Fixing material for medical device implant and preparation method of fixing material

Abstract

The invention discloses a fixing material for a medical device implant and a preparation method of the fixing material. The fixing material comprises a degradable polymer base material with a porous structure and a drug sustained-release coating uniformly loaded on the degradable polymer base material, holes in the degradable polymer are uniformly distributed in a staggered array manner, the holes are circular, the aperture of each hole is 0.5-3.0 mm, and the distance between every two adjacent holes is 0.5-3.0 mm; the drug sustained-release coating contains antibiotics; the antibiotics comprise rifampicin, minocycline hydrochloride, vancomycin, tobramycin, gentamicin and the like. The fixing material disclosed by the invention takes the degradable polymer as a base material, and the drug-loading coating containing antibiotics and anti-inflammatory substances is loaded on the surface, so that slow release of antibiotic drugs can be realized, and the fixing material has excellent antibacterial and anti-inflammatory performance and complete absorbability in a biological environment; and the problems of biocompatibility, infection risk and secondary operation removal of the implant for a long time are solved.

Description

Technical Field

[0001] This invention relates to the technical field of biomedical materials, and in particular to a fixation material for medical device implants and its preparation method. Background Technology

[0002] Breast implants, including those used in breast reconstruction and augmentation surgery, require stable and safe implants after surgery. Postoperative stability and safety are crucial factors in determining surgical success, necessitating a reliable fixation material to stabilize them in the pre-defined anatomical position and prevent complications such as displacement, rotation, deformation, or "double bubble" deformity. Traditional fixation methods primarily rely on the formation and suturing of the tissue capsule; however, this approach has significant limitations: First, operational complexity: manual suturing of implants is time-consuming, requires high surgical skill, and stress concentration at suture points can lead to tissue cutting or fixation failure. Insufficient mechanical fixation between the implant and the surrounding tissue can easily result in displacement, rotation, or deformation, affecting aesthetic outcomes. Second, biocompatibility and integration: some materials may trigger excessive inflammatory responses or fibrous capsule contracture, resulting in poor integration with surrounding breast tissue and unsatisfactory long-term stability. Furthermore, the surgical wound provides a breeding ground for bacterial biofilm formation, and implant-related infections, especially delayed infections, have become a major clinical challenge. While some existing non-absorbable synthetic meshes or biomaterials can provide immediate support, their permanent presence may lead to long-term foreign body reactions, chronic inflammation, or even tissue erosion. Thirdly, poor morphological adaptability: traditional materials may struggle to perfectly conform to the complex three-dimensional curvature between the implant and the chest wall, resulting in incomplete fixation or the presence of cavities.

[0003] The long-term efficacy of orthopedic implants (such as artificial joint prostheses, spinal internal fixation systems, and trauma internal fixation devices) rests on two major clinical goals: achieving stable bone-implant integration and preventing implant-related infections. Current clinical practice primarily addresses these goals through two technical approaches: one is surface functionalization of the implant, including porous metal coatings and bioceramic coatings, which aim to achieve biological fixation by promoting bone ingrowth. However, this strategy lacks the ability to control the risk of perioperative infection, and its surface structure is easily affected by instrument scraping or soft tissue contamination during surgery, which may impair the potential for bone integration. The other approach is perioperative antibiotic intervention, including systemic intravenous administration and local application of antibiotic-loaded bone cement. The former carries risks of systemic drug distribution, insufficient target site concentration, and the selection of drug-resistant bacteria. While the latter can release drugs locally, the polymethyl methacrylate carrier is non-degradable, and the drug release kinetics exhibit a burst followed by a sharp drop, making it unsuitable for uncemented prostheses or most internal fixation scenarios.

[0004] During the critical perioperative window for establishing a stable interface between the implant and the host bone, existing technologies reveal a prominent clinical contradiction: surface design to promote osseointegration and interventions to prevent infection are disconnected in terms of time, space, and carrier. Particularly noteworthy is that the implantation procedure itself disrupts the local microvascular network of the implantation bed, creating temporary areas of weakened blood supply. Simultaneously, the implant, as a foreign surface, initiates a competitive process of protein adsorption and bacterial colonization within minutes of contact with the tissue. Current surface modification techniques or systemic drug delivery protocols cannot provide a controlled local microenvironment at the implant-bone interface at the moment of implantation to simultaneously achieve mechanical isolation of contaminants and continuous drug intervention. Even the most advanced implant surface drug-eluting coating technologies face complex challenges due to their bonding with the prosthesis, including coating adhesion strength, intraoperative wear and failure, and the potential interference of degradation products with the bone healing process.

[0005] To overcome these shortcomings, an ideal tissue fixation material must possess two core characteristics: excellent antibacterial and anti-inflammatory efficacy and bioabsorbability. While some absorbable materials on the market can be degraded and absorbed within months, their inherent hydrophobicity and bioinertness make them ineffective at inhibiting microbial adhesion and colonization. Their antibacterial function typically relies on postoperative systemic antibiotic administration, failing to provide continuous, localized active protection at the implantation site. Furthermore, the degradation rate of these materials often does not match the tissue regeneration process, potentially leading to premature loss of mechanical support. The degradation process may also create an acidic microenvironment, hindering tissue integration and exacerbating inflammatory responses.

[0006] Therefore, there is an urgent need in this field for a novel fixation material specifically designed for implants that offers excellent initial stability and long-term biointegration while greatly simplifying surgical procedures. Simultaneously, it should provide broad-spectrum, long-lasting antibacterial and anti-inflammatory protection, reducing the risk of infection and inflammation. Furthermore, its degradation kinetics should be synchronized with the ingrowth of new tissue and the replacement of mechanical properties, ensuring that after complete absorption, it is replaced by healthy autologous tissue, achieving true bio-fixation without any risk of long-term foreign body residue. Summary of the Invention

[0007] To address the shortcomings of existing technologies, this invention discloses a fixation material for medical device implants. It uses a biodegradable polymer as the base material and loads a drug-loaded coating containing antibiotics and anti-inflammatory substances on the surface, enabling sustained drug release. It has excellent antibacterial and anti-inflammatory properties and is fully absorbable in a biological environment, solving the long-standing problems of biocompatibility, infection risk, and secondary surgical removal faced by implant materials.

[0008] The specific technical solution is as follows:

[0009] A fixation material for medical device implants includes a biodegradable polymer substrate with a porous structure and a drug-releasing coating uniformly loaded on the biodegradable polymer substrate;

[0010] The pores in the biodegradable polymer are arranged in a uniform, staggered array. The pores are circular with a diameter of 0.5 to 3.0 mm and a spacing of 0.5 to 3.0 mm between adjacent pores.

[0011] The drug-releasing coating contains antibiotics;

[0012] The antibiotics include one or more of rifampin, minocycline hydrochloride, vancomycin, tobramycin, and gentamicin.

[0013] Preferably, the drug sustained-release coating further contains anti-inflammatory substances;

[0014] The anti-inflammatory substances include one or more of gallic acid, dexamethasone, and triamcinolone.

[0015] Further optimization:

[0016] The antibiotic loading capacity is 100~300 µg / cm³. 2 ;

[0017] More preferably, the antibiotic is selected from rifampin and minocycline hydrochloride;

[0018] The drug loading of rifampin is 90~110 µg / cm³. 2 ;

[0019] The drug loading of minocycline hydrochloride is 90–110 µg / cm³. 2 .

[0020] Further optimization:

[0021] The drug loading of the anti-inflammatory substance is 0.1~30 µg / cm³. 2 ;

[0022] More preferably, the anti-inflammatory substance is selected from gallic acid.

[0023] In this invention, the biodegradable polymer with a porous structure is obtained by creating pores in a nonwoven fabric made of the biodegradable polymer, or by weaving a thread made of the biodegradable polymer into a mesh structure.

[0024] Preferably, the suture material is selected from one or more of polyglycolic acid suture, polyglycolic acid-lactide copolymer suture, polydioxanone suture, polycaprolactone suture, polylactic acid and its stereoisomers suture; the nonwoven fabric is selected from one or more of polylactic acid nonwoven fabric, polyglycolic acid nonwoven fabric, polylactic acid-lactide copolymer nonwoven fabric, polycaprolactone nonwoven fabric, collagen / gelatin nonwoven fabric, chitin / chitosan nonwoven fabric.

[0025] When choosing nonwoven fabric as the substrate for perforation, the inventors systematically screened and evaluated nonwoven fabric samples with different processes and basis weights using polylactic acid (PLA) as raw material. The screening covered PLA nonwoven fabrics prepared using various mainstream processes such as spunlace, hot-air bonding, and spunbond bonding, with basis weights ranging widely from 15 to 90 g / m². Through subjective tactile evaluation and comparative analysis of objective physical properties of the series of samples, a significant correlation was found between the material's softness and basis weight and molding process. The spunlace process imparts both good bulkiness and fiber flexibility to the nonwoven fabric, while the hot-air bonding process provides a uniquely soft touch through thermal fusion bonding.

[0026] Based on considerations of the balance between the softness and overall performance of nonwoven fabrics, further optimization was performed:

[0027] The nonwoven fabric prepared from the biodegradable polymer is selected from spunlace PLA nonwoven fabric with a basis weight of 30-40 g / m² or hot-air PLA nonwoven fabric with a basis weight of 20-30 g / m².

[0028] Further optimization:

[0029] The nonwoven fabric prepared from the biodegradable polymer is selected from spunlace PLA nonwoven fabric with a basis weight of 35 g / m² or hot-air PLA nonwoven fabric with a basis weight of 22 g / m².

[0030] Experiments revealed that spunlace PLA nonwoven fabric with a basis weight of 35 g / m² exhibits a delicate and soft touch while providing sufficient mechanical support; hot-air PLA nonwoven fabric with a basis weight of 22 g / m² achieves ultra-high softness due to its lightweight and uniquely soft properties. These two specific specifications of polylactic acid nonwoven fabric were identified as the optimal substrates for achieving the objectives of this invention.

[0031] This material possesses a porous mesh structure with optimized pore size range, porosity, and connectivity, making it suitable for the migration, proliferation, and extracellular matrix deposition of host fibroblasts and vascular endothelial cells. This active tissue ingrowth mechanism allows the material to transform from a 'physical barrier' into a 'living tissue integration layer,' forming a strong biological anchor with surrounding breast adipose tissue, glands, and muscle tissue, achieving a smooth transition from short-term mechanical fixation to long-term biological fixation. The inventors prepared samples with different pore sizes (0.5–3.0 mm), different pore shapes (circular, square, rhomboid), different pore spacings (0.5–3.0 mm), and different array arrangements (uniform and staggered arrays). Evaluation through freehand tensile deformation tests revealed that the perforation parameters significantly affect the local stress concentration and overall ductility of the material. Simultaneously, by simulating the loading of a specific product, the impact of different pore structures on product fixation and ease of use was evaluated. The comprehensive evaluation results show that circular holes, due to their most uniform stress distribution at the edges, can effectively suppress crack propagation, thus providing optimal deformation controllability during manual stretching.

[0032] In a specific parameter combination, a further optimization is made with a pore size of 1.0 mm, a spacing of 1.0 mm between adjacent pores, and an alternating arrangement design. This configuration ensures that the material has sufficient air permeability and a low initial modulus, while the alternating arrangement effectively disperses stress, giving the nonwoven fabric high ductility and tear resistance. In addition, the pore structure of this density provides an ideal anchoring point for loaded products, achieving a good balance between loading stability and ease of use.

[0033] When the substrate is obtained by weaving a mesh structure from a thread made of a biodegradable polymer, it is further preferred that the thread made of the biodegradable polymer is selected from PGA 3-0~6-0 medical sutures; and the weaving machine used is a warp knitting machine.

[0034] In this invention, the medical device implant includes a breast prosthesis implant or an orthopedic implant;

[0035] The breast implants include silicone implants, saline implants, smooth implants, textured implants, round implants, or anatomical implants;

[0036] The orthopedic implants include artificial hip joints, knee joint prostheses, spinal interbody fusion devices, bone plates, or screws.

[0037] When the implant is an orthopedic implant, the fixation material disclosed in this invention is a temporary functional carrier independent of the implant itself. Its inner surface can closely adhere to the orthopedic implant, while its outer surface is designed as a functional interface that facilitates implantation. The antibiotics and anti-inflammatory substances loaded in the fixation material serve as therapeutic agents, and their degradation cycle is precisely designed to ensure complete absorption within 6 weeks to 6 months post-surgery, perfectly covering the infection risk period. Through this design, this invention achieves the spatiotemporal orderly integration of three major functions: acting as a physical barrier to prevent interface contamination at the moment of implantation; continuously releasing therapeutic agents as a local drug reservoir during the critical post-operative period; and completely degrading without interfering with long-term bone integration after fulfilling its purpose. This provides a complete solution for various orthopedic implant surgeries, from "implant protection" to "post-operative treatment" to "self-removal." This invention also discloses a method for preparing the aforementioned fixation material for medical device implants, including:

[0038] Step 1: Mix antibiotics, selectively added anti-inflammatory substances, absorbable polymers, and mixed solvents to obtain a spraying solution;

[0039] Step 2: Using a biodegradable polymer with a porous structure as the substrate, the spraying solution prepared in Step 1 is sprayed onto the surface of the substrate using an air spraying process.

[0040] Step 3: Shape the sprayed substrate and cut it into a shape that matches the implant of the medical device to obtain the fixation material.

[0041] In step one:

[0042] Preferably, the antibiotic is selected from rifampin and minocycline hydrochloride;

[0043] Preferably, the mass ratio of rifampin to minocycline hydrochloride in the antibiotic is (0.3~3.0):1; more preferably, it is 1:1.

[0044] Preferably, the anti-inflammatory active substance is selected from gallic acid;

[0045] Preferably, the absorbable polymer is selected from one or more of polyglycolic acid, polylactide, polyglycolic acid-co-lactide, polyglycolic acid-co-caprolactone, and tyrosine polyarylate.

[0046] Preferably, the mass ratio of antibiotic to absorbable polymer is 1:(1~4); by adjusting the mass ratio of the two, the release rate of antibiotic can be controlled to achieve different sustained-release effects to match different application scenarios.

[0047] Preferably, the mass ratio of the anti-inflammatory substance to the antibiotic is (0.01~1):1; more preferably, it is 0.1:1. Preferably, the concentration of the antibiotic in the spray solution is 1~20 g / mL; more preferably, it is 5~10 g / mL.

[0048] Preferably, the mixed solvent is selected from hexafluoroisopropanol and ethyl acetate, and the volume percentage of hexafluoroisopropanol in the mixed solvent is not less than 10%.

[0049] More preferably, the volume percentage of hexafluoroisopropanol in the mixed solvent is 10-90%; more preferably, the volume percentage of hexafluoroisopropanol is 50-90%.

[0050] Preferably, in step two, the air spraying:

[0051] The spray flow rate is selected from 0.4~1.0 mL / min, the feed speed is selected from 6~10 mm / s, the pressure of air path 1 and air path 2 are independently selected from 0.1~1.0 MPa, the pressure of liquid path is selected from 0.05~0.2 MPa, and the nozzle height is selected from 20~40 mm.

[0052] Preferably, in step three, the fixation material provided by the present invention has high flexibility in size and shape. When used to fix breast implants, the material can be prefabricated in various specifications, such as, but not limited to, rectangular, crescent-shaped, anatomical, or bra-shaped, and is available in different size series such as small (10cm×15cm or smaller), medium (15cm×20cm), and large (20cm×25cm or larger) to accommodate implants with different capacities and base widths, while facilitating rapid selection during surgery.

[0053] More preferably, the fixation material possesses excellent intraoperative cutability and three-dimensional plasticity. Surgeons can easily cut and shape the sheet-like material of this invention into any desired form, depending on the type (round or anatomical), volume, base width of the selected breast implant, and the patient's unique thoracic anatomy. Whether used to fully cover the implant or to reinforce key areas (such as the lower pole of the implant to support the inframammary fold, or to cover the upper pole to prevent rotation of anatomical implants), precise, stable, and personalized fixation can be achieved.

[0054] The fixation material involved in this invention is applicable to various breast implants commonly used in clinical practice, including round or anatomical (teardrop) implants filled with silicone or saline, as well as products with different surface properties such as smooth, textured, or nano-textured surfaces. This material can be flexibly cut or shaped into various sizes and shapes (such as rectangular, crescent-shaped, or customized contours) according to intraoperative needs to perfectly fit the shape and volume of different implants, providing reliable support for their initial stability and long-term integration within the breast environment, while greatly simplifying the surgical procedure.

[0055] When used for fixing orthopedic implants, the material can also be prefabricated in various sizes, including small sizes (10cm×15cm or smaller) suitable for implants with a diameter ≤ 6mm (such as cancellous bone screws), medium sizes (15cm×20cm) suitable for implants with a cross-section of 6-15mm (such as bone plates, small joint prostheses), and large sizes (20cm×25cm or larger) suitable for implants with a cross-section > 15mm (such as hip and knee joint prostheses), etc., to accommodate implants of different shapes and sizes, and to facilitate rapid selection during surgery.

[0056] The fixation material involved in this invention employs a highly adaptable pre-formed design, perfectly conforming to the shapes of various orthopedic implants commonly used in clinical practice, including but not limited to artificial joint prostheses (acetabular cups, femoral stems, tibial supports), spinal internal fixation systems (pedicle screws, fusion devices), and trauma internal fixation devices (bone plates, intramedullary nails). This material possesses excellent flexibility and manipulation, enabling rapid and tight wrapping with the implant through simple techniques while maintaining structural integrity, without displacement or damage during implantation. This design significantly simplifies the surgical procedure, reduces intraoperative preparation time, and provides an efficient and convenient intraoperative protection solution for various orthopedic implant surgeries.

[0057] Compared with the prior art, the present invention has the following beneficial effects:

[0058] This invention provides a fixation material for breast implants or orthopedic implants, comprising a mesh substrate made of a fully absorbable polymer, and a drug coating loaded with antibiotics such as rifampin and minocycline, and anti-inflammatory substances such as gallic acid, on its surface. By combining the highly effective antibiotic and anti-inflammatory drug-loaded coating with a fully absorbable polymer matrix, multiple synergistic benefits are achieved, significantly improving the safety and reliability of implant fixation.

[0059] Firstly, this invention provides potent and precise local antibacterial and anti-inflammatory protection. By loading antibiotics onto the surface and superficial layer of the material, antibacterial drugs can be released at the surgical site after implantation, forming an effective local bactericidal barrier and greatly reducing the risk of early postoperative infection. Compared to systemic antibiotic treatment, this local administration method can achieve higher drug concentrations in the target area with a lower total drug amount, resulting in a more direct and efficient effect, while reducing the potential risks of systemic side effects and toxicity. The controllable release characteristics of the coating can also adjust the release kinetics according to clinical needs, ensuring continuous protection during the critical window period. Compared to surface coating technologies that only focus on osseointegration, this invention creatively constructs a three-in-one functional paradigm of "intraoperative immediate protection - local continuous treatment - postoperative autonomous removal". Through an independent biodegradable encapsulation system, it provides a physical isolation barrier in the critical early stage of implant placement, effectively preventing interface contamination and mechanical damage to the coating during the surgical procedure. Compared to permanent foreign body carriers such as antibiotic bone cement, this system can be completely biodegraded after fulfilling its mission of local drug release, avoiding the risk of chronic inflammation or secondary infection caused by long-term foreign body retention, and truly achieving the clinical ideal of "zero residue".

[0060] Secondly, this invention completely eliminates the potential for long-term complications caused by residual fixation materials. Its core matrix is ​​composed of a fully absorbable polymer, which, after fulfilling its temporary mechanical support function, can be safely and completely degraded and absorbed in the body through hydrolysis, ultimately being metabolized into water and carbon dioxide and excreted. This characteristic ensures that no permanent synthetic foreign bodies remain in the body, fundamentally avoiding long-term problems such as chronic inflammation, tissue erosion, aggravated capsular contracture, or increased palpability that may result from long-term material retention. It achieves the ideal pathway of "intervention-support-degradation-disappearance," fundamentally improving patient safety.

[0061] Third, this invention achieves precise spatiotemporal control of local drug delivery. Compared to systemic drug delivery, this system delivers high concentrations of antibiotics directly to the implant-bone interface, the most susceptible "target area" for infection, greatly improving local drug bioavailability while avoiding systemic toxicity.

[0062] Fourth, this invention demonstrates exceptional ease of operation during surgery. Its superior cutibility and flexibility allow surgeons to precisely and instantly cut it into the desired shape according to the implant size and the patient's individual anatomy, ensuring a close fit to the irregular curves of the implant surface and tissue cavities. The material maintains its structural integrity even in humid environments and possesses suitable initial adhesion and an operating window, enabling rapid temporary fixation upon tissue contact while allowing for fine-tuning of its position. This significantly simplifies the fixation process, reduces reliance on complex suturing techniques, and effectively shortens surgical time.

[0063] The preparation method disclosed in this invention optimizes parameters such as the basis weight and pore structure of the substrate, and the resulting fixation material has excellent flexibility, stretchability and mechanical compatibility similar to human tissue. Attached Figure Description

[0064] Figure 1 A photograph of the fixation material prepared in Example 1;

[0065] Figure 2 Microscopic image of the fixation material prepared in Example 1;

[0066] Figure 3 A photograph of the fixation material prepared in Example 7;

[0067] Figure 4 The tensile stress-strain curves of the nonwoven fabrics in Examples 1-3 and the uncoated nonwoven fabric are shown.

[0068] Figure 5 The in vitro release curves of rifampin in PBS buffer are shown for the immobilization materials prepared in Examples 1-3, respectively.

[0069] Figure 6 The in vitro release curves of minocycline hydrochloride in PBS buffer are shown for the immobilization materials prepared in Examples 1-3, respectively.

[0070] Figure 7 The in vivo release curves of rifampin in SD rats from the fixation materials prepared in Examples 1-3 are shown.

[0071] Figure 8 The in vivo release curves of minocycline hydrochloride in SD rats from the fixation materials prepared in Examples 1-3 are shown.

[0072] Figure 9 The images show the antibacterial activity of the fixatives prepared in Examples 1-3 against Staphylococcus aureus, with a nonwoven fabric (marked as blank) without spraying treatment as a comparison. Detailed Implementation

[0073] To make the objectives, technical solutions, and advantages of the embodiments of the present invention clearer, the technical solutions in the embodiments of the present invention will be clearly and completely described below. Where specific conditions are not specified in the embodiments, conventional conditions or conditions recommended by the manufacturer shall apply. Reagents or instruments whose manufacturers are not specified are all conventional products that can be purchased commercially. The features and performance of the present invention will be further described in detail below with reference to the embodiments.

[0074] Example 1

[0075] 1) Cut 35 g / m² PLA spunlace nonwoven fabric into 20cm×20cm pieces and laser-drill circular holes with 1mm hole diameter and 1mm hole spacing in an interlaced array.

[0076] 2) Using a pipette, mix 14 mL of hexafluoroisopropanol and 6 mL of ethyl acetate in a 50 mL centrifuge tube. Weigh 0.084 g of rifampin and minocycline hydrochloride into light-protected 50 mL centrifuge tubes. Weigh 0.168 g of PLGA polymer into a centrifuge tube containing the two antibiotics. Pour the prepared hexafluoroisopropanol and ethyl acetate mixture into the centrifuge tube containing the polymer and antibiotics. Vortex at 1000 rpm for 30 min to dissolve all the solutes, thus preparing a high-release-rate coating solution with a mass ratio of PLGA, rifampin, and minocycline hydrochloride of 2:1:1. Turn on the computer, air pump, and air spraying power. Transfer the spraying solution into a 10 mL syringe and fix it to the syringe pump. Place the cut and perforated PLA nonwoven fabric into the appropriate spraying position. Set the parameters for spraying: spraying flow rate: 0.4 mL / min, feed speed: 6 mm / s, pressure of air path 1 and air path 2: 0.2 MPa, pressure of liquid path: 0.05 MPa, nozzle height: 30 mm.

[0077] 3) Place the sprayed PLA nonwoven fabric in a vacuum drying oven and vacuum dry for 3 hours. Then, use a sealing machine to hot-press the edges and cut it into 8cm × 20cm shapes to obtain the fixation material for implants.

[0078] Example 2

[0079] The preparation process is basically the same as in Example 1, with the only difference being:

[0080] In step 2), the mass of PLGA polymer is replaced with 0.392 g to prepare a medium-release rate coating solution with a mass ratio of PLGA, rifampin and minocycline hydrochloride of 4.7:1:1.

[0081] Example 3

[0082] The preparation process is basically the same as in Example 1, with the only difference being:

[0083] In step 2), the mass of PLGA polymer is replaced with 0.588 g to prepare a low-release-rate coating solution with a mass ratio of PLGA, rifampin and minocycline hydrochloride of 7:1:1.

[0084] Example 4

[0085] The preparation process is basically the same as in Example 1, with the only difference being:

[0086] In step 1), the 35 g / m² PLA spunlace nonwoven fabric is replaced with 22 g / m² PLA hot-air nonwoven fabric, and the hole structure on the nonwoven fabric is exactly the same as in Example 1.

[0087] Example 5

[0088] The preparation process is basically the same as in Example 2, with the only difference being:

[0089] In step 1), the 35 g / m² PLA spunlace nonwoven fabric is replaced with 22 g / m² PLA hot-air nonwoven fabric, and the hole structure on the nonwoven fabric is exactly the same as in Example 2.

[0090] Example 6

[0091] The preparation process is basically the same as in Example 3, with the only difference being:

[0092] In step 1), the 35 g / m² PLA spunlace nonwoven fabric is replaced with 22 g / m² PLA hot-air nonwoven fabric, and the hole structure on the nonwoven fabric is exactly the same as in Example 3.

[0093] Example 7

[0094] 1) PGA 5-0 medical sutures with a diameter of 0.15mm are woven into a mesh structure sheet using a warp knitting machine. The knitting machine is a tongue needle warp knitting machine, model number E16, using a single comb for knitting, and the knitting process is three-needle open warp forging.

[0095] Steps 2) to 3) are exactly the same as in Example 1.

[0096] Example 8

[0097] 1) PGA 5-0 medical sutures with a diameter of 0.15mm are woven into a mesh structure sheet using a warp knitting machine. The knitting machine is a tongue needle warp knitting machine, model number E16, using a single comb for knitting, and the knitting process is three-needle open warp forging.

[0098] Steps 2) to 3) are exactly the same as in Example 2.

[0099] Example 9

[0100] 1) PGA 5-0 medical sutures with a diameter of 0.15mm are woven into a mesh structure sheet using a warp knitting machine. The knitting machine is a tongue needle warp knitting machine, model number E16, using a single comb for knitting, and the knitting process is three-needle open warp forging.

[0101] Steps 2) to 3) are exactly the same as in Example 3.

[0102] Example 10

[0103] The preparation process is basically the same as in Example 1, with the only difference being:

[0104] In step 2), 0.008 g of gallic acid is weighed and placed together with rifampicin and minocycline hydrochloride in a 50 mL centrifuge tube protected from light to prepare a coating solution with a mass ratio of PLGA, rifampicin, minocycline hydrochloride and gallic acid of 20:10:10:1.

[0105] Comparative Example 1

[0106] The preparation process is basically the same as in Example 1, with the only difference being:

[0107] In step 1), circular holes are laser-drilled with a 5mm aperture, a 5mm hole spacing, and an interlaced array.

[0108] Comparative Example 2

[0109] The preparation process is basically the same as in Example 1, with the only difference being:

[0110] In step 1), square holes are laser-drilled with a side length of 1mm and a spacing of 1mm, in an alternating array.

[0111] Comparative Example 3

[0112] The preparation process is basically the same as in Example 1, with the only difference being:

[0113] In step 1), diamond-shaped holes are laser-drilled with a side length of 1mm and a spacing of 1mm, arranged in an alternating row and row array.

[0114] Comparative Example 4

[0115] The preparation process is basically the same as in Example 1, with the only difference being:

[0116] In step 1), circular holes are made by laser drilling with a diameter of 1 mm and a spacing of 1 mm, arranged in a neat array.

[0117] Evaluation of PLA substrates with different pore structures prepared in Example 1 and Comparative Examples 1-4 through freehand tensile deformation tests revealed that perforation parameters significantly affect local stress concentration and overall ductility of the material. Simultaneously, the impact of different pore structures on product fixation and ease of use was evaluated by simulating loading specific products. Comprehensive evaluation results show that circular pores, due to their most uniform edge stress distribution, effectively suppress crack propagation, thus providing optimal deformation control during freehand stretching. In contrast, rhomboid pores, square pores, and larger-diameter circular pores exhibit significant deformation and are difficult to restore to their original shape. Among specific parameter combinations, the staggered arrangement design with a pore diameter of 1 mm and a pore spacing of 1 mm demonstrates the best overall performance: this configuration ensures sufficient air permeability and a low initial modulus while the staggered arrangement effectively disperses stress, giving the nonwoven fabric high ductility and tear resistance; furthermore, this density of pore structure provides ideal anchoring points for loaded products, achieving a good balance between loading stability and ease of use.

[0118] Performance testing:

[0119] I. Mechanical property testing

[0120] 1.1 Piercing Strength Test

[0121] According to GB / T 19976-2005, the bursting strength of the fixing materials prepared in Examples 1-3 and Example 10, as well as the perforated nonwoven fabric without spraying treatment (hereinafter referred to as nonwoven fabric), was tested. The sample was clamped in the circular sample holder of the fixed base, and the spherical push rod was pushed vertically towards the sample at a constant moving speed, causing the sample to deform until it broke. The bursting strength, bursting elongation, and bursting elongation data are listed in Table 1 below.

[0122] Table 1

[0123]

[0124] The comparison revealed that the breaking strength of the coated fixative material prepared after spraying was significantly higher than that of the uncoated sample. Meanwhile, the breaking strength of the coated fixative material with added gallic acid was similar to that of the uncoated material.

[0125] 1.2 Tear Strength Test

[0126] Tear strength tests were conducted on the fixing materials prepared in Examples 1-3 and Example 10, as well as the perforated nonwoven fabric without coating treatment (hereinafter referred to as nonwoven fabric), according to GB / T 3917.3-2009. A trapezoid was drawn on the sample, and the clamps of the strength tester were used to clamp the two non-parallel sides of the trapezoid. A continuously increasing force was applied to the sample, causing the tear to propagate along the width of the sample. The maximum tear strength was measured and listed in Table 2 below.

[0127] Table 2

[0128]

[0129] The comparison revealed that the tear strength of the coated fixation material prepared after spraying was significantly higher than that of the uncoated sample. Meanwhile, the maximum tear strength of the coated fixation material with added gallic acid was similar to that of the uncoated material.

[0130] 1.3 Tensile stress test

[0131] Tensile stress was tested on the fixed materials prepared in Examples 1-3 and Example 10, as well as the perforated nonwoven fabric (hereinafter referred to as nonwoven fabric) without spraying treatment, using a universal tensile testing machine. A sample with a width of 10 mm and a length of 100 mm was clamped in a fixture. A 10 N sensor, a speed of 10 mm / min, and a sample size of 20 mm were used to test the tensile stress. The tensile stress / strain curves are shown below. Figure 4 As shown, the specific data is listed in Table 3 below.

[0132] Table 3

[0133]

[0134] The comparison revealed that the tensile stress of the coated fixative material prepared after spraying was significantly higher than that of the uncoated sample. Meanwhile, the tensile stress of the coated fixative material with added gallic acid was similar to that of the uncoated material.

[0135] The mechanical properties of this material are matched to those of natural breast soft tissue. This 'mechanical compatibility' avoids undue pressure on the implant or surrounding tissue due to excessive rigidity, and also prevents ineffective support due to excessive softness. It adapts to the movement of the breast itself (such as breathing and changes in body position), providing dynamic and lasting stability and giving patients a more natural postoperative feel.

[0136] II. Drug Release Test

[0137] 2.1 In vitro release experiment:

[0138] The encapsulation materials prepared in Examples 1-3 were cut into samples with a diameter of 1 cm and placed in 20 mL of PBS buffer solution (pH=7.4). The samples were then shaken (110 rpm) in a 37°C incubator. Samples were taken at 0.5, 2, 8, 24, and 48 h for UV spectrophotometry to determine the drug concentration released into the medium. The percentage of drug release at different times was determined based on the drug's standard curve.

[0139] Among them, the in vitro release curve of rifampin is as follows: Figure 5 As shown, the in vitro release curve of minocycline hydrochloride is as follows: Figure 6 As shown in the figure. Observations revealed that rifampin and minocycline hydrochloride had similar in vitro release characteristics, releasing approximately 20%, 50%, and 80% of the drug respectively after 2 hours in vitro, and the release continued for up to two days.

[0140] 2.2 In vivo release experiment:

[0141] SD rats (healthy, adult, weighing 200-300 g) were selected for in vivo drug release experiments. The fixation materials prepared in Examples 1-3 were cut into 1 cm diameter samples using a 1 cm diameter mold. Subcutaneous implantation was performed through a 1-2 cm chest incision. The 1 cm diameter sheet of fixation material was placed directly into the subcutaneous tissue of the chest, and the subcutaneous tissue was sutured with silk sutures, ensuring a neat and aligned wound to avoid dead space. After implantation, the wound was wiped clean with iodine swabs to remove blood and tissue debris. Rats were sacrificed on days 1, 3, 5, 7, and 14. The substrate was removed, its surface moisture was dried, and excess tissue was removed. Antibiotics on the substrate surface were washed away with a 1:1 (v / v) methanol-water mixture, and the residual drug loading in the coating was detected using a UV-Vis spectrophotometer.

[0142] Among them, the in vivo release curve of rifampin is as follows: Figure 7 As shown, the in vitro release curve of minocycline hydrochloride is as follows: Figure 8 As shown. Observations revealed that the in vitro release of rifampin was similar to that of minocycline hydrochloride. The fixation material prepared in Example 1 released more than 80% of the drug after one day, the fixation material prepared in Example 2 released about 50% after one day, and the encapsulation material prepared in Example 3 released about 10% after one day. All three materials continued to release the drug for up to 14 days.

[0143] The standard curve is obtained by plotting it in the following way:

[0144] High performance liquid chromatography (HPLC) conditions: Column: C18 column; Mobile phase: Methanol: Acetonitrile: 0.075 mol / L potassium dihydrogen phosphate solution: 1 mol / L citric acid (30:30:36:4) as the mobile phase, pH adjusted to approximately 7 with 10 mol / L NaOH; Detection wavelength: 254 nm; Injection volume: 10 µL.

[0145] Method for establishing standard curves (methanol / water solution) for rifampin and minocycline hydrochloride: Weigh 10 mg of minocycline hydrochloride and rifampin separately on an electronic balance and place them in centrifuge tubes. Add them to a mixed solution of methanol and deionized water (volume ratio 1:1), stir to dissolve, and prepare a solution with a mass concentration of 1 mg / mL. Quantitatively dilute the solution to the following mass concentrations: 0.5, 1, 5, 10, 15, 20 µg / mL. Detect the absorbance values ​​of minocycline hydrochloride and rifampin at 254 nm and plot the standard curves.

[0146] Method for establishing standard curves (PBS solution) for rifampin and minocycline hydrochloride: Weigh 10 mg of minocycline hydrochloride and rifampin separately on an electronic balance and place them in centrifuge tubes. Add PBS buffer solution (pH=7.4), stir to dissolve, and prepare a solution with a mass concentration of 1 mg / mL. Quantitatively dilute the solution to the following mass concentrations: 0.5, 1, 5, 10, 15, 20 µg / mL. Detect the absorbance values ​​of minocycline hydrochloride and rifampin at 254 nm and plot the standard curves.

[0147] III. Antibacterial Test

[0148] Prepare the bacterial culture: Pick a single colony with an inoculation loop and place it in a 15 mL centrifuge tube containing 5 mL of sterile LB liquid medium (or add 15 mL of medium to a 50 mL centrifuge tube). Incubate overnight (15-20 h) at 37°C and 150 rpm on a shaker.

[0149] Dilution and Spreading: Dilute Staphylococcus aureus solution to 0.5 McFarland turbidity (10⁸ CFU / mL) with sterile physiological saline / LB liquid medium. Spread Staphylococcus aureus solution evenly on LB agar plates using a sterile cotton swab and let stand for 5 min. After standing, place a 1 cm diameter piece of the coating material prepared in Example 1 and a perforated nonwoven fabric without spraying treatment as a blank control. Then, let the agar plates stand in a clean bench for 30 min. Place the plates in a self-sealing bag (unsealed) and invert them in a 37°C oven for overnight incubation. Observe the size of the inhibition zone.

[0150] Antibacterial results such as Figure 9 As shown, compared with the blank sample, the coated PLA nonwoven fabric has significant antibacterial properties. Meanwhile, the antibacterial effect of the coating with gallic acid added in Example 10 is similar to that of the coating in Example 1.

[0151] The above description is merely a specific embodiment of the present invention, but the scope of protection of the present invention is not limited thereto. The specific examples used above to illustrate the present invention are only for the purpose of helping to understand the present invention and are not intended to limit the present invention. Those skilled in the art to which this invention pertains can make several simple deductions, modifications, substitutions, or combinations based on the concept of the present invention. These deductions, modifications, substitutions, or combinations also fall within the scope of the claims of the present invention.

Claims

1. A fixation material for medical device implants, characterized in that, It includes a biodegradable polymer substrate with a porous structure and a drug sustained-release coating uniformly loaded on the biodegradable polymer substrate; The pores in the biodegradable polymer are arranged in a uniform, staggered array. The pores are circular with a diameter of 0.5 to 3.0 mm and a spacing of 0.5 to 3.0 mm between adjacent pores. The drug-releasing coating contains antibiotics; The antibiotics include one or more of rifampin, minocycline hydrochloride, vancomycin, tobramycin, and gentamicin.

2. The fixation material for medical device implants according to claim 1, characterized in that, The drug-release coating also contains anti-inflammatory substances; The anti-inflammatory substances include one or more of gallic acid, dexamethasone, and triamcinolone.

3. The fixation material for medical device implants according to claim 2, characterized in that: The antibiotic loading capacity is 100~300 µg / cm³. 2 ; The drug loading of the anti-inflammatory substance is 0.1~30 µg / cm³. 2 .

4. The fixation material for medical device implants according to claim 1, characterized in that: The porous biodegradable polymer is obtained by creating holes in a nonwoven fabric made from the biodegradable polymer, or by weaving a thread made from the biodegradable polymer into a mesh structure. The suture material is selected from one or more of the following: polyglycolic acid suture, polyglycolic acid-lactide copolymer suture, polydioxanone suture, polycaprolactone suture, polylactic acid and its stereoisomer suture. The nonwoven fabric is selected from one or more of the following: polylactic acid nonwoven fabric, polyglycolic acid nonwoven fabric, polylactic acid-glycolic acid copolymer nonwoven fabric, polycaprolactone nonwoven fabric, collagen / gelatin nonwoven fabric, and chitin / chitosan nonwoven fabric.

5. The method for preparing a fixation material for medical device implants according to claim 4, characterized in that: The nonwoven fabric prepared from the biodegradable polymer is selected from spunlace PLA nonwoven fabric with a basis weight of 30-40 g / m² or hot-air PLA nonwoven fabric with a basis weight of 20-30 g / m²; the filament prepared from the biodegradable polymer is selected from PGA3-0 ~ 6-0 medical sutures. The knitting machine used is a warp knitting machine.

6. The fixation material for medical device implants according to claim 1, characterized in that, The medical device implants include breast implants or orthopedic implants; The breast implants include silicone implants, saline implants, smooth implants, textured implants, round implants, or anatomical implants; The orthopedic implants include artificial hip joints, knee joint prostheses, spinal interbody fusion devices, bone plates, or screws.

7. A method for preparing a fixation material for medical device implants according to any one of claims 1 to 6, characterized in that, include: Step 1: Mix antibiotics, selectively added anti-inflammatory substances, absorbable polymers, and mixed solvents to obtain a spraying solution; Step 2: Using a biodegradable polymer with a porous structure as the substrate, the spraying solution prepared in Step 1 is sprayed onto the surface of the substrate using an air spraying process. Step 3: Shape the sprayed substrate and cut it to obtain the fixation material for medical device implants.

8. The method for preparing a fixation material for medical device implants according to claim 7, characterized in that, In step one: The absorbable polymer is selected from one or more of polyglycolic acid, polylactide, polyglycolic acid-co-lactide, polyglycolic acid-co-caprolactone, and tyrosine polyarylate. The mixed solvent is selected from hexafluoroisopropanol and ethyl acetate, and the volume percentage of hexafluoroisopropanol in the mixed solvent is not less than 10%.

9. The method for preparing a fixation material for medical device implants according to claim 7, characterized in that, In step one: The mass ratio of antibiotics to absorbable polymers is 1:(1~4). The mass ratio of anti-inflammatory substances to antibiotics was (0.01~1):1; The concentration of antibiotics in the spray solution is 1~20 g / mL; In the mixed solvent, hexafluoroisopropanol accounts for 10-90% by volume.

10. The method for preparing a fixation material for medical device implants according to claim 7, characterized in that, In step two, the air spraying: The spray flow rate is selected from 0.4~1.0 mL / min, the feed speed is selected from 6~10 mm / s, the pressure of air path 1 and air path 2 are independently selected from 0.1~1.0 MPa, the pressure of liquid path is selected from 0.05~0.2 MPa, and the nozzle height is selected from 20~40 mm.

Images