was observed.

In order to analyze the previously observed inflammatory response, the thickness of the film formed around the implant was measured (FIG. 8b). As a result of the measurement, at the 6th week of transplantation, it was observed that the acellular allogeneic dermis formed a film of about 250μm similar to the collagen-coated mesh. was able to confirm However, at week 20, it was confirmed that a thick film of about 340 µm was formed in the acellular allodermal group, whereas the film of 250 µm, similar to that of week 6, was maintained in the mesh group regardless of whether or not collagen was coated. Judging from the results of the Mason Trichrome staining photograph observed at the 20th week, it can be inferred that this result is caused by an excessive inflammatory reaction.

Next, in order to compare the biodegradation behavior of the implants, the thickness change of each implant for 20 weeks was measured ( FIG. 8c ). At week 6, 99% of the acellular allogeneic dermis remained close to the thickness of the first implanted implant, but about 78% remained in the mesh group, confirming a decrease in thickness of about 22%. This trend was maintained until week 12, and the thickness of the acellular allogeneic dermis was maintained at 92%, but the thickness of the mesh was maintained at about 70%. However, at the 20th week, the acellular allogeneic dermis had a rapid decrease in thickness compared to the 12th week, and only about 45% of the original thickness remained, confirming that rapid biodegradation occurred within 8 weeks. As shown in Fig. 8b, it can be seen that, at week 20, the thickest film around the graft was formed due to an inflammatory reaction according to the rapid biodegradation of the acellular allogeneic dermis inserted into the tissue.

3-2. Blood vessel formation inside the biodegradable mesh

To evaluate the angiogenesis-inducing ability of the mesh for transplantation of the present invention, whether or not collagen was coated, immunostaining was performed on the previously collected tissue, and the cell nucleus was 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) using a confocal microscope (LSM700, Carl Zeiss, Oberkochen, Germany). Observation (Fig. 9a) and (Fig. 9a) the number of blood vessels (arterioles) per area was comparatively quantified (Fig. 9b).

As can be seen from the fluorescence micrograph of Fig. 9a, uneven distribution of blood vessels was observed in the acellular allogeneic dermis over the 12th and 20th weeks, whereas the tissue into which the mesh was inserted went to the inside of the mesh regardless of the collagen coating. It was confirmed that the blood vessels were uniformly distributed. This phenomenon can be inferred from the local distribution of blood vessels due to the reduced cross-sectional area due to the acellular allogeneic dermis, which is rapidly decomposed at the 20th week and decreases in thickness.

Arterioles of SD rats are known to have a diameter of 20 to 40 μm, and the number of blood vessels satisfying the diameter condition of arterioles per unit area (mm ² ) was quantified through immunofluorescence staining ( FIG. 9b ). At 12 weeks after implant insertion, it was confirmed that about 16 blood vessels were observed in the acellular allogeneic dermis and mesh, whereas in the case of the collagen-coated mesh, about 23 blood vessels, 40% more than this, were distributed inside. did This trend was maintained until week 20, confirming that the collagen coating actively induces the formation of blood vessels inside the mesh.

Example 4: Preparation and Characterization of Collagen Sponge-Polymer Mesh Conjugate

4-1. Collagen Sponge Making

As another aspect of the present invention, the present inventors have prepared a collagen-containing sponge bonded with a polymer mesh, atelo collagen (Type 1, medical device grade, Dalimthysen, Korea) extracted from pig dermis in 0.5M acetic acid. It was dissolved at a concentration of 3.0 wt%. After being placed in a brass mold, immersed in liquid nitrogen (-196° C.) to freeze and freeze-dried for 24 hours according to the method described above in Example 1. Thereafter, the dried collagen sponge was subjected to dehydrothermal treatment (DHT) in an oven at 120° C. for 24 hours to prepare a collagen sponge ( FIG. 10 ).

4-2. Fabrication of Conjugate of Collagen Sponge Polymer Mesh through 3D Printing

In order to reinforce the physical properties of the prepared collagen sponge, the sponge is fixed on a 3D printing stage, and under the printing conditions applied for polymer mesh production in Example 1, the diameter of each strand is 0.4 mm and the spacing between the strands is 2.0 mm in the form of a mesh A PCL-collagen conjugate was prepared by direct printing on the sponge (FIG. 10).

Next, the surface and cross-sectional shape of the mesh structure contacted through 3D printing on the collagen sponge was observed using an electron microscope (FIG. 11). As a result, pores of 20 - 200 μm were formed on the surface of the collagen sponge, and it was confirmed through cross-sectional observation that the printed PCL and collagen form a stably bonded structure.

4-3. Analysis of Physical Strength of Collagen Sponge-Polymer Mesh Conjugate

In order to compare and analyze the physical strength of the prepared collagen sponge-polymer mesh conjugate, tensile strength and bonding strength with a suture used for fixing the human body were measured, respectively. As can be seen from the tensile strength measurement result of FIG. 12, compared to the simple collagen sponge, the conjugate of the present invention bonded with the PCL mesh had about 20 times higher tensile strength and about 70 times better tensile modulus, and the modulus of elasticity. Also, it was confirmed that it was about 10 times higher. Next, as a result of passing the suture through the collagen sponge and the collagen sponge-PCL mesh conjugate of the present invention, respectively, and analyzing the bonding strength with the suture, it was observed that the strength significantly increased from about 56.26KPa to 496.15KPa compared to the simple collagen sponge. did. Through these results, it was confirmed that the conjugate of the present invention in which the PCL polymer was in contact with the collagen sponge provides a stable fixation function in the human body while dramatically improving the physical properties of the existing collagen sponge.

As described above in detail a specific part of the present invention, for those of ordinary skill in the art, this specific description is only a preferred embodiment, and it is clear that the scope of the present invention is not limited thereto. Accordingly, the substantial scope of the present invention will be defined by the appended claims and their equivalents.

Claims (22)

delete delete delete delete delete delete delete delete delete delete delete delete delete delete A method for manufacturing a double-structured porous scaffold for human insertion, comprising embossing a polymer having biocompatibility on the surface of a collagen sponge obtained by freeze-drying a collagen solution or dispersion in a mesh shape, the embossed Method characterized in that it is performed by outputting the polymer in a mesh shape on the surface of the collagen sponge using a 3D printer.
delete delete 16. The method of claim 15, wherein the polymer is polycaprolactone (PCL), poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(lactic-co-glycolic acid) (PLGA), poly(PLCL) (L-Lactide-co-ε-caprolactone)) and combinations thereof.
19. The method of claim 18, wherein the polymer is polycaprolactone (PCL).
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