Medical material and method for producing the same

Abstract

To provide a medical material suitable for injection into bone defect sites that produces excellent treatment outcomes.SOLUTION: A medical material comprises: porous fibers formed from a bioabsorbable polymer, wherein the porous fibers have a short average fiber length and include a dispersed apatite precursor; and a hydrogel. The average fiber length of the porous fibers is in the range of 100 μm to 400 μm. The bioabsorbable polymer is selected from polylactic acid, polyglycolic acid, lactic acid-glycolic acid copolymer, and polycaprolactone.SELECTED DRAWING: Figure 1

Description

[Technical field]

[0001] The present invention relates to a medical material for injection into a bone defect and a method for producing the same. [Background technology]

[0002] In recent years, a method of injecting gel into the affected area has been attracting attention as a treatment for bone defects. Compared to conventional autologous bone grafts and artificial bone grafts, this method is less invasive and is expected to reduce the risk of infection and shorten recovery time (Non-Patent Documents 1, 2).

[0003] However, since it takes time to acquire mechanical properties, there is concern that bone formation will take time. Therefore, there is a need for a scaffold that can acquire mechanical properties quickly to promote bone formation and can improve cell activity such as cell adhesion, proliferation, and differentiation.

[0004] For example, it has been reported that porous PLLA / DCPA fibers were produced by adding dicalcium phosphate anhydrous (CaHPO4, DCPA) particles, a precursor of hydroxyapatite (HAp), the main component of bone, to porous PLLA fibers made of poly-L-lactic acid (PLLA), a bioabsorbable polymer, and coating the fiber surface with apatite quickly and without peeling off (Non-Patent Document 3).

[0005] Also, for example, it has been reported that coating a PLLA fiber scaffold with apatite promoted differentiation of osteoblast-like cells (MC3T3-E1) (Non-Patent Document 4). [Prior art documents] [Non-patent literature]

[0006] [Non-Patent Document 1] Yamada, Y., Jae SB, Ozawa R, Nagasaka, T., Okazaki, Y., Hata, K., Ueda, M., Bone regeneration following injection of mesenchymal stem cells and fibrin glue with a biodegradable scaffold, Journal of Cranio-Maxillofacial Surgery, Vol. 23, (2003), pp. 27-33. [Non-Patent Document 2] Qingpu, H., Paul, A., De, B., Kevin, MS, Injectable scaffolds for tissue regeneration, Journal of Materials Chemistry, Vol. 14, (2004), pp.1915-1923. [Non-Patent Document 3] Matsuo, H., Nakamachi, E., Yamamoto, K., Morita Y., European Society of Biomechanics, (2019), pp. 1-2. [Non-Patent Document 4] Whited, BM, Whitney JR, Hofmann MC, Xu , Y., Rylander MN, Pre-osteoblast infiltration and differentiation in highly porous apatite-coated PLLA electrospun scaffolds, Biomaterials, Vol. 32, No. 9, (2011), pp. 2294-2304. Summary of the Invention [Problem to be solved by the invention]

[0007] However, the above-mentioned medical materials for treating bone defects do not have sufficient mechanical strength.

[0008] The present invention has been made in consideration of the above problems, and an object of the present invention is to provide a medical material for injection into a bone defect site, which can provide an excellent therapeutic effect, and a method for producing the same. [Means for solving the problem]

[0009] The medical material according to the present invention is characterized by comprising porous fibers of a bioabsorbable polymer having a short average fiber length and containing dispersed apatite precursors, and a hydrogel.

[0010] The method for producing a medical material according to the present invention is characterized by comprising the steps of: extruding a bioabsorbable polymer solution mixed with an apatite precursor by electrospinning to produce porous fibers of the bioabsorbable polymer; freeze-pulverizing the porous fibers to produce porous fibers having a short average fiber length; and mixing the short-average porous fibers with a hydrogel. Effect of the Invention

[0011] According to the present invention, a medical material that provides excellent therapeutic effects can be obtained. [Brief description of the drawings]

[0012] [Figure 1] FIG. 1 shows SEM images of porous PLLA / DCPA fibers with fiber diameters of 5, 10, and 15 μm before freeze-milling. [Diagram 2] FIG. 1 shows SEM images of porous PLLA short microfibers after five grinding cycles with different grinding times, where (a) is a grinding time of 1 second, (b) is a grinding time of 3 seconds, (c) is a grinding time of 5 seconds, and (d) is a grinding time of 10 seconds. In each figure, Bar: 50 μm, and magnification is ×500. [Diagram 3]This figure shows the fiber length of porous PLLA short microfibers for each grinding number and grinding time, where (a) is a grinding time of 1 second, (b) is a grinding time of 3 seconds, (c) is a grinding time of 5 seconds, and (d) is a grinding time of 10 seconds. [Figure 4] FIG. 1 shows SEM images of porous PLLA / DCPA fibers, of which (a) is the surface and (b) is the cross section. Bar: 5 μm in each image, magnification ×5000. [Diagram 5] FIG. 1 shows SEM images of porous PLLA short microfibers, of which (a) is before immersion in SBF, (b) is after 1 day of immersion in SBF, (c) is after 3 days of immersion in SBF, and (d) is after 7 days of immersion in SBF. In each figure, Bar: 10 μm, magnification ×1000. [Figure 6] FIG. 1 shows SEM images of porous PLLA short microfibers with different fiber diameters and grinding times. [Figure 7] FIG. 13 shows the fiber diameter of porous PLLA short microfibers before and after immersion in SBF. [Figure 8] FIG. 13 shows images of the control and porous PLLA / DCPA short fiber mixed agarose gels under each condition before and after 7 days of immersion in SBF. [Figure 9] FIG. 13 shows alizarin stained images of control and porous PLLA / DCPA short fiber mixed agarose gels after immersion in SBF for 1, 3, and 7 days. [Figure 10] FIG. 1 shows SEM images of short microfibers in each agarose gel after immersion in SBF for 1, 3, and 7 days. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0013] Hereinafter, an embodiment of the present invention will be described in detail with reference to the attached drawings. However, the embodiment is intended to facilitate understanding of the principles of the present invention, and the scope of the present invention is not limited to the embodiment described below. Other embodiments in which a person skilled in the art appropriately replaces the configuration of the embodiment described below are also included in the scope of the present invention.

[0014] The medical material according to the present invention is a medical material for injection into a bone defect, and is characterized by comprising a hydrogel and porous fibers of a bioabsorbable polymer having a short average fiber length and containing dispersed apatite precursors.

[0015] In the present invention, since the average fiber length of the porous fibers is short, the porous fibers can be easily mixed with the hydrogel, and the medical material can be easily injected into the bone defect.

[0016] In the present invention, since the porous fibers of a bioabsorbable polymer having a short average fiber length and containing dispersed apatite precursors are used, and the hydrogel are used, a network structure is formed within the hydrogel.

[0017] Furthermore, the porous fibers of the bioabsorbable polymer containing dispersed apatite precursors are bonded to each other by the precipitated apatite to form a network structure, which results in a stronger structure and increased mechanical strength.

[0018] In addition, the precipitated apatite is expected to contribute to improving cell activity.

[0019] The average fiber length of the porous fibers is short, and is not particularly limited, but is, for example, 100 μm to 400 μm.

[0020] The bioabsorbable polymer is not particularly limited, but examples thereof include polylactic acid, polyglycolic acid, lactic acid-glycolic acid copolymer, polycaprolactone, polyglycerol sebacic acid, polyhydroxyalkanoic acid, polybutylene succinate, or derivatives thereof, and preferably polylactic acid.

[0021] Monomers constituting the polymer in polylactic acid include, for example, L-lactic acid and D-lactic acid, but are not particularly limited. In addition, there are no particular limitations on the optical purity, molecular weight, composition ratio of L- and D-lactic acid, or sequence of the polymer, and a stereocomplex of poly-L-lactic acid and poly-D-lactic acid may be used. A polymer containing a large amount of L-lactic acid is preferable, and it is more preferable that the L-lactic acid is 100%.

[0022] The molecular weight of the bioabsorbable polymer is not particularly limited, but is, for example, 1×10 3 ~5×10 6 and preferably 1×10 4 ~1×10 6 , more preferably 5 × 10 4 ~5×10 5 In addition, the terminal structure of the polymer and the catalyst for polymerizing the polymer can be selected arbitrarily.

[0023] The porous fibers of the bioabsorbable polymer have, for example, an average fiber diameter of 0.01 to 10 μm and an average apparent density of 10 to 300 kg / m 3 It is.

[0024] Considering the rate of absorption in the body, the porosity of the porous fiber of the bioabsorbable polymer is preferably 20 to 80%. If the porosity is less than 20%, the number of cells that can be attached is small and there is a risk that the apatite coating will peel off. If the porosity is more than 80%, the number of cells that can be attached is large, but the mechanical strength is low and it may not be preferable as a scaffold material. The porosity can be calculated from the average apparent density and the inherent density of the polymer.

[0025] The apatite precursor is either anhydrous calcium hydrogen phosphate (DCPA), monocalcium phosphate, dicalcium phosphate, tricalcium phosphate, tetracalcium phosphate or calcium octaphosphate, preferably anhydrous calcium hydrogen phosphate (DCPA).

[0026] The hydrogel is selected from the group consisting of collagen, agarose, alginic acid, gelatin, hydroxypropyl cellulose, CMC, fibrinogen, thrombin, cellulose, glycosaminoglycan, chitin, chitosan, hyaluronic acid, starch, fibroin, polyglycolic acid, polylactide, polyglycolide, polycaprolactone, polyethylene oxide, and polyethylene glycol.

[0027] The porous fibers of the bioabsorbable polymer of the present invention can contain cell growth factors, such as FGF (fibroblast growth factor), EGF (epidermal growth factor), PDGF (platelet-derived growth factor), TGF-β (transforming growth factor β), NGF (nerve growth factor), HGF (hepatocyte growth factor), and BMP (bone morphogenetic protein).

[0028] The porous fiber of the bioabsorbable polymer of the present invention can be used, for example, in orthopedic surgery, dental oral surgery, and bone defect repair after craniotomy or thoracotomy. Specifically, in orthopedic surgery, the porous body of the present invention can be injected into bone defects caused by trauma such as bone defects after bone tumor resection or fractures to promote bone regeneration. In dental oral surgery, the porous body of the present invention can be injected into bone defects caused by periodontal disease, alveolar clefts, tooth extraction sockets, and the like, and excellent bone regeneration effects can be confirmed in a short period of time.

[0029] A method for producing the medical material according to the present invention will now be described.

[0030] In the first step, a bioabsorbable polymer solution containing an apatite precursor is extruded by electrospinning to produce porous fibers of the bioabsorbable polymer.

[0031] That is, a solution is prepared by dissolving a bioabsorbable polymer in an organic solvent, and then the hydroxyapatite precursor is added to the solution and mixed. The organic solvent is not particularly limited, but examples thereof include halogenated solvents such as 1,3-dioxolane, dichloromethane, chloroform, dichloroethane, dichloroethylene, carbon tetrachloride, trichloroethane, hexafluoro-2-propanol, and hexafluoroacetone.

[0032] Next, the polymer dissolved in the organic solvent is extruded using an electrospinning device to obtain a porous fiber of a bioabsorbable polymer. The obtained porous fiber contains a dispersed hydroxyapatite precursor. The electrospinning device is not particularly limited, but may be, for example, an extrusion section for extruding the polymer dissolved in the organic solvent, a nozzle section for discharging the molten polymer, a voltage application section for applying a voltage to form nanofibers, and a collection section for collecting the obtained nanofibers. Methods for forming nanofibers from the dissolved polymer include a hot melt electrospinning method, a wet electrospinning method, a blow spinning method, and a force spinning method.

[0033] In the second step, the porous fibers are freeze-pulverized to produce porous fibers with a short average fiber length.

[0034] Since the porous fiber is freeze-pulverized, it is preferable that the bioabsorbable polymer forming the porous fiber has a glass transition temperature Tg of 0° C. or higher or is a crystalline resin. This is because freeze-pulverization is facilitated by satisfying any of these conditions. The freeze-pulverization method is not particularly limited, but it is preferable to immerse the prepared porous fiber in water, cool it, and freeze it, thereby sealing the porous fiber in an ice block, and then freeze-pulverize the ice block. This is because freezing makes it possible to pulverize the porous fibers of the bioabsorbable polymer while maintaining their porosity.

[0035] In the third step, the short porous fibers are mixed with the hydrogel. That is, the short porous fibers are mixed with PBS or the like to prepare a suspension, and the suspension is mixed with the hydrogel solution. This makes it possible to manufacture the medical material according to the present invention.

[0036] Although the present invention is a medical material for injection into a bone defect, characterized by having a porous fiber made of a bioabsorbable polymer having a short average fiber length and containing dispersed apatite precursor, and a hydrogel, it is also possible to configure the medical material as a medical material having a porous fiber made of a bioabsorbable polymer having a short average fiber length and a hydrogel. It is also possible to configure the medical material as a medical material consisting of a porous fiber made of a bioabsorbable polymer having a short average fiber length and containing dispersed apatite precursor without using a hydrogel. EXAMPLES

[0037] (1) Preparation of porous PLLA / DCPA short fibers (1-1) Preparation of DCPA particles Anhydrous calcium hydrogen phosphate (DCPA) particles were prepared using CaCl2 and Na2HPO4·12H2O. DCPA is an acidic to neutral pH material and has a Ca 2+ and PO4 3- A stable precipitation occurs when the ratio of CaCl2 and Na2HPO4·12H2O is mixed to a concentration of 1.0. A solution was prepared by dissolving 6.00 g of CaCl2 and 19.38 g of Na2HPO4·12H2O in 300 ml of distilled water to obtain a Ca / P ratio of 1.0. The solution was stirred for 1 hour with a stirrer to react, and then washed twice with 300 ml of distilled water to remove NaCl produced as a by-product. After washing, the resulting particles were wet-milled to prepare DCPA particles.

[0038] (1-2) Preparation of spinning solution A spinning solution was prepared by mixing the DCPA particles prepared in (1-1) above, poly-L-lactic acid (PLLA, Mn=173,300, LUMINY L 130, Corbion), and polyethylene oxide (PEO, Alfa Aesar) with a solvent. The solvent used was 500 g of 1,3-dioxolane (DOL, 13643-45, Nacalai Tesque, Inc.) with 75 μl of triethylamine (TEA, T0424, Tokyo Chemical Industry Co., Ltd.) added as a stabilizer. First, the DCPA particles and the solvent were stirred for 30 minutes using a stirring and degassing device (ARE-310, Shinki Co., Ltd.) to suppress the aggregation of the DCPA particles. After stirring, PLLA and PEO were added and stirred for 30 minutes. The solution concentration was 6 w / v%, and the ratio of PLLA / PEO was 90 / 10 to make the fibers porous. The mass ratio of DCPA to PLLA was 5 / 1.

[0039] (1-3) Fiber fabrication by electrospinning The nanofiber electrospinning unit used for spinning was NEU-11-TEA-2622 type (Kato Tech Co., Ltd.). Each solution prepared in (1-2) above was filled into a 20 mL syringe (SS-20ESZ, slip tip (side mouth), φ21.7×123.6 mm, Terumo Co., Ltd.). An 18 G injection needle with a flatly polished tip (NN-1838R, φ1.2×38 mm, Terumo Co., Ltd.) was used as the nozzle. A hollow aluminum shaft with a diameter of φ9 was used as the target, and the target rotation speed was 100 rpm, the spinning distance was 200 mm, and the spinning time was 8 minutes. In addition, the Taylor cone formed at the needle tip during spinning was used as a guide for spinning, and the applied voltage was 7.5-8.0 kV and the discharge amount was 100 μl / min.

[0040] To evaluate the effects of fiber diameter and length, porous PLLA / DCPA fiber scaffolds with fiber diameters of 5, 10, and 15 μm were prepared. After spinning, the morphology was observed by SEM. Figure 1 shows SEM images of the porous PLLA / DCPA fibers with fiber diameters of 5, 10, and 15 μm. It was confirmed that the surface and interior of the fiber had a porous structure under all conditions. In addition, it was confirmed that DCPA particles were contained from the surface to the interior of the porous PLLA / DCPA fiber.

[0041] (1-4) Preparation of short microfibers by freeze-grinding method The porous PLLA fiber scaffold prepared in (1-3) above was cut into 10 mm squares, placed in a mold (35 × 35 × 6 mm, thickness 2 mm) prepared by thermally curing polydimethylsiloxane (PDMS), and immersed in 99.5% ethanol (054-00466, Fujifilm Wako Pure Chemical Industries, Ltd.). PDMS was prepared by mixing resin SILPOT 184W / C and hardener CATALYST SILPOT 184 (KYK Technologies, Inc.) in a ratio of 10:1, and the cloudy resin was used after vacuum degassing. The 99.5% ethanol was removed, and the scaffold was immersed in DW (distilled water) and frozen to obtain ice with a side length of 35 mm and a thickness of 6 mm, which contained the fiber scaffold. In order to facilitate crushing of the fiber scaffold, it was frozen using liquid nitrogen and crushed using a freeze crusher (HTPH-01, AS ONE Corporation). The conditions were 1, 3, 5, and 10 seconds for the grinding time, and 1 to 5 grinding times. When the grinding times were two or more, the fibers that had been freeze-ground once were transferred together with the ice to a PDMS mold, the ice was melted, and the fibers were frozen in an environment of -80°C, and then freeze-ground again for the same grinding time. After grinding, the fibers were dried using a freeze dryer (FDU-1200, Tokyo Rikakikai Co., Ltd.) to obtain porous PLLA short microfibers. The morphology of the short fibers was observed by SEM. The accelerating voltage during SEM observation was 8 kV, and the SS (spot size) was 45. The fiber length of the short fibers was measured by Image J. Similarly, the porous PLLA / DCPA short microfibers were prepared by freeze-grounding the porous PLLA / DCPA fiber scaffolds prepared in (1-3) above with a grinding time of 10 seconds and a grinding time of one time.

[0042] (2-1) Apatite Precipitation Ability Evaluation In order to evaluate the apatite precipitation ability of the prepared porous PLLA / DCPA short microfibers in vivo, we performed an immersion experiment using simulated body fluid (SBF). SBF is registered with the International Organization for Standardization as a standard fluid for evaluating the apatite precipitation ability of artificial material surfaces in vivo. SBF was prepared according to the following procedure.+ 142 mM,K + 5.0 mM, Mg 2+ 1.5 mM Ca 2+ 2.5 mM Cl - 148.8 mM,HCO3 - 4.2 mM,HPO4 2- 1.0 mM and SO4 2- NaCl (191-01665, Fujifilm Wako Pure Chemical), NaHCO3 (199-01301, Fujifilm Wako Pure Chemical), KCl (28514-75, Nacalai Tesque), K2HPO4·3H2O (28728-85, Nacalai Tesque), MgCl2·6H2O (20909-55, Nacalai Tesque), CaCl2, and Na2SO4 (31916-15, Nacalai Tesque) were dissolved in distilled water at 36.5 ± 1.5 °C in that order to a concentration of 0.5 mM. After dissolution, SBF was prepared using Tris (hidroxymethyl) aminomethane (T1503, SIGMA-ALDRICH) and 1 M-HCl to a temperature of 36.5 °C and pH 7.40. 1 M-HCl was prepared by dissolving 35% concentrated hydrochloric acid (HCl, 7647-01-0, Sigma-Aldrich Japan Co., Ltd.) in distilled water and diluting the solution with distilled water.

[0043] The porous PLLA / DCPA short microfibers prepared in (1-4) above were placed in a 2 ml tube and immersed in 1 ml of SBF at 37°C for 1, 3, and 7 days. 500 μl of SBF was replaced every day. After immersion in SBF, 900 μl of SBF was removed, and the fibers were washed twice with 1 ml of distilled water to remove salts attached to the surface of the fibers. In order to prevent the fibers from being accidentally sucked up, the fibers were centrifuged using a centrifuge (Tabletop Cooled Centrifuge 3500, KUBOTA) when replacing the SBF and washing with distilled water, and the SBF or distilled water remaining as the supernatant was removed. The conditions were a rotation speed of 13,000 rpm and a time of 10 min. After washing, the fibers were dried using a freeze dryer, and the morphology of the fibers was observed by SEM. The accelerating voltage during SEM observation was 8 kV and SS was 45. In addition, the fiber diameters of the fibers before and after immersion were measured using a Sigma scan pro. Measurements were taken at three points per fiber, and the average value was used as the representative value.

[0044] (2-2) Statistical processing In the SBF immersion experiment, normality was tested using the fiber diameter data for each immersion period. If normality was observed, one-way analysis of variance was performed, followed by the Tukey-Kramer test for multiple comparisons. On the other hand, if normality was not observed, the nonparametric Kruskal-Wallis test was performed. The significance level was set at 5%.

[0045] (2-3) Results Figures 2 and 3 show the results of the fiber length of the porous PLLA short microfibers depending on the number of times of grinding and the grinding time. Figure 2 shows SEM images of the porous PLLA short microfibers depending on the grinding time after five grindings, where (a) is a grinding time of 1 second, (b) is a grinding time of 3 seconds, (c) is a grinding time of 5 seconds, and (d) is a grinding time of 10 seconds. In each figure, the bar is 50 μm and the magnification is ×500. Figure 3 shows the results of the fiber length of the porous PLLA short microfibers depending on the number of times of grinding and each grinding time, where (a) is a grinding time of 1 second, (b) is a grinding time of 3 seconds, (c) is a grinding time of 5 seconds, and (d) is a grinding time of 10 seconds. At a grinding time of 1 second, the average fiber lengths of the first to fifth grindings were 301, 274, 269, 224, and 181 μm, respectively. At a grinding time of 3 seconds, the average fiber lengths after grinding counts 1 to 5 were 282, 189, 102, 122, and 105 μm, respectively. At a grinding time of 5 seconds, the average fiber lengths after grinding counts 1 to 5 were 316, 225, 179, 169, and 150 μm, respectively. At a grinding time of 10 seconds, the average fiber lengths after grinding counts 1 to 5 were 290, 213, 165, 124, and 120 μm, respectively. At all grinding times, the fiber length tended to become shorter with an increase in the grinding count. In addition, the variation in fiber length tended to decrease. Based on this, porous PLLA short microfibers were obtained by freeze-grinding the porous PLLA fibers produced by the electrospinning method. In addition, we succeeded in controlling the fiber length to the order of tens to hundreds of μm by adjusting the conditions such as the number of freeze-grinding times and the time.

[0046] An SEM image of the porous PLLA / DCPA short microfibers after milling is shown in Figure 4. It was confirmed that both ends of the fiber were not crushed and the porous structure on the surface and inside of the fiber was maintained.

[0047] Figure 5 shows SEM images of porous PLLA / DCPA short microfibers of each fiber diameter before and after immersion. After 1 day of SBF immersion, precipitated apatite was present in particulate form, and most of the fiber surface was covered with apatite. After 3 days of SBF immersion, more apatite precipitated than after 1 day of SBF immersion, and the entire fiber surface was covered with a smooth apatite coating. It was also confirmed that the apatite-coated fibers were bonded together. After 7 days of SBF immersion, the unevenness of the apatite coating on the fiber surface was reduced compared to after 3 days of SBF immersion, and it was confirmed that the apatite-coated fibers were bonded together in the same way as after 3 days of SBF immersion. Figure 6 shows the morphology of porous PLLA short microfibers for each fiber diameter and grinding time.

[0048] The fiber diameter before and after SBF immersion is shown in Figure 7. Before SBF immersion, and after 1, 3, and 7 days, the diameters were 6.1±0.8, 7.9±1.1, 8.3±0.9, and 9.2±0.8 μm, respectively. There was a tendency for the fiber diameter to increase with increasing SBF immersion days.

[0049] (3) Evaluation of bone matrix production (3-1) Preparation of porous PLLA / DCPA short microfiber mixed agarose gel A porous PLLA / DCPA short microfiber suspension was prepared by suspending the porous PLLA / DCPA short fibers prepared in 1-4 in PBS. The prepared suspension was mixed 1:1 with agarose gel solution with a concentration of 2.0 w / v% to prepare a porous PLLA / DCPA short fiber mixed agarose solution with a final concentration of 1.0 w / v%. The amount of porous PLLA / DCPA short fibers mixed was 0 (control), 0.2, 0.4, and 0.8 w / v% relative to the amount of agarose solution. The agarose gel solution with a concentration of 2.0 w / v% was prepared by dissolving agarose powder (A4018-5G, SIGMA-ALDRICH) in PBS. The prepared porous PLLA / DCPA short fiber mixed agarose solution and agarose solution were dropped in an amount of 80 μl into a homemade polycarbonate mold, sealed with a cover glass, and then solidified for 30 minutes at 4°C. A porous PLLA / DCPA short microfiber mixed agarose gel with a diameter of 6 mm and a thickness of 2 mm was prepared by hollowing out the mold with a 6 mm diameter biopsy trephine.

[0050] (3-2) Apatite Precipitation Ability Evaluation Each gel prepared in 3-1 was placed in one well of a 24-well plate and soaked in 1 ml of SBF for 1, 3, and 7 days at 37°C. SBF was replaced by 1 ml every day. After soaking in SBF, the gels were washed twice with 1 ml of distilled water to remove salts adhering to the surface of the gels.

[0051] After washing, each gel was cut in half with a scalpel blade. The cut gels were placed in one well of a 24-well plate and stained with alizarin to observe the distribution of Ca due to the precipitation of hydroxyapatite in the gel. After staining, the staining reagent was removed and the gels were photographed with a digital camera.

[0052] The surfaces of the fibers in the 0.2, 0.4, and 0.8 w / v% gels were observed by SEM. After immersion in SBF, the gels were finely chopped with a scalpel blade and freeze-dried. The freeze-dried gels were coated with osmium to a thickness of 40 nm using an osmium plasma coater to impart electrical conductivity. The morphology of the fibers emerging from each gel was then observed by SEM.

[0053] Figure 8 shows images of the control and porous PLLA / DCPA short fiber mixed agarose gels under each condition before and after 7 days of SBF immersion. Before SBF immersion, the fibers in the 0.2, 0.4, and 0.8 w / v% porous PLLA / DCPA short fiber mixed agarose gels were uniformly present throughout the gel, and the agarose gels tended to become whiter as the amount of fiber mixed increased. On the other hand, after 7 days of SBF immersion, the color of the control agarose gel was transparent, but the entire agarose gel appeared to be cloudy under the 0.2, 0.4, and 0.8 w / v% conditions. There was also little difference in the degree of cloudiness of the 0.2, 0.4, and 0.8 w / v% agarose gels. Figure 9 shows alizarin-stained images of the control and porous PLLA / DCPA short fiber mixed agarose gels after 1, 3, and 7 days of SBF immersion. In the control, the color of the agarose gel was transparent regardless of the number of days of immersion. On the other hand, in the 0.2, 0.4, and 0.8 w / v% conditions, the agarose gel was stained red overall after 3 days of SBF immersion, confirming that hydroxyapatite had precipitated throughout, and after 7 days of SBF immersion, there was a tendency for the area of ​​the gel stained deep red to increase with increasing amounts of fiber mixed in.

[0054] Figure 10 shows SEM images of short microfibers in each agarose gel after 1, 3, and 7 days of SBF immersion. After 1 day of SBF immersion, no apatite precipitation was observed on the fiber surface under any of the conditions. After 3 days of SBF immersion, apatite was observed to have precipitated on the fiber surface under all of the conditions, covering most of the fiber surface. After 7 days of SBF immersion, it was confirmed that the fiber surface was more covered with apatite under all of the conditions compared to after 3 days of SBF immersion. In addition, adhesion between fibers with apatite coating was occasionally observed. [Industrial Applicability]

[0055] It can be used to treat bone defects.

Claims

1. A porous fiber made of a bioabsorbable polymer having a short average fiber length and containing an apatite precursor dispersed therein; A hydrogel, A medical material for injection into a bone defect, comprising:

2. 2. The medical material according to claim 1, wherein the porous fibers have an average fiber length of 100 μm to 400 μm.

3. 2. The medical material according to claim 1, wherein the bioabsorbable polymer is any one of polylactic acid, polyglycolic acid, lactic acid-glycolic acid copolymer, and polycaprolactone.

4. 3. The medical material according to claim 1, wherein the apatite precursor is any one of anhydrous calcium hydrogen phosphate (DCPA), monocalcium phosphate, dicalcium phosphate, tricalcium phosphate, tetracalcium phosphate, and calcium octaphosphate.

5. 2. The medical material according to claim 1, characterized in that the hydrogel is selected from the group consisting of collagen, agarose, alginic acid, gelatin, hydroxypropyl cellulose, CMC, fibrinogen, thrombin, cellulose, glycosaminoglycan, chitin, chitosan, hyaluronic acid, starch, fibroin, polyglycolic acid, polylactide, polyglycolide, polycaprolactone, polyethylene oxide, and polyethylene glycol.

6. A step of extruding a bioabsorbable polymer solution mixed with an apatite precursor by electrospinning to prepare a porous fiber of the bioabsorbable polymer; freeze-pulverizing the porous fibers to produce porous fibers having a short average fiber length; mixing the short porous fibers with a hydrogel; 1. A method for producing a medical material to be injected into a bone defect, comprising the steps of:

7. The method for producing a medical material according to claim 6, characterized in that after the step of producing the porous fiber, the produced porous fiber is immersed in water, cooled and frozen to encapsulate the porous fiber within an ice block, and the ice block is freeze-crushed to produce porous fibers having a short average fiber length.

8. The method for producing a medical material according to claim 6, characterized in that the porous fibers have an average fiber length of 100 μm to 400 μm.

9. 7. The method for producing a medical material according to claim 6, wherein the bioabsorbable polymer is any one of polylactic acid, polyglycolic acid, lactic-co-glycolic acid copolymer, and polycaprolactone.

10. 7. The method for producing a medical material according to claim 6, wherein the apatite precursor is any one of calcium hydrogen phosphate anhydrous (DCPA), monocalcium phosphate, dicalcium phosphate, tricalcium phosphate, tetracalcium phosphate, and calcium octaphosphate.

11. The method for producing a medical material according to claim 6, characterized in that the hydrogel is selected from the group consisting of collagen, agarose, alginic acid, gelatin, hydroxypropyl cellulose, CMC, fibrinogen, thrombin, cellulose, glycosaminoglycan, chitin, chitosan, hyaluronic acid, starch, fibroin, polyglycolic acid, polylactide, polyglycolide, polycaprolactone, polyethylene oxide, and polyethylene glycol.

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