NOVEL POROUS SCAFFOLD AND METHOD FOR MANUFACTURING SAME

Abstract

The present invention relates to a porous scaffold having excellent tissue engineering properties, and a method for manufacturing same. The scaffold of the present invention can be manufactured by a simple process, and exhibits high tensile strength and biocompatibility, as well as an excellent cell engraftment rate, and thus can be useful as a support composition for various of human transplantation, for example, as a support for artificial ligaments or abdominal wall reinforcement.

Claims

1. A method for preparing a porous scaffold, comprising: (a) producing a polymer mesh having pores with an area of 0.1 to 0.5 mm.sup.2 and strands each having a diameter of 0.1 to 0.3 mm from a solution of a first polymer; and (b) coating the surface of the produced polymer mesh with a solution of a second polymer having biocompatibility.

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

3. (canceled)

4. The method of claim 1, wherein the second polymer having biocompatibility is collagen.

5. The method of claim 4, wherein the collagen solution has a concentration of 0.2 to 0.8% (v/v).

6. The method of claim 1, further comprising performing plasma treatment on a surface of the polymer mesh between the step (a) and the step (b).

7. The method of claim 6, wherein the plasma treatment is performed for 45 to 90 seconds.

8. A porous scaffold comprising: (a) a first polymer mesh having pores with an area of 0.1 to 0.5 mm.sup.2 and strands each having a diameter of 0.1 to 0.3 mm; and (b) a second polymer having biocompatibility with which the surface of the first polymer mesh is coated.

9. The porous scaffold of claim 8, wherein the first polymer is selected from the group consisting of polycaprolactone (PCL), poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(lactic-co-glycolic acid) (PLGA), poly(L-lactide-co-ε-caprolactone) (LCL), and combinations thereof.

10. (canceled)

11. The porous scaffold of claim 8, wherein the second polymer having biocompatibility is collagen.

12. A support composition for human body transplantation comprising the porous scaffold of claim 8.

13. The support composition of claim 12, wherein the support composition is 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, or abdominal wall reinforcement.

14. A method for tissue reconstruction comprising transplanting the support composition of claim 12 in vivo.

15. A method for preparing a dual structure porous scaffold comprising embossing a first polymer having biocompatibility into a mesh form on the surface of a support containing a second polymer having biocompatibility.

16. The method of claim 15, wherein the second polymer having biocompatibility is collagen.

17. The method of claim 16, wherein the support containing collagen is a collagen sponge.

18. The method of claim 15, wherein the first polymer is selected from the group consisting of polycaprolactone (PCL), poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), poly(lactic-co-glycolic acid) (PLGA), poly(L-lactide-co-ε-caprolactone) (LCL), and combinations thereof.

19. (canceled)

20. The method of claim 15, wherein the embossing is performed by outputting the first polymer in a mesh form using a three-dimensional printer on the surface of the second polymer-containing support.

21. The method of claim 15, wherein the mesh form includes strands each having a diameter of 0.3 to 0.5 mm and a spacing between the strands of 0.1 to 0.3 mm.

22. A dual structure porous scaffold including the following: (a) a support containing a second polymer having biocompatibility; and (b) a first polymer mesh which is bonded to the surface of the support and has biocompatibility.

23. A method for tissue reconstruction comprising transplanting the dual structure porous scaffold of claim 22 in vivo.

Description

BRIEF DESCRIPTION OF DRAWINGS

[0060] FIG. 1 shows the results of observing a polymer mesh according to the present disclosure prepared using a three-dimensional printer with an optical microscope.

[0061] FIG. 2 is optical photographs showing air bubbles on the surface of the polymer mesh generated after the polymer mesh according to the present disclosure is subjected to plasma surface treatment for various time periods and then coated with collagen.

[0062] FIG. 3 is a picture showing the macroscopic shapes of the collagen-coated meshes.

[0063] FIG. 4 is electron micrographs showing the microscopic shapes of the collagen-coated meshes.

[0064] FIG. 5 is a drawing showing the results of analyzing the physical strength values of a collagen-coated mesh for transplantation and acellular allogeneic dermis.

[0065] FIGS. 6A and 6B show the results of analyzing elements present on the surfaces of a mesh which is not coated with collagen (FIG. 6A) and a mesh which is coated with 0.5% collagen (FIG. 6B) using Energy Dispersive X-Ray Spectroscopy (EDS) (EDAX, USA).

[0066] FIGS. 7A and 7B show the results of observing active cells after cell culture (FIG. 7A) and the results of quantifying the observation results (FIG. 7B) in order to compare the cellular reactivities of the meshes depending on whether or not the meshes are coated with collagen.

[0067] FIGS. 8A, 8B and 8C show the results of staining the collected tissues with Masson's Trichrome by collecting tissues after transplanting the meshes into the acellular allogeneic dermis and the experimental animal dermis for 6, 12, and 20 weeks respectively in order to verify the biological safety of the meshes depending on whether or not the meshes are coated with collagen (FIG. 8A), and based on the staining results, shows the results of quantifying the thickness values of the films according to the inflammatory reaction (FIG. 8B) and the thickness values of the implants according to the biodegradation (FIG. 8C) respectively.

[0068] FIGS. 9A and 9B show the results of performing immunofluorescence staining on the tissues obtained after transplanting the meshes into the acellular allogeneic dermis and the animal dermis (FIG. 9A) in order to verify the distribution and number of blood vessels (arterioles) inside the meshes depending on whether or not the meshes are coated with collagen, and the results of quantifying the numbers of blood vessels (FIG. 9B) respectively.

[0069] FIG. 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.

[0070] FIG. 11 is electron micrographs showing a fine shape in which a polymer mesh is printed on a collagen sponge and bonded thereto.

[0071] FIG. 12 is diagrams showing the results of analyzing the physical properties of a collagen sponge and a structure in which a polymer mesh is printed on the collagen sponge and bonded thereto.

MODE FOR INVENTION

[0072] Hereinafter, the present disclosure will be described in more detail through examples. These examples are only for illustrating the present disclosure in more detail, and it will be obvious to those skilled in the art that the scope of the present disclosure is not limited by these examples according to the subject matter of the present disclosure.

EXAMPLE

Example 1: Preparation of Biodegradable Polymer Mesh

1-1. Fabrication of Polymer Mesh

[0073] A three-dimensional printer (Biobots, USA) was used in order to prepare a three-dimensional polymer structure, and the three-dimensional printing technique can easily adjust the size of a mesh depending on conditions such as nozzle diameter, temperature, discharge pressure, and nozzle movement speed. The present inventors selected a mesh form including strands each having a diameter of 0.2 mm and a spacing of 1.0 mm between the strands as the design that can most stably support the damaged ligament and abdominal wall (FIG. 1), and polycaprolactone (Sigma Aldrich, USA) was used as a raw material polymer.

[0074] In order to fabricate a polymer mesh, the diameter of the nozzle was set to 0.1 to 0.5 mm, the nozzle temperature was set to 80 to 90° C, the discharge pressure was set to 50 to 100 psi, and the nozzle movement speed was set to 2 to 5 mm/s. The polycaprolactone mesh prepared under these conditions was processed into a circular specimen having a diameter of 1.5 cm through a punching operation, washed with 70% ethanol for about 30 minutes in order to remove foreign substances, and then dried at room temperature for 2 hours.

1-2. Coating of Polymer Mesh with Collagen

[0075] The surface of the mesh was coated with collagen in order to impart biocompatibility to the polycaprolactone mesh prepared by three-dimensional printing. In order to homogeneously coat collagen, the present inventors introduced a pretreatment process that imparts hydrophilicity by subjecting polycaprolactone with strong hydrophobicity before coating to surface treatment using plasma. First of all, a collagen solution was prepared by dissolving atelocollagen (type 1, medical device grade, Dalim Tissen Co., Ltd., Korea) extracted from porcine dermis in 0.5 M acetic acid at a concentration of 0.5% at 4° C for 12 hours.

Exploration of Optimal Plasma Treatment Time

[0076] In order to select the optimal plasma treatment time for the most efficient collagen coating, the caprolactone mesh having been dried after washing was placed on a slide glass, and then treated using a plasma surface treatment machine (PDC-32G Plasma Cleaner, Harrick Plasma, USA) for 0, 15, 30, 45, and 60 seconds under medium vacuum conditions of 1.0 to 0.1 Torr. After the surface treatment process, 250 μl of a collagen solution per specimen was put therein to coat the mesh surface with collagen at 4° C for 30 minutes, and the collagen-coated mesh was observed with an optical microscope (EVOS® XL Core Cell Imaging System, Thermo Fisher scientific, USA) (FIG. 2). As shown in FIG. 2, it could be observed that the collagen-coated polycaprolactone mesh that had not been subjected to plasma surface treatment had not only an uneven collagen coating due to the strong hydrophobicity of the surface thereof, but also many bubbles generated on the surface thereof. It could be observed that, as the plasma treatment time was gradually increased at intervals of 15 seconds, the bubbles tended to decrease. When plasma was applied for 60 seconds, it could be observed that a uniform collagen coating film was formed on the surface of the mesh.

Exploration of Optimal Collagen Concentration

[0077] Thereafter, the present inventors tried to evaluate the optimal concentration of collagen with which the surface is coated by considering physical properties, biocompatibility, etc. that the polymer mesh should have as a human body insert. For this, collagen solutions were prepared by dissolving atelocollagen in 0.5 M acetic acid at various concentrations (0.1, 0.5, 0.75, and 1.0%) at 4° C for 12 hours, plasma surface treatment was performed for 60 seconds, and then 250 μl of the collagen solution was put into each of the mesh specimens to carry out a coating operation at 4° C for 30 minutes. Each sample that had been subjected to the coating operation was cooled to −70° C for 12 hours and then dried using a freeze dryer (FreeZone 12 plus, Labconco, USA) for 24 hours in order to create a porous surface structure of collagen with which the surface thereof is coated. Thereafter, a neutralization operation was performed in order to remove acetic acid present in the form of a salt inside freeze-dried collagen. For this, after the specimen that had been freeze-dried was washed 4 times for 15 minutes using anhydrous alcohol (ethanol absolute, Merck KGaA, Germany), 0.5 M NaOH (Duksan General Science, Korea) was dissolved in 70% ethanol, and then the neutralization operation of acetic acid was performed 4 times for 15 minutes. Thereafter, in order to remove the residual amount of NaOH present in the specimen, the collagen-coated mesh was sequentially washed 4 times for 15 minutes using 50% ethanol, 30% ethanol, and tertiary distilled water. After the collagen-coated mesh that had been washed was cooled to −70° C for 12 hours, and dried using a freeze dryer for 24 hours as mentioned above, images were obtained using a digital camera (EOS 500D, Canon, Japan) (FIG. 3). As a result, macroscopic shapes in which collagen blocked pores while being laminated in the form of a sponge on the mesh surface were observed from the group in which the meshes were coated with collagen at a concentration of 0.5% or more, this phenomenon worsened as the concentration of collagen increased, but the shapes of the meshes coated with collagen at a 0.1% concentration were completely preserved.

[0078] Next, the surface shapes of the collagen-coated meshes were observed using an electron microscope (FE-SEM, MERLIN, Zeiss, Germany) in order to observe the micro-shapes of the collagen-coated meshes (FIG. 4). As a result, when the meshes were 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 to 0.75%, the pores of collagen formed by freeze drying were decreased from about 500 μm to 20 μm. However, in the specimen coated with 1.0% collagen, the entire mesh surface was covered with collagen so that the pores could not be observed. The porous structure of collagen thus formed is a structure useful for initial cell attachment and the formation of blood vessels into the mesh when inserted into the human body. Judging from the results of surface observation using an electron microscope, it was determined that the mesh coated with 0.5% collagen, which was confirmed to have pores of about 150 to 300 μm, would be most suitable as a biodegradable mesh for transplantation.

Example 2: Analysis of Biodegradable Mesh Properties

2-1. Analysis of Physical Strength of Biodegradable Meshes

[0079] The tensile strength values were measured in order to analyze the physical strength values of the biodegradable meshes for transplantation fabricated in the present disclosure. In order to secure analysis results with higher reliability, acellular allogeneic dermis (CG Derm, Korea) commercially available as a ready-made article for the purpose of reconstruction of soft tissues of the human body was set as a comparison group, and the strength thereof was compared with that of the mesh for transplantation developed by this research team. To this end, after each specimen was processed into a 1 cm×5 cm rectangle and soaked in physiological saline for 30 minutes, the tensile strength was measured while the specimen was pulled at a speed of 1 mm per second using an all-around test analyzer (Universal Testing Systems, Instron 3360, USA). As a result, acellular allogeneic dermis that was the ready-made article showed a lower elastic force than the mesh for transplantation according to the present disclosure until it showed a tensile modulus of 50%, but showed the highest tensile strength of 15.27 MPa at the point of showing a tensile modulus of 124% (FIG. 5). On the other hand, it could be confirmed that the elastic modulus and tensile modulus of the mesh for transplantation according to the present disclosure were 2 times and 5 times higher than those of acellular allogeneic dermis, respectively, so that the elastic restoring force thereof was remarkably excellent. This high elastic restoring force of the mesh for transplantation according to the present disclosure shows that the mesh has very excellent properties as a human body insert for providing physical reinforcement to the ligaments, abdominal wall region, or the like.

2-2. Qualitative Analysis of Biodegradable Meshes

[0080] Elements present on the surface of the mesh for transplantation according to the present disclosure depending on whether or not the mesh is coated with collagen were analyzed using Energy Dispersive X-Ray Spectroscopy (EDS) (EDAX, USA). As a result, it could be confirmed that only carbon and oxygen components were detected in the polycaprolactone mesh that was not coated with collagen, whereas nitrogen in the peptide was detected in the specimen whose surface was coated with collagen so that 12.71% of the nitrogen element in the total element ratio was existed (FIG. 6).

2-3. Cellular Reactivity of Biodegradable Mesh

[0081] Human dermal-derived fibroblasts (LONZA, USA) were cultured on the mesh surface in order to evaluate the reactivity between the collagen-coated mesh for transplantation and cells in an in vitro environment. After the previously prepared circular specimens having a diameter of 1.5 cm were placed on a 24-well tissue culture plate (TCP, Corning, USA), 70% ethanol was put thereinto, and a sterilization operation was performed for 30 minutes under a UV lamp. Thereafter, 50,000 fibroblasts (passage number 4) were seeded in each specimen and cells were seeded even in TCP as a control group, and then each of the fibroblasts was cultured at 37° C under 5% carbon dioxide conditions using a medium in which 10 v/v % Fetal bovine serum (FBS) (Gibco, USA) and 1 v/v % antibiotic (Gibco, USA) were mixed with Dulbecco's Modified Eagle Medium (DMEM) (low glucose, Gibco, USA) for 7 days. At this time, in order to analyze behaviors of the cells, the survival/proliferation behaviors of the cells were comparatively analyzed by performing live and dead assays (Thermo Fisher Scientific, USA) on the 1st and 7th days after the start of culture. To this end, after each specimen was washed three times with a phosphate buffer solution (PBS, Gibco, USA) at the end of the culture, calcein AM and ethidium homodimer-1 (EthD-1) in the live and dead assay kit were diluted to concentrations of 2 μM and 4 μM respectively, the diluted solutions were put into each specimen, and the cells were stained at room temperature for 30 minutes, and then the stained cells were observed using a confocal fluorescence microscope (LSM700, Zeiss, Germany) (FIG. 7A), and the observed stained cells were quantitatively analyzed (FIG. 7B). As a result, on the first day of culture, 20 to 30 cells with high activity per unit area (1 mm.sup.2) were observed in all three groups of specimens, and there seemed no difference between the groups. However, on the 7th day of culture, it was observed that the number of cells per unit area in the collagen-coated mesh group for transplantation was 7 times higher than that of the collagen-uncoated group and about 3 times higher than that of TCP, suggesting that obvious cellular response results could be observed depending on the presence or absence of collagen. This appears to be the result obtained since the porous collagen structure existing between the meshes provides enough space for the cells to attach and proliferate. Accordingly, it can be seen that when the scaffold according to the present disclosure is inserted into the human body after tissue dissection, the attachment 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 Meshes

3-1. Inflammatory Responses and Biodegradation Behaviors of Biodegradable Meshes

[0082] In order to evaluate the inflammatory responses and biodegradation behaviors depending on whether or not the mesh for transplantation according to the present disclosure is coated with collagen, after transplanting meshes together with acellular allogeneic dermis (thickness: 1.5 mm, MegaDerm, L&C Bio, Korea) on the dorsal skin of Sprague Dawley (SD) rats (6 weeks old, male N=4, Orient Bio, Korea) and euthanizing the rats at week 6, week 12, and week 20 to collect tissues, the collected tissues were stained with Masson's Trichrome (Sigma Aldrich, USA) to observe the cross sections of the tissues with an optical microscope (CX43, Olympus, Tokyo, Japan) (FIG. 8A). Further, the inflammatory responses of the transplantation periphery (FIG. 8B) and the biodegradation degrees of the implants (FIG. 8C) were analyzed.

[0083] As shown in FIG. 8A, the epidermis, dermis, and subcutaneous tissues of the skin were all clearly observed at the interfaces of normal tissues for 20 weeks, and it could be observed in the groups into which the meshes and acellular allogeneic dermis were inserted that the implants were inserted under the dermal tissues without moving the position. However, unlike the acellular allogeneic dermis, it could be confirmed that the tissues were filled between the porous structures formed by the meshes in all groups in which the meshes were inserted, but in the acellular allogeneic dermis, films were formed thick due to excessive inflammatory responses at week 20, and a delamination phenomenon from the tissues was observed.

[0084] In order to analyze the previously observed inflammatory responses, the thickness values of the films formed on the implant periphery were measured (FIG. 8B). As a result of the measurement, it could be observed that the acellular allogeneic dermis formed a film of about 250 μm similar to that of the collagen-coated mesh in the 6th week of transplantation, and as it progressed to the 12th week, a 200 to 280 μm film was formed in all groups of implants so that similar numerical values could be confirmed. However, it could be confirmed that a thick film of about 340 μm was formed in the acellular allogeneic dermis group at week 20, whereas the film of 250 μm, similar to that of week 6, was maintained in the mesh group regardless of whether or not the meshes were coated with collagen. Judging from the results of the Masson's Trichrome staining photographs observed at week 20, such results may be inferred as a phenomenon caused by excessive inflammatory responses.

[0085] Next, in order to compare the biodegradation behaviors of the implants, changes in the thickness of each of the implants for 20 weeks were measured (FIG. 8C). At week 6, 99% of the acellular allogeneic dermis remained close to the thickness of the first inserted implant, but about 78% of the acellular allogeneic dermis remained in the mesh group, confirming a decrease in the 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 thicknesses of the meshes were maintained at about 70%. However, as a rapid decrease in the thickness of the acellular allogeneic dermis occurred at week 20 compared to week 12 so that only about 45% of the original thickness remained, it could be confirmed that rapid biodegradation occurred within 8 weeks. As shown in FIG. 8B, it can be seen from such results that the thickest film was formed on the transplantation periphery due to an inflammatory response according to the rapid biodegradation of the acellular allogeneic dermis inserted into the tissues at week 20.

3-2. Blood Vessel Formation Ability Inside the Biodegradable Mesh

[0086] After performing immunostaining on the previously collected tissues in order to evaluate the ability to induce angiogenesis depending on whether or not the mesh for transplantation according to the present disclosure is coated with collagen, staining the cell nucleuses with 4′,6-diamidino-2-phenylindole (DAPI, Blue signal, Sigma Aldrich, USA), and staining vascular endothelial cells with CD31 (Red signal, Thermo Fisher Scientific, Waltham, Mass., USA), the stained cell nucleuses and vascular endothelial cells were observed using a confocal microscope (LSM700, Carl Zeiss, Oberkochen, Germany) (FIG. 9A), and the numbers of blood vessels (arterioles) per area were comparatively quantified (FIG. 9B).

[0087] As can be seen from the fluorescence micrographs of FIG. 9A, uneven distribution of blood vessels was observed within the acellular allogeneic dermis over the week 12 and week 20, whereas it could be confirmed that the blood vessels were uniformly distributed to the inside of the meshes regardless of whether or not the meshes were coated with collagen in the tissues into which the meshes were inserted. Such a phenomenon may be inferred from the local distribution of blood vessels due to the reduction of the cross-sectional area caused by the acellular allogeneic dermis, which is rapidly decomposed and decreases in thickness at week 20.

[0088] Arterioles of the SD rats are known to have a diameter of 20 to 40 μm, and the numbers of blood vessels satisfying the diameter conditions of arterioles per unit area (mm.sup.2) was quantified through immunofluorescence staining (FIG. 9B). At week 12 after implant insertion, about 16 similar blood vessels were observed in the acellular allogeneic dermis and meshes, whereas it was confirmed that about 23 blood vessels, which are 40% more than the acellular allogeneic dermis and meshes, were distributed inside the collagen-coated mesh. This trend was maintained by week 20, confirming that the collagen coating actively induced the formation of blood vessels into the meshes.

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

4-1. Collagen Sponge Fabrication

[0089] As another aspect of the present disclosure, the present inventors dissolved atelocollagen (Type 1, medical device grade, Dalim Tissen Co., Ltd, Korea) extracted from porcine dermis in 0.5M acetic acid at a concentration of 3.0% by weight in order to fabricate a collagen-containing sponge bonded with a polymer mesh. Thereafter, after putting the dissolved atelocollagen in a brass mold, the brass mold was immersed in liquid nitrogen (−196° C) to freeze, and then 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. Preparation of Conjugate of Collagen Sponge Polymer Mesh through Three-Dimensional Printing

[0090] A PCL-collagen conjugate was fabricated by fixing the sponge to a three-dimensional printing stage in order to reinforce the physical properties of the prepared collagen sponge, and directly printing PCL on the sponge in a mesh form including strands each having a diameter of 0.4 mm and a spacing between the strands of 2.0 mm under the printing conditions applied for polymer mesh fabrication in Example 1 (FIG. 10).

[0091] Next, the surface shapes and cross-sectional shapes of the mesh structure in which PCL was bonded onto the collagen sponge through three-dimensional printing were observed using an electron microscope (FIG. 11). As a result, pores of 20 to 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 formed a stably bonded structure.

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

[0092] In order to compare and analyze the physical strengths of the prepared collagen sponge-polymer mesh conjugate, tensile strengths, and bonding strengths with a stitching fiber used when fixing the human body were respectively measured. As can be seen from the tensile strength measurement result of FIG. 12, it could be seen that the tensile strength of the conjugate according to the present disclosure to which the PCL mesh was bonded was about 20 times higher than that of the simple collagen sponge, and the tensile modulus thereof was also about 70 times superior to that of the simple collagen sponge, and it could be confirmed that the elastic modulus of the conjugate according to the present disclosure to which the PCL mesh was bonded was also about 10 times higher than that of the simple collagen sponge. Next, as a result of passing the stitching fiber through the collagen sponge and the collagen sponge-PCL mesh conjugate according to the present disclosure respectively and analyzing the bonding strengths with the stitching fiber, it was observed that the strengths remarkably increased from about 56.26 KPa to 496.15 KPa compared to the simple collagen sponge. Through these results, it could be confirmed that the conjugate according to the present disclosure in which the PCL polymer was bonded onto the collagen sponge provides a stable fixation function within the human body while dramatically improving the physical properties of the existing collagen sponge.

[0093] As the specific parts of the present disclosure have been described in detail above, these specific descriptions are only preferred embodiments for those of ordinary skill in the art, and it is clear that the scope of the present disclosure is not limited thereto. Accordingly, the substantial scope of the present disclosure will be defined by the appended claims and their equivalents.