Preparation method and application of a multi-stage biomimetic nerve membrane structure nerve conduit

Abstract

The application discloses a preparation method and application of a nerve conduit with a multistage biomimetic nerve membrane structure and belongs to the technical field of tissue engineering materials. The nerve conduit with the multistage biomimetic nerve membrane structure comprises a biomimetic inner membrane support with a secondary guiding structure and a biomimetic outer membrane support with a double-layer structure. In the nerve conduit with the multistage biomimetic nerve membrane structure, the filler for playing a growth guiding role has a multistage guiding structure, the primary guiding is that cells are directionally migrated along the fiber direction in the directional fiber membrane, are elongated in shape on the fiber and are accelerated in differentiation, and the secondary guiding is that directional microchannels constrain the growth of nerve fibers, promote the maturation of bundling and the alignment of the growth direction. The application solves the problem of insufficient contact guiding efficiency existing in the prior art and overcomes the contradiction between the filling density of guiding fibers, the growth space and the guiding effect.

Description

Technical Field

[0001] This invention relates to the field of tissue engineering materials technology, and in particular to a method for preparing and applying a multi-level biomimetic nerve membrane structure for nerve conduits. Background Technology

[0002] Peripheral nerve injury (PNI) is a common neurological disorder, usually caused by traumatic events, surgical resection, or iatrogenic injury. PNI often leads to motor or sensory nerve dysfunction accompanied by neuropathic pain. It has a high incidence and poor prognosis, severely impacting patients' quality of life and imposing a huge social and family burden. Clinically, for nerve defects with short gaps (<8mm), end-to-end anastomosis is often used for repair. However, for nerve defects with long gaps (>2-3cm), autologous transplantation is the "gold standard." However, its clinical application is limited due to numerous problems such as secondary surgery, limited availability of donor nerves, functional damage at the donor site, size mismatch, and the risk of neuroma formation. Currently, the development of nerve guide conduits (NGCs) provides an attractive alternative for the repair of long-gap nerve defects. Although more than 10 nerve bridging products have entered the clinical application stage both domestically and internationally, these products are mostly used for the repair of small, straight, short-gap nerve defects. For functional repair of long-distance severe nerve defects or complex nerve defects spanning joints, currently the only clinically available products worldwide are decellularized allogeneic grafts (from Axogen, Inc., USA). China Zhongda Medical Two products (acellular allogeneic nerve repair materials) possess long-distance repair capabilities. According to research and clinical feedback, these products achieve superior nerve function recovery because they retain a relatively complete nerve membrane system and have a continuous longitudinal guiding structure similar to natural peripheral nerves. However, similar to autologous grafts, their length and diameter are limited by the availability of donors, often resulting in size mismatches and high costs. Furthermore, due to the decellularization process, these products have poor mechanical properties, weak resistance to compression, and lack cross-joint repair capabilities, leading to potential failure during implantation. Therefore, constructing nerve conduits with long-distance and cross-joint repair capabilities, and capable of large-scale production, remains a pressing clinical challenge.

[0003] The successful experience with decellularized allogeneic grafts demonstrates that precisely mimicking the membrane structure of natural peripheral nerves holds promise for solving this clinical challenge from a structural biomimetic perspective. In fact, researchers have already conducted extensive work on simulating the structure and properties of the nerve endometrium, constructing structures such as oriented micro / nanofibrils and microchannels to guide nerve cell orientation and promote nerve regeneration. Among these, Academician Xiaosong Gu's team pioneered representative research on oriented micro / nanofibril structures. In 2005, the team developed a two-component artificial nerve graft composed of a chitosan shell and oriented internal fibers of polyglycolic acid (PGA). This conduit enables nerve reconstruction and functional recovery, and to some extent restores the motor function of the operated limb. The team received approval for clinical use of their product in 2020. Peripheral nerve repair grafts are a continuation of this technological approach. This product also consists of a shell conduit and internal fibers, with the internal fibers made of poly(glycolic acid lactide) (PGLA). This is currently the only nerve conduit product with an internal guiding structure and clinical approval. Similarly, in 2021, Deling Kong's research group developed a two-component nerve conduit composed of a poly(glycolic acid lactide) copolymer (PLCL) shell and poly(p-dioxanone) oriented internal fibers (PDS). This composite conduit demonstrated near-autologous nerve regeneration and functional recovery in repairing an 11mm sciatic nerve defect in rats. However, during surgical implantation and nerve regeneration, it was found that while oriented micro / nanofiber fibers can effectively increase the contact guiding area and achieve cell-oriented guidance, their filling density is extremely difficult to control: slightly higher density leads to reduced lumen porosity, hinders nerve migration, and may trigger excessive inflammatory responses; slightly lower density causes accumulation on one side of the conduit due to gravity, affecting the stability of the guiding structure, leading to guidance failure and axonal mismatch. This risk gradually increases with the length and diameter of the repaired nerve.

[0004] Regarding microchannel structures, in 2012, Ahmet Bozkurt et al. prepared a neural conduit with an open porous structure by molding and freeze-drying porcine collagen. This conduit contained continuous, longitudinal microchannel structures. By seeding Schwann cells into the scaffold, it showed near-autologous graft repair effects in repairing a 20mm sciatic nerve defect in rats. In 2018, Wei Zhu et al. used 3D printing technology to construct a neural conduit with a four-channel structure, achieving effective directional guidance of regenerated nerves in a 4mm sciatic nerve defect model in mice. Microchannel structures are not affected by gravity and therefore do not accumulate, offering advantages in structural stability compared to micro / nanofiber structures. However, their relatively small surface area hinders nerve cell adhesion. Furthermore, the sizes of microchannel structures constructed in current studies are mostly concentrated in the range of hundreds of micrometers to several millimeters, which is far from the typical size of nerve cells (tens of micrometers). Therefore, efficient directional guidance is difficult to achieve, and axonal mismatch problems are also likely to occur. Therefore, there is an urgent need to construct a novel biomimetic structure for the neuroendomembrane that combines high guidance precision, strong guidance stability, and large guidance area, thereby maximizing its contact guidance efficiency.

[0005] In terms of mimicking the structure and characteristics of the epineurium, currently approved neural conduits for clinical use include... Peripheral nerve sleeve tube, Artificial nerve sheaths and similar devices mimic the function of the epineurium, and their structure is that of a hollow conduit. However, clinical application feedback and reports indicate that some products experience failures during use: First, after implantation, the conduit is compressed by surrounding muscles. Due to insufficient initial mechanical properties of the conduit shell or a sharp decline in mechanical properties after degradation, the conduit is prone to rupture or reduction in lumen volume before nerve repair is complete, thus losing its nerve protection function and hindering nerve growth. Second, because the conduit shell lacks a bending-resistant design, a "kick" can easily form in the middle of the conduit during joint movement in cross-joint repairs. This causes compression and deformation of the internal lumen, hindering nerve regeneration and potentially leading to suture slippage from both ends of the conduit, inducing secondary injury. Clinically, to avoid this, joint fixation is often used to prevent relative movement. However, prolonged fixation can easily cause joint stiffness, tissue adhesions, muscle atrophy, and other problems. In recent years, researchers have conducted relevant studies to address these issues, but a mature solution still does not exist. To enhance the compressive strength of nerve conduits, in 2021, Shujun Zhang et al. used a woven silk and magnesium wire inner layer and a freeze-dried silk fibroin and chitosan solution outer layer to obtain a nerve conduit with good compressive strength. Compared with a conduit without a woven inner layer, this composite structure conduit better maintained the three-dimensional morphology of the conduit during the repair of a 10mm sciatic nerve defect in rats, demonstrating better nerve regeneration and functional recovery. In 2023, Yongcong Fang et al. developed a multi-scale, three-layered nerve conduit, in which the 100-micron-level polylactide-caprolactone (PCL) fiber structure endowed the conduit with good mechanical stability. The regenerated nerve showed good sphenoid myelin sheath formation, increased muscle weight, and improved sciatic nerve functional index. To improve the bending resistance of nerve conduits, in 2019, Quan Qi et al. prepared a flexible helical nerve conduit using a receiver rod with a helical structure via electrospinning. This conduit maintained its shape and good flexibility even after three months in vivo. Compared to traditional electrospun nerve conduits, this material has advantages such as lower intraoperative tension and applicability for cross-articular nerve repair, but its regenerative efficacy remains poor. Among clinically used products, only Neuroflex, manufactured by Stryker Corporation in the United States, is available worldwide. TMWhile possessing bending resistance, its maximum bending angle is 60°. For repairing large-angle transarticular nerve defects (such as those at the metacarpophalangeal joints), this angle is insufficient for clinical needs, and there is a lack of relevant product information and clinical evaluation data for reference. Furthermore, it is worth noting that the aforementioned nerve conduits with bending resistance designs all achieve this resistance through the construction of folded or spiral structures. However, these structures are perpendicular to the direction of nerve regeneration, which is detrimental to axonal orientation and extension, further increasing the likelihood of nerve mismatch. Therefore, there is an urgent need to construct a novel biomimetic structure for the epineurium that combines compression resistance, bending resistance, and ease of manipulation, thereby minimizing the risk of failure in clinical settings. Summary of the Invention

[0006] The purpose of this invention is to provide a method for preparing and applying a multi-level biomimetic neural conduit structure to solve the aforementioned problems in the background art.

[0007] To achieve the above objectives, the present invention provides the following technical solution:

[0008] One of the technical solutions of the present invention is to provide a nerve conduit with a multi-level biomimetic nerve membrane structure, comprising a biomimetic inner membrane scaffold with a secondary guiding structure and a biomimetic outer membrane scaffold with a double-layer structure;

[0009] The preparation method of the biomimetic endometrial scaffold with a secondary guiding structure includes the following steps:

[0010] (1) Spinning is performed using spinning solution to obtain spun fibers, and the spun fibers are collected on the receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°; then a second winding process is performed on the surface of the first mesh fiber membrane to prepare a second oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°, thus obtaining a receiver with oriented microfiber membrane material;

[0011] (2) The sacrificial layer material is loaded onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material to obtain an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold.

[0012] (3) Spinning is carried out using spinning solution to obtain spun fibers. Then, the spun fibers are wound three times on the surface of the oriented microfiber membrane loaded with oriented sacrificial fiber scaffold to prepare a third oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°. Then, a fourth winding process is carried out to prepare a fourth mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°, thus obtaining a multilayer fiber membrane that wraps the oriented sacrificial fiber scaffold.

[0013] (4) Cut the multilayer fiber membrane that encapsulates the directional sacrificial fiber scaffold along the receiver axis, and then remove the directional sacrificial fiber scaffold, leaving directional microchannels formed by the sacrificial fibers in the fiber membrane, to obtain the biomimetic endometrial scaffold with a secondary guiding structure.

[0014] Preferably, the thickness ratio of the first reticular fiber membrane to the second oriented fiber membrane is 0.5-1, and the thickness of the first reticular fiber membrane is 50-100 μm; the thickness ratio of the third oriented fiber membrane to the fourth reticular fiber membrane is 1-2, and the thickness of the fourth reticular fiber membrane is 50-100 μm.

[0015] Preferably, the fiber-forming substance in the spinning solution is one or more of polycaprolactone, polylactide, poly(lactide-glycolic acid) copolymer, polyglycolic acid, polyhydroxy fatty acid ester, poly(lactide-caprolactone) copolymer, polydioxanone, silk fibroin, chitosan, gelatin and collagen.

[0016] Preferably, in step (2): the loading method is 3D printing, specifically including the following steps: loading the sacrificial layer material into the 3D printer, performing 3D printing, loading the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, and during 3D printing, the sacrificial fibers are spirally loaded on the surface of the receiver by moving the receiver along the axial direction and rotating the receiver to obtain an oriented microfiber membrane loaded with an oriented sacrificial fiber support.

[0017] Preferably, the sacrificial layer material is a water-soluble polymer or a temperature-sensitive polymer; the printing air pressure of the 3D printing is 0.1-50 kPa, the barrel heating temperature is 30-200℃, the receiver rotation speed is 1-400 rpm, the receiver axial movement speed is 1-100 mm / sec, and the printing time is 1-100 min.

[0018] Preferably, the diameter of the sacrificial fiber is 100-1000 μm; the proportion of the oriented sacrificial fiber scaffold in the multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold is 50-70%.

[0019] Preferably, the method for preparing the biomimetic outer membrane scaffold with the double-layer structure includes the following steps:

[0020] A first spinning fiber is obtained by spinning using a first spinning solution, and the first spinning fiber is collected on a receiver for a first winding process to obtain an inner fiber layer with a fiber diameter of 0.1-10 μm and an inter-fiber angle of 5-15°. A second spinning fiber is obtained by spinning using a second spinning solution, and then the second spinning fiber is used to perform a second winding process on the surface of the inner fiber layer to wind an outer fiber layer with a fiber diameter of 10-100 μm and an inter-fiber angle of 15-45°, thus obtaining the biomimetic outer membrane scaffold with a double-layer structure.

[0021] The second technical solution of the present invention provides a method for preparing a nerve conduit with the above-mentioned multi-level biomimetic nerve membrane structure, comprising the following steps:

[0022] After the bionic endometrial scaffold with a secondary guiding structure is rolled up along the fiber direction in the directional fiber membrane, it is inserted into the middle cavity of the bionic outer membrane scaffold with a double-layer structure to obtain the nerve conduit of the multi-level bionic nerve membrane structure.

[0023] The third technical solution of the present invention provides an application of the above-mentioned multi-level biomimetic nerve membrane structure in the field of nerve conduits.

[0024] The technical principle of this invention is as follows:

[0025] To address the shortcomings of existing nerve conduit products, such as short repairable distance, poor efficacy, and weak resistance to compression and bending, this invention utilizes a wet spinning-3D printing combined system to fabricate a nerve conduit that integrates both a biomimetic inner membrane system and a biomimetic outer membrane system. On one hand, during the fabrication of the inner membrane biomimetic structure, the extruded fibers can be independently arranged on a roller collector, forming a fiber membrane with adjustable arrangement angle, selectable range, and adjustable thickness. On the other hand, during the fabrication of the outer membrane biomimetic structure, the extruded fibers can be continuously collected onto a rod-shaped collecting rod with adjustable diameter, forming a fiber braided tube with adjustable arrangement angle and selectable wall thickness. The fiber membrane, after being rolled up and inserted into the fiber braided tube, forms a loosely porous, multi-level biomimetic nerve conduit with microchannels. Specifically, the biomimetic inner membrane system is woven from fibers with a diameter of 1-50 μm; the layered fiber membranes can be arranged at different angles, with a total of four layers collected. The microstructure of the product is as follows: Figure 4 As shown, the cross-section of the fiber membrane exhibits a "multi-layer sandwich" structure. The two outermost layers have interwoven fibers forming a mesh structure (fiber angles of 15-45°), creating an interlocking structure that distributes load under tension, pressure, or torsion, preventing structural damage and deformation. The two innermost layers have fibers arranged at an angle of 0-5° along the long axis of the duct, exhibiting an oriented pattern. 3D-printed sacrificial material is filled between these two oriented fiber layers. After the sacrificial material is removed, microchannels exist between these two oriented fiber layers.

[0026] Regarding the guidance of nerve growth, the nerve conduit of this invention possesses a two-level guiding function. The first-level guiding function refers to the directional migration of cells along the oriented fibers in the nerve conduit, elongating their shape on the fibers and accelerating differentiation. The second-level guiding function refers to the fact that after a large number of cells migrate to the scaffold, accompanied by cell proliferation and increased number, multiple regenerated nerve fibers can aggregate in the directional microchannels to form thicker nerve bundles. The microchannels, on the one hand, constrain nerve fiber growth and promote aggregation, and on the other hand, guide the alignment and growth of nerve bundles. Specifically, during the fiber collection process, sacrificial material with a diameter of 100-1000 μm is extruded using 3D printing and filled into a multilayer fiber membrane. After spinning, the multilayer fiber membrane embeds these sacrificial materials; subsequently, by eluting the sacrificial material, a directional microchannel structure (arrangement angle along the long axis of the conduit is 0-5°) is formed in the multilayer fibers, and the channel diameter is maintained between 50-500 μm. The biomimetic endometrial system with directional microchannels formed within the fibrous membrane is composed of a multi-level guiding structure consisting of 50-500 μm directional microchannels (arranged at an angle of 0-5° along the long axis of the duct) and 1-50 μm directional microfibers (arranged at an angle of 0-5° along the long axis of the duct). The directional microfibers, acting as the primary guiding structure, enable high-precision orientation guidance of nerve cells, accelerating nerve growth and extending the repair distance. The directional microchannels, acting as the secondary guiding structure, provide space for aligned nerve growth, increase the contact guidance area, improve the stability of the microfiber arrangement in three-dimensional space, and prevent microfiber accumulation. Simultaneously, this structure also provides axial and radial mechanical support.

[0027] The biomimetic outer membrane system is essentially a layered fiber sheath with two layers tightly bonded together. A tungsten carbide rod is used to collect the fibers. After collecting a certain thickness of fibers, the tungsten carbide rod is removed to obtain a hollow conduit. The diameter of the tungsten carbide rod determines the inner diameter of the lumen, allowing for customization for nerves of different sizes. Specifically, the inner layer fibers are cross-wound and collected onto the tungsten carbide rod, with fiber diameters between 0.1-10 μm and fiber angles between 5-15°. The inner layer fiber diameter selected in this invention has a high surface area to volume ratio, enabling the woven fabric to have good air permeability. The smaller cross angle increases the mutual binding between fibers, improving the structure's tensile and bending resistance, and contributing to increased axial (length direction) strength of the tubular structure. Therefore, the inner layer of the biomimetic outer membrane structure uses finer fibers woven in a tighter arrangement, which improves the tear resistance and durability of the fiber sheath. The high-density weaving leaves smaller pores, ensuring good permeability while acting as a barrier. This allows for the transport of nutrients and metabolic waste while preventing excessive ingrowth of surrounding fibroblasts into the lumen, which could hinder internal nerve regeneration. Fibers are then collected to cover the inner layer; at this point, the fiber diameter is between 10-100 μm, and the angle between fibers is 15-45°. Larger diameter fibers provide better structural support, making the fabric stronger and improving the durability of the fiber tube. The larger cross angle allows the fibers to move relative to each other under stress, thus absorbing more energy rather than directly transmitting stress, enhancing radial (diameter direction) strength, and helping to resist external pressure and textile deformation. Therefore, the double-layer structure of the biomimetic outer membrane system simultaneously enhances the axial and radial mechanical strength of the tube, improving the overall resistance to bending, compression (radial pressure), kinking (axial pressure), and tension (axial tensile force) of the fiber tube.

[0028] The beneficial technical effects of the present invention are as follows:

[0029] 1. In the multi-level biomimetic nerve membrane structure of the present invention, the filler that plays a growth guiding role has a multi-level guiding structure. The first-level guidance involves the directional migration of cells along the fiber direction in the directional fiber membrane, elongating their shape on the fibers and accelerating differentiation. The second-level guidance involves directional microchannels constraining the growth of nerve fibers, promoting bundle maturation, and aligning the growth direction. Compared with existing hollow conduit products, the present invention can repair nerve defects of approximately 5 cm in length, including sensory nerves, motor nerves, and complex nerves, while effectively improving the electrical conduction function and behavioral recovery of regenerated nerves. Compared with existing simple directional microchannel structures, the present invention has a more precise cell guiding structure, which can significantly promote the directional spreading of cells and the directional elongation of axons. It has a higher contact guiding area and stronger spatial stability, which can promote cell infiltration and avoid the accumulation of fibers in three-dimensional space due to gravity. It also has the dual growth guiding advantages of directional microchannels and directional microfibers. The present invention solves the problem of insufficient contact guiding efficiency in existing solutions and overcomes the contradiction between the guiding fiber filling density, growth space, and guiding effect.

[0030] The conduit shell (biomimetic outer membrane scaffold) of this invention has a tightly integrated double-layer structure. The inner layer is a dense, cross-woven fibrous structure, which endows the conduit with barrier properties. This structure ensures the transport of nutrients and metabolic waste while preventing surrounding fibroblasts from ingrowing into the conduit and hindering nerve regeneration. It also improves the axial mechanical properties of the conduit. The outer layer is a loose, cross-woven fibrous structure, which gives the conduit resistance to compression and bending. This structure can prevent the conduit from failing due to compression when implanted in a muscular environment, and can also prevent failure due to bending when repairing cross-joint sites. It also improves the radial mechanical properties of the conduit.

[0031] 3. The manufacturing process of this catheter is highly controllable, and the density, diameter and angle of the microchannels and microfibers in the inner core, as well as the diameter, fiber angle and adhesion degree of the microfibers in the outer shell, can be precisely controlled; at the same time, the diameter and wall thickness of the catheter can also be controlled. Attached Figure Description

[0032] To more clearly illustrate the technical solutions in the embodiments of the present invention or the prior art, the drawings used in the embodiments will be briefly introduced below. Obviously, the drawings described below are only some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0033] Figure 1 This is a schematic diagram of the peripheral nerve membrane structure and representative scanning electron microscope (SEM) images of each membrane layer.

[0034] Figure 2 This is a schematic diagram of the single-fiber wet spinning platform of the present invention.

[0035] Figure 3 This is a schematic diagram of the fabrication process of the biomimetic endometrial stent of the present invention.

[0036] Figure 4 These are stereomicroscopic and SEM images of the nerve conduit of the multi-level biomimetic nerve membrane structure in Embodiment 3 of the present invention.

[0037] Figure 5 Schematic diagrams of different biomimetic inner membrane topologies used in cell experiments.

[0038] Figure 6 Two-dimensional characterization of the non-guided structure membrane scaffold, the microchannel structure membrane scaffold, the microfiber structure membrane scaffold, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0039] Figure 7 The study examines the regulation of neuronal cell morphology by non-guided structure membrane scaffolds, microchannel structure membrane scaffolds only, microfiber structure membrane scaffolds only, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0040] Figure 8 Schematic diagrams of stents for catheters with only densely cross-arranged microfiber structures, catheters with only loosely cross-arranged microfiber structures, and catheters with a densely cross-arranged microfiber structure inner layer plus a loosely cross-arranged microfiber structure outer layer.

[0041] Figure 9 Compression cycle curves (A) and compressive stress changes (B) of dense membrane structure catheters, loose membrane structure catheters, and double-layer membrane structure catheters of Example 3.

[0042] Figure 10 A schematic diagram (A) of using the Transwell system to evaluate the prevention of fibroblast infiltration by a biomimetic outer membrane scaffold (double-walled tube), statistical results of the porosity of the fibrous membranes that make up each wall layer (C), crystal violet staining results (B) and statistical results (D) after fibroblasts penetrate each membrane layer.

[0043] Figure 11 A schematic diagram (A) and statistical results (B) for evaluating the nutrient permeability of the biomimetic outer membrane scaffold of Example 3 using the Transwell system.

[0044] Figure 12 The image shows the bending resistance of the multi-level biomimetic neural membrane structure in Embodiment 3 of the present invention.

[0045] Figure 13 The diagram shows the structure of the membrane scaffold without a guide structure, the membrane scaffold with only a microchannel structure, the membrane scaffold with only a microfiber structure, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0046] Figure 14 Stereoscopic and SEM images (A) and internal porosity statistics (B) of the non-guided structure membrane scaffold, the microchannel structure membrane scaffold, the microfiber structure membrane scaffold, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0047] Figure 15 H&E and DAPI staining images of the non-guided structure membrane scaffold, the microchannel structure membrane scaffold, the microfiber structure membrane scaffold, and the microchannel + microfiber structure membrane scaffold of Example 3 after subcutaneous implantation in rats 7 days and 14 days later (Figure A), and the corresponding cell infiltration count statistics (Figure B).

[0048] Figure 16 This study evaluates the efficacy of the biomimetic outer membrane scaffold of Example 3, the microchannel + microfiber structure membrane scaffold of Example 3, and autologous nerve in repairing a 2cm sciatic nerve defect model in rats. In the figures, A shows immunofluorescence imaging of the regenerated nerve tissue, B shows transmission electron microscopy imaging of the distal myelin sheath structure of the regenerated nerve, and C, D, and E represent the electrophysiological function recovery assessments of the biomimetic outer membrane scaffold of Example 3, the microchannel + microfiber structure membrane scaffold of Example 3, and the autologous nerve, respectively.

[0049] Figure 17 This study evaluates the effectiveness of the microchannel + microfiber structure membrane scaffold and autologous nerve in a 5cm sciatic nerve defect repair model in a large animal model, as described in Example 3. A shows intraoperative observation, images taken 6 months post-operation, and intraoperative TEM imaging; B shows MRI evaluation; C shows electrophysiological function testing; and D shows gait analysis and motor function recovery.

[0050] Figure 18 A heatmap showing the comprehensive scoring of five key performance indicators for various nerve grafts that can be used to repair long-segment nerve defects. Detailed Implementation

[0051] Various exemplary embodiments of the present invention will now be described in detail. This detailed description should not be considered as a limitation of the present invention, but rather as a more detailed description of certain aspects, features, and embodiments of the present invention. It should be understood that the terminology used in this invention is merely for describing particular embodiments and is not intended to limit the present invention.

[0052] Furthermore, regarding the numerical ranges in this invention, it should be understood that each intermediate value between the upper and lower limits of the range is also specifically disclosed. Any stated value or intermediate value within a stated range, as well as each smaller range between any other stated value or intermediate value within said range, are also included in this invention. The upper and lower limits of these smaller ranges may be independently included or excluded from the range.

[0053] Unless otherwise stated, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art. While only preferred methods and materials have been described herein, any methods and materials similar or equivalent to those described herein may be used in the implementation or testing of this invention. It should be noted that any aspects of this invention not described in detail are conventional practices in the art and are not the focus of this invention.

[0054] The terms “comprising,” “including,” “having,” “containing,” etc., used in this invention are all open-ended terms, meaning that they include but are not limited to.

[0055] This invention discloses a neural conduit with a multi-level biomimetic neural membrane structure, which includes a biomimetic inner membrane scaffold with a secondary guiding structure and a biomimetic outer membrane scaffold with a tightly integrated double-layer structure.

[0056] Furthermore, the biomimetic endometrial scaffold with a secondary guiding structure is a multilayer fiber membrane encapsulating a sacrificial template, obtained through layer-by-layer assembly. Its preparation method includes the following steps:

[0057] (1) Constructing a layered fiber membrane using single-fiber wet spinning technology: Using chemically pure reagents (acetic acid, formic acid, water, tetrahydrofuran, dimethyl sulfoxide, dichloromethane, chloroform, acetic acid, acetone, trifluoroethanol, or hexafluoroisopropanol) as solvents, prepare a solution of a certain concentration (1-60 g / mL) of degradable polymers (including one or more of polycaprolactone (PCL), polylactide (PLA), poly(lactide-glycolic acid) copolymer (PLGA), polyglycolic acid (PGA), polyhydroxyalkanoate (PHA), poly(lactide-caprolactone) copolymer (PLCL), and poly(p-dioxane-hexanone (PDS)) or natural materials (one or more of silk fibroin, chitosan, gelatin, and collagen) to obtain a spinning solution; load the spinning solution into a syringe and further install it on a micro-injection pump; use anhydrous ethanol, anhydrous methanol, vegetable oil, or glycerol as a coagulation medium. In a coagulation bath, the syringe needle is placed inside the coagulation bath, and a cylindrical receiver is mounted on a triaxial receiver (capable of movement in the x-direction plus rotation of the axis). The injection pump flow rate is set to 0.1-20 mL / h, the receiver rotation speed to 1-800 rpm, the receiver x-direction (axial) movement speed to 1-100 mm / sec, and the spinning time to 1-600 min for wet spinning. The resulting spun fibers are collected on the receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°, with a collected thickness of 50-100 μm. Then, a second winding process is performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°, with a collected thickness of 100-150 μm, resulting in a receiver with oriented microfiber membrane material.

[0058] By adjusting parameters such as polymer concentration, needle diameter, receiver rotation speed and movement speed, and propulsion flow rate, mesh fiber membranes (first layer) and oriented fiber membranes (second layer) with controllable fiber diameters (1-50 μm) can be prepared. The definition of an oriented fiber membrane: the fibers composing this layer are arranged approximately parallel to each other with an angle between 0-5°. The definition of a mesh fiber membrane: the fibers composing this layer are interwoven in a mesh-like pattern with an angle between 15-45°.

[0059] (2) Constructing a directional sacrificial fiber scaffold using 3D printing technology: Add one or more of water-soluble polymer materials (polyvinyl alcohol (PVA), polyethylene oxide (PEO), polyethylene glycol (PEG), polyvinylpyrrolidone (PVP), and polyacrylic acid (PAA)) or thermosensitive polymer materials (polyoxyethylene polyoxypropylene ether (Pluronic F-127) and / or poly(N-isopropylacrylamide) (PNIPAM)) into the barrel of the 3D printer and further install it on the 3D printer; install the receiver with the directional microfiber membrane material on the triaxial receiver of the 3D printer (which can realize x-axis movement plus axis rotation). Set the printing air pressure to 0.1-50 kPa, the barrel heating temperature to 30-200℃, the receiver rotation speed to 1-400 rpm, the receiver x-direction moving speed to 1-100 mm / sec, and the printing time to 1-100 min for 3D printing. This will load the printed sacrificial fibers onto the surface of the oriented microfiber membrane on the receiver with the oriented microfiber membrane material. During 3D printing, the sacrificial fibers are spirally loaded on the receiver surface by moving the receiver along the axial direction (i.e., the x-direction) and rotating the receiver, thus obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold.

[0060] Directional sacrificial fiber scaffolds with controllable diameters (100-1000 μm) can be fabricated by adjusting parameters such as needle diameter, needle shape, receiver rotation speed and movement speed, and printing air pressure. A directional sacrificial fiber scaffold is defined as follows: the linear scaffold (sacrificial fibers) extruded using 3D printing technology are arranged almost parallel to the fiber direction of the directional microfiber membrane, with an angle between them between 0-5°. The sacrificial fibers can be easily removed subsequently by washing with water or temperature control, leaving space.

[0061] The diameter of the sacrificial fiber is determined by the nozzle orifice of the 3D printer. Common nozzle diameters are between 0.2-1 mm, and the extruded fiber diameter is smaller than the nozzle diameter. Adding an electric field between the nozzle and the substrate for electrostatic stretching (electrostatic fusion printing) further stretches the extruded fiber, achieving higher resolution fiber deposition (fineer fibers). However, considering that the sacrificial template is modeled after the perimembranes of a nerve, with a diameter between 50-500 μm, this means that after removing the sacrificial template, microchannels with pore sizes of 50-500 μm need to be formed. Since the sacrificial template shrinks after cooling, the surrounding fiber layer may not adhere tightly to the outside of the template. Therefore, the sacrificial template support is set slightly larger than the preset orifice diameter, typically between 100-1000 μm.

[0062] (3) The spinning fibers are wound three times on the surface of the oriented microfiber membrane loaded with oriented sacrificial fiber scaffold to wind a third oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°, and the thickness is collected to be 100-150 μm. Then, a fourth winding process is performed to wind a fourth mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°, and the thickness is collected to be 50-100 μm, thereby wrapping the oriented sacrificial fiber scaffold in it, and obtaining a multilayer fiber membrane wrapped with the oriented sacrificial fiber scaffold.

[0063] (4) The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis, and then washed with water to remove the water-soluble sacrificial fiber scaffold or cooled to below the phase transition temperature to remove the temperature-sensitive sacrificial fiber scaffold, thereby leaving directional microchannels with a diameter of 50-1000μm formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers of 1-50μm further increase the porosity of the fiber membrane; thus forming a biomimetic endometrial scaffold with a secondary guiding structure that combines directional guiding fibers (primary guidance) and directional microchannels (secondary guidance).

[0064] Furthermore, the method for preparing the biomimetic outer membrane scaffold with a tightly bound bilayer structure includes the following steps:

[0065] A fiber inner layer with a fine fiber diameter and small cross angle is constructed using single-fiber wet spinning technology: A spinning solution of a certain concentration (1-60 g / mL) is prepared using chemically pure reagents (acetic acid, formic acid, water, tetrahydrofuran, dimethyl sulfoxide, dichloromethane, chloroform, acetic acid, acetone, trifluoroethanol, or hexafluoroisopropanol) as solvents. This solution is composed of one or more of the following: polycaprolactone (PCL), polylactide (PLA), poly(lactide-glycolic acid) copolymer (PLGA), polyglycolic acid (PGA), polyhydroxyalkanoate (PHA), poly(lactide-caprolactone) copolymer (PLCL), and poly(p-dioxane-hexanone) (PDS)) or natural materials (one or more of the following: silk fibroin, chitosan, gelatin, and collagen). The spinning solution is loaded into a syringe and further mounted on a micro-injection pump. The solution is then spun with anhydrous... Ethanol, anhydrous methanol, vegetable oil, or glycerol are used as coagulation baths. The syringe needle is placed inside the coagulation bath, and a rod-shaped receiver is mounted on a triaxial receiver (capable of movement in the x-direction plus rotation of the axis). The injection pump flow rate is set to 0.1-2 mL / h, the receiver rotation speed to 1-2000 rpm, the receiver x-direction movement speed to 1-1000 mm / sec, and the spinning time to 1-300 min for wet spinning. The resulting spun fibers are collected on the receiver and wound once to obtain an inner fiber layer with a fiber diameter of 0.1-10 μm and an inter-fiber angle of 5-15°. Then, a second winding process is performed on the surface of the inner fiber layer to wind an outer fiber layer with a fiber diameter of 10-100 μm and an inter-fiber angle of 15-45°. After peeling, the biomimetic outer membrane scaffold with a tightly bonded double-layer structure is obtained.

[0066] By adjusting parameters such as polymer concentration, needle diameter, receiver rotation speed and moving speed, and propulsion flow rate, a tubular inner fiber layer with a fiber diameter of 0.1-10 μm and an inter-fiber angle of 5-15° can be prepared, as well as a tubular outer fiber layer with a fiber diameter of 10-100 μm and an inter-fiber angle of 15-45° can be prepared.

[0067] This invention also discloses a method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0068] After the biomimetic inner membrane scaffold is rolled up along the fiber direction of the oriented fiber membrane, it is inserted into the central cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure. The porosity of the remaining volume of the cavity in the conduit is set to ≤30% to ensure a close fit while retaining a certain amount of elastic space.

[0069] The single-fiber wet spinning technology involved in the preparation process, such as Figure 2As shown, the spinning solution squeezed from the needle enters the coagulation bath and undergoes phase separation. The separated solvent diffuses into the coagulation bath and sinks to the bottom due to its higher density. The separated polymer is pulled out of the liquid surface and collected on the receiver. One fiber is produced from one outlet. By adjusting the transverse and rotational speeds of the receiver, fibers with different arrangement angles can be obtained. The single fiber is continuously woven to form a complete fiber membrane.

[0070] Unless otherwise specified, "room temperature" in this invention refers to 10-30°C.

[0071] All raw materials used in the following embodiments and comparative examples of the present invention are commercially available products.

[0072] Example 1

[0073] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0074] Preparation of the biomimetic outer membrane scaffold: First, 1.0 g of PCL with a number-average molecular weight of 40,000 was weighed and added to 10 mL of chloroform. The solution was stirred and dissolved overnight at room temperature to obtain a 10% PCL solution. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiving rod with a diameter of 2 mm was mounted on a triaxial receiver. The PCL solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 50 mL of vegetable oil. By adjusting the y-axis of the receiver, the outlet of the coagulation bath was positioned 10 cm away from the receiving rod. The injection pump speed was set to 0.1 mL / h, the receiving rod rotation speed to 500 rpm, the x-axis movement speed to 0.5 mm / sec, the spinning time to 10 min, and the inner layer thickness to 100 μm, resulting in an inner fiber layer with a fiber diameter of 7 ± 2.5 μm and an inter-fiber angle of 10 ± 5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 1 cm. The injection pump speed was set to 5 mL / h, the receiving rod rotation speed to 50 rpm, the x-axis movement speed to 2 mm / sec, the spinning time to 20 min, and the outer layer thickness to 100 μm, resulting in a fiber outer layer with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 35 ± 5°. The fiber was then peeled off, and the vegetable oil on the membrane material was removed using anhydrous ethanol. The tubular biomimetic outer membrane scaffold was then vacuum-dried for later use.

[0075] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a 20mm diameter cylindrical receiver rod was mounted on a triaxial receiver. A 10% PCL solution was drawn into a syringe, which was then attached to a microinjection pump. The syringe needle was placed in a coagulation bath containing 50mL of vegetable oil, with the coagulation bath outlet positioned 3cm from the receiver rod. The injection pump's feed rate was set to 0.5mL / h, the receiver rod's rotation speed to 50rpm, and the x-axis movement speed to 1m. Wet spinning was performed at a speed of m / sec and a spinning time of 30 min. The resulting spun fibers were collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 15±2.5 μm and an inter-fiber angle of 35±5°, with a collected thickness of 75±25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 15±2.5 μm and an inter-fiber angle within 5°, with a collected thickness of 125±25 μm, resulting in a receiver with oriented microfiber membrane material. Subsequently, 2.0g of PEG with a number-average molecular weight of 8000 was weighed and added to the printer barrel. The printer nozzle was placed 1cm away from the receiver rod. The printing air pressure was set to 1kPa, the barrel heating temperature to 60℃, the receiver rod rotation speed to 10rpm, the x-axis moving speed to 0.1mm / sec, and the printing time to 5min. 3D printing was performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with oriented microfiber membrane material, thereby obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter was 200μm and the fiber angle with the oriented microfiber membrane was less than 5°. Further, the above-mentioned PCL solution was used for single-fiber wet spinning. The injection pump speed was set to 0.5 mL / h, the receiving rod speed to 100 rpm, the x-axis moving speed to 1 mm / sec, and the spinning time to 60 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 15 ± 2.5 μm and an inter-fiber angle of less than 5°, with a collected thickness of 125 ± 25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 15 ± 2.5 μm and an inter-fiber angle of 35 ± 5°, with a collected thickness of 75 ± 25 μm. This process encapsulated the oriented sacrificial fiber scaffold, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45 ± 15%). The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis. Then, ultrapure water is used to remove the PEG sacrificial material, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increase the porosity of the fiber membrane.The plant oil on the membrane material was removed using anhydrous ethanol, and the biomimetic inner membrane scaffold was then vacuum dried for later use.

[0076] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0077] Example 2

[0078] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0079] Preparation of the biomimetic outer membrane scaffold: First, 2.0 g of PLGA with a number-average molecular weight of 20,000 was weighed and added to 10 mL of dichloromethane. The solution was stirred and dissolved overnight at room temperature to obtain a 20% PLGA solution. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiver rod with a diameter of 3 mm was mounted on a triaxial receiver. The PLGA solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 100 mL of anhydrous ethanol. By adjusting the y-axis of the receiver, the coagulation bath outlet was positioned 15 cm away from the receiver rod. The injection pump speed was set to 0.2 mL / h, the receiver rod rotation speed to 300 rpm, the x-axis movement speed to 0.8 mm / sec, the spinning time to 15 min, and the inner layer thickness to 150 μm. This resulted in an inner fiber layer with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of 10 ± 5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 5 cm. The injection pump speed was set to 4 mL / h, the receiving rod rotation speed to 60 rpm, the x-axis movement speed to 4 mm / sec, the spinning time to 25 min, and the outer layer thickness to 250 μm, resulting in a fiber outer layer with a fiber diameter of 0.25 ± 5 μm and an inter-fiber angle of 35 ± 5°. The fiber was then peeled off to obtain the biomimetic outer membrane scaffold, which was then vacuum-dried for later use.

[0080] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a cylindrical receiver rod with a diameter of 25 mm was mounted on a triaxial receiver. A 20% PLGA solution was drawn into a syringe, which was then attached to a microinjection pump. The syringe needle was placed in a coagulation bath containing 100 mL of anhydrous ethanol, with the coagulation bath outlet positioned 10 cm from the receiver rod. The injection pump's feed rate was set to 1 mL / h, the receiver rod's rotation speed to 200 rpm, and the x-axis movement speed to 2 m / s. Wet spinning was performed at a speed of m / sec and a spinning time of 50 min. The resulting spun fibers were collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 7.5±2.5 μm and an inter-fiber angle of 35±5°, with a collected thickness of 75±25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 7.5±2.5 μm and an inter-fiber angle within 5°, with a collected thickness of 125±25 μm, resulting in a receiver with oriented microfiber membrane material. Subsequently, 3.0g of PEO with a number-average molecular weight of 100,000 was weighed and added to the printer barrel. The printer nozzle was placed 1cm away from the receiver rod. The printing air pressure was set to 2kPa, the barrel heating temperature to 65℃, the receiver rod rotation speed to 15rpm, the x-axis moving speed to 0.3mm / sec, and the printing time to 10min. 3D printing was performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with oriented microfiber membrane material, thereby obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter is 300μm and the fiber angle with the oriented microfiber membrane is less than 5°. Further, the above-mentioned PLGA solution was used for single-fiber wet spinning. The injection pump feed rate was set to 1 mL / h, the receiving rod rotation speed to 200 rpm, the x-axis movement speed to 2 mm / sec, and the spinning time to 130 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of less than 5°, with a collected thickness of 125 ± 25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of 35 ± 5°, with a collected thickness of 75 ± 25 μm. This process encapsulated the oriented sacrificial fiber scaffold within the membrane, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45 ± 15%). The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis to form a membrane. Then, ultrapure water is used to remove the PEO sacrificial material, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increase the porosity of the fiber membrane, resulting in a biomimetic inner membrane scaffold, which is then vacuum dried for later use.

[0081] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0082] Example 3

[0083] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0084] Preparation of the biomimetic outer membrane scaffold: First, 1g of PLCL with a number-average molecular weight of 80,000 was weighed and added to 10mL of hexafluoroisopropanol. The solution was stirred and dissolved overnight at room temperature to obtain a 10% PLCL solution. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiver rod with a diameter of 4mm was mounted on a triaxial receiver. The PLCL solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 200mL of anhydrous ethanol. By adjusting the y-axis of the receiver, the coagulation bath outlet was positioned 25cm from the receiver rod. The injection pump speed was set to 1mL / h, the receiver rod rotation speed to 500rpm, the x-axis movement speed to 1.5mm / sec, the spinning time to 10min, and the inner layer thickness to 100μm, resulting in an inner fiber layer with a fiber diameter of 7.5±2.5μm and an inter-fiber angle of 10±5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 2 cm. The injection pump speed was set to 8 mL / h, the receiving rod rotation speed to 100 rpm, the x-axis movement speed to 8 mm / sec, the spinning time to 20 min, and the outer layer thickness to 200 μm. This resulted in a fiber outer layer with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 35 ± 5°. The fiber was then peeled off to obtain the biomimetic outer membrane scaffold, which was vacuum dried for later use.

[0085] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a cylindrical receiver rod with a diameter of 100 mm was mounted on a triaxial receiver. A 10% PLCL solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 200 mL of anhydrous ethanol, with the coagulation bath outlet positioned 30 cm from the receiver rod. The injection pump's feed rate was set to 1.5 mL / h, the receiver rod's rotation speed to 600 rpm, and the x-axis movement speed... Wet spinning was performed at a speed of 3 mm / sec and a spinning time of 60 min. The resulting spun fibers were then collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 15±5 μm and an inter-fiber angle of 35±10°, with a collected thickness of 75±25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 15±5 μm and an inter-fiber angle within 5°, with a collected thickness of 125±25 μm, thus obtaining a receiver with oriented microfiber membrane material. Subsequently, 2.0g of PVA with a number-average molecular weight of 50,000 was weighed and added to the printer barrel. The printer nozzle was positioned 1cm away from the receiver rod. The printing air pressure was set to 1.5kPa, the barrel heating temperature to 180℃, the receiver rod rotation speed to 5rpm, the x-axis moving speed to 0.1mm / sec, and the printing time to 10min. 3D printing was then performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, thereby obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter was 300μm and the fiber angle with the oriented microfiber membrane was less than 5°. Further, the above-mentioned PLCL solution was used for single-fiber wet spinning. The injection pump speed was set to 1.5 mL / h, the receiving rod speed to 600 rpm, the x-axis moving speed to 3 mm / sec, and the spinning time to 180 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 15±5 μm and an inter-fiber angle of less than 5°, resulting in a thickness of 125±25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 15±5 μm and an inter-fiber angle of 35±10°, resulting in a thickness of 75±25 μm. This process encapsulated the oriented sacrificial fiber scaffold within the membrane, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45±15%). The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis to form a membrane. Then, ultrapure water is used to remove the PVA sacrificial material, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increase the porosity of the fiber membrane, resulting in a biomimetic inner membrane scaffold, which is then vacuum dried for later use.

[0086] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0087] Example 4

[0088] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0089] Preparation of the biomimetic outer membrane scaffold: First, 2.5 g of PDS with a number-average molecular weight of 20,000 was weighed and added to 10 mL of tetrahydrofuran. The solution was stirred and dissolved overnight at room temperature to obtain a PDS solution with a concentration of 25 g / mL. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiving rod with a diameter of 5 mm was mounted on a triaxial receiver. The PDS solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 500 mL of anhydrous methanol. By adjusting the y-axis of the receiver, the outlet of the coagulation bath was positioned 10 cm away from the receiving rod. The injection pump speed was set to 4 mL / h, the receiving rod rotation speed to 400 rpm, the x-axis movement speed to 2 mm / sec, the spinning time to 20 min, and the inner layer thickness to 200 μm. This resulted in an inner fiber layer with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 10 ± 5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 1 cm. The injection pump speed was set to 12 mL / h, the receiving rod rotation speed to 80 rpm, the x-axis movement speed to 6 mm / sec, the spinning time to 40 min, and the outer layer thickness to 350 μm. This resulted in a fiber outer layer with a diameter of 35 ± 10 μm and an inter-fiber angle of 30 ± 15°. The fiber was then peeled off to obtain the biomimetic outer membrane scaffold, which was then vacuum-dried for later use.

[0090] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a 60mm diameter cylindrical receiver rod was mounted on a triaxial receiver. 25g / mL of PDS solution was drawn into a syringe, which was then attached to a microinjection pump. The syringe needle was placed in a coagulation bath containing 500mL of anhydrous ethanol, with the coagulation bath outlet positioned 15cm from the receiver rod. The injection pump's feed rate was set to 2mL / h, the receiver rod's rotation speed to 300rpm, and the x-axis movement speed... Wet spinning was performed at a speed of 1 mm / sec and a spinning time of 100 min. The resulting spun fibers were then collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 30 ± 15°, with a collected thickness of 75 ± 25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 25 ± 5 μm and an inter-fiber angle within 5°, with a collected thickness of 125 ± 25 μm, thus obtaining a receiver with oriented microfiber membrane material. Subsequently, 2.0g of Pluronic F-127 with a number-average molecular weight of 12600 was weighed and added to the printer barrel. The printer nozzle was positioned 0.5cm away from the receiver rod. The printing air pressure was set to 0.5kPa, the barrel heating temperature to 25℃, the receiver rod rotation speed to 10rpm, the x-axis movement speed to 0.5mm / sec, and the printing time to 15min. 3D printing was performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, thereby obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter was 400μm and the fiber angle with the oriented microfiber membrane was less than 5°. Further, the above-mentioned PDS solution was used for single-fiber wet spinning. The injection pump speed was set to 2 mL / h, the receiving rod speed to 300 rpm, the x-axis moving speed to 1 mm / sec, and the spinning time to 100 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 25±5 μm and an inter-fiber angle of less than 5°, resulting in a thickness of 125±25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 25±5 μm and an inter-fiber angle of 30±15°, resulting in a thickness of 75±25 μm. This process encapsulated the oriented sacrificial fiber scaffold within the membrane, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45±15%).The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold was cut open and laid flat along the receiver axis to form a membrane. Then, the Pluronic F-127 sacrificial material was removed using an ice-water mixture of ultrapure water, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increased the porosity of the fiber membrane, resulting in a biomimetic inner membrane scaffold, which was then vacuum dried for later use.

[0091] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0092] Example 5

[0093] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0094] Preparation of the biomimetic outer membrane scaffold: First, 1.0 g of PHA with a number-average molecular weight of 150,000 was weighed and added to 10 mL of dimethyl sulfoxide. The solution was stirred and dissolved overnight at room temperature to obtain a PHA solution with a concentration of 10 g / mL. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiving rod with a diameter of 1.5 mm was mounted on a triaxial receiver. The above PDS solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 50 mL of anhydrous methanol. By adjusting the y-axis of the receiver, the outlet of the coagulation bath was positioned 20 cm away from the receiving rod. The injection pump speed was set to 2 mL / h, the receiving rod rotation speed to 800 rpm, the x-axis movement speed to 3 mm / sec, the spinning time to 15 min, and the inner layer thickness to 100 μm, resulting in a fiber inner layer with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of 10 ± 5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 4 cm. The injection pump speed was set to 10 mL / h, the receiving rod rotation speed to 100 rpm, the x-axis movement speed to 10 mm / sec, the spinning time to 25 min, and the outer layer thickness to 200 μm. This resulted in a fiber outer layer with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 30 ± 15°. The fiber was then peeled off to obtain the biomimetic outer membrane scaffold, which was then vacuum-dried for later use.

[0095] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a cylindrical receiver rod with a diameter of 15 mm was mounted on a triaxial receiver. A 10 g / mL PHA solution was drawn into a syringe, which was then attached to a microinjection pump. The syringe needle was placed in a coagulation bath containing 50 mL of anhydrous ethanol, with the coagulation bath outlet positioned 10 cm from the receiver rod. The injection pump's feed rate was set to 1 mL / h, the receiver rod's rotation speed to 500 rpm, and the x-axis movement speed to 1 m / s. Wet spinning was performed at a speed of m / sec and a spinning time of 30 min. The resulting spun fibers were collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 7.5±2.5 μm and an inter-fiber angle of 30±15°, with a collected thickness of 75±25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 7.5±2.5 μm and an inter-fiber angle within 5°, with a collected thickness of 125±25 μm, resulting in a receiver with oriented microfiber membrane material. Subsequently, 3.0g of PNIPAM with a number-average molecular weight of 300,000 was weighed and added to the printer barrel. The printer nozzle was positioned 0.8cm away from the receiver rod. The printing air pressure was set to 1.5kPa, the barrel heating temperature to 35℃, the receiver rod rotation speed to 15rpm, the x-axis moving speed to 0.5mm / sec, and the printing time to 5min. 3D printing was then performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, thereby obtaining an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter was 300μm and the fiber angle between the sacrificial fiber and the oriented microfiber membrane was less than 5°. Further, the above-mentioned PHA solution was used to continue single-fiber wet spinning. The injection pump speed was set to 1 mL / h, the receiving rod speed to 500 rpm, the x-axis moving speed to 1 mm / sec, and the spinning time to 60 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of less than 5°, with a collected thickness of 125 ± 25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 7.5 ± 2.5 μm and an inter-fiber angle of 30 ± 15°, with a collected thickness of 75 ± 25 μm. This process encapsulated the oriented sacrificial fiber scaffold, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45 ± 15%).The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis to form a membrane. Then, the PNIPAM sacrificial material is removed using an ice-water mixture of ultrapure water, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increase the porosity of the fiber membrane, resulting in a biomimetic inner membrane scaffold, which is then vacuum dried for later use.

[0096] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0097] Example 6

[0098] A method for preparing a neural conduit with a multi-level biomimetic neural membrane structure, comprising the following steps:

[0099] Preparation of the biomimetic outer membrane scaffold: First, 1.0 g of Col was weighed and added to 100 mL of 0.5% acetic acid solution. The solution was stirred and dissolved overnight at room temperature to obtain a Col solution with a concentration of 1 g / mL. The biomimetic outer membrane scaffold was then prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiving rod with a diameter of 3.0 mm was mounted on a triaxial receiver. The Col solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 200 mL of anhydrous ethanol. By adjusting the y-axis of the receiver, the outlet of the coagulation bath was positioned 5 cm away from the receiving rod. The injection pump speed was set to 3 mL / h, the receiving rod rotation speed to 100 rpm, the x-axis movement speed to 2 mm / sec, the spinning time to 30 min, and the inner layer thickness to 200 μm, resulting in a fiber inner layer with a fiber diameter of 45 ± 5 μm and an inter-fiber angle of 10 ± 5°. The distance between the coagulation bath outlet and the receiving rod was adjusted to 1 cm. The injection pump speed was set to 9 mL / h, the receiving rod rotation speed to 50 rpm, the x-axis movement speed to 8 mm / sec, the spinning time to 60 min, and the outer layer thickness to 400 μm. This resulted in a fiber outer layer with a diameter of 45 ± 5 μm and an inter-fiber angle of 35 ± 10°. The fiber was then peeled off to obtain the biomimetic outer membrane scaffold, which was vacuum dried for later use.

[0100] Fabrication of a biomimetic endometrial scaffold: A biomimetic endometrial scaffold was fabricated in a fume hood using a combined single-fiber wet spinning and 3D printing technique. First, a cylindrical receiver rod with a diameter of 100 mm was mounted on a triaxial receiver. A 1 g / mL Col solution was drawn into a syringe, which was then attached to a microinjection pump. The syringe needle was placed in a coagulation bath containing 200 mL of anhydrous ethanol, with the coagulation bath outlet positioned 10 cm from the receiver rod. The injection pump's feed rate was set to 1.5 mL / h, the receiver rod's rotation speed to 200 rpm, and the x-axis movement speed... Wet spinning was performed at a speed of 1.5 mm / sec and a spinning time of 100 min. The resulting spun fibers were then collected onto a receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 45±5 μm and an inter-fiber angle of 35±10°, with a collected thickness of 75±25 μm. Then, a second winding process was performed on the surface of the first mesh fiber membrane to wind a second oriented fiber membrane with a fiber diameter of 45±5 μm and an inter-fiber angle within 5°, with a collected thickness of 125±25 μm, thus obtaining a receiver with oriented microfiber membrane material. Subsequently, 3.0g of PVP with a number-average molecular weight of 100,000 was weighed and added to the printer barrel. The printer nozzle was positioned 1cm away from the receiver rod. The printing air pressure was set to 2.5kPa, the barrel heating temperature to 130℃, the receiver rod rotation speed to 10rpm, the x-axis movement speed to 0.2mm / sec, and the printing time to 15min. 3D printing was then performed to load the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, resulting in an oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold. The sacrificial fiber diameter was 400μm and the fiber angle with the oriented microfiber membrane was less than 5°. Further, the above-mentioned Col solution was used for single-fiber wet spinning. The injection pump speed was set to 1.5 mL / h, the receiving rod speed to 200 rpm, the x-axis moving speed to 1.5 mm / sec, and the spinning time to 200 min. The resulting spun fibers were then wound three times on the surface of a oriented microfiber membrane loaded with an oriented sacrificial fiber scaffold to form a third oriented fiber membrane with a fiber diameter of 45 ± 5 μm and an inter-fiber angle of less than 5°, resulting in a thickness of 125 ± 25 μm. Then, a fourth winding process was performed to form a fourth mesh fiber membrane with a fiber diameter of 45 ± 5 μm and an inter-fiber angle of 35 ± 10°, resulting in a thickness of 75 ± 25 μm. This process encapsulated the oriented sacrificial fiber scaffold within the membrane, resulting in a multilayer fiber membrane encapsulating the oriented sacrificial fiber scaffold (with the oriented sacrificial fiber scaffold accounting for 45 ± 15%). The multilayer fiber membrane encapsulating the directional sacrificial fiber scaffold is cut open and laid flat along the receiver axis to form a membrane. Then, ultrapure water is used to remove the PVP sacrificial material, thereby leaving directional microchannels formed by the sacrificial fibers in the fiber membrane. The gaps between the fibers further increase the porosity of the fiber membrane, resulting in a biomimetic inner membrane scaffold, which is then vacuum dried for later use.

[0101] After the biomimetic endometrial scaffold is rolled up along the fiber direction in the directional fiber membrane, it is loaded (inserted) into the middle cavity of the biomimetic outer membrane scaffold, thus obtaining a nerve conduit with a multi-level biomimetic nerve membrane structure.

[0102] Effect verification

[0103] Figure 1 This is a schematic diagram of the peripheral nerve membrane structure and representative scanning electron microscope (SEM) images of each membrane layer.

[0104] Figure 1 The representative scanning electron microscope images of each membrane layer are specifically images of the typical microstructural features of the rat sciatic nerve after decellularization, which are used to compare the morphological biomimetic basis of each membrane layer design in the biomimetic conduit.

[0105] Depend on Figure 1 It is evident that the main membranous system of natural nerves consists of the endoneurium (En), perineurium (P), and epipineurium (Ep). The endoneurium is a thin, loose connective tissue membrane surrounding the axon and myelin sheath. It is composed of parallel-arranged fibers, further forming axially extending microchannel structures, and plays a crucial role in guiding nerve growth and resisting axial tensile forces. The perineurium is a densely woven connective tissue membrane surrounding each nerve fiber bundle, acting as a barrier to substances entering and exiting the surrounding nerve fibers. The epipineurium is a thick, loose connective tissue membrane surrounding multiple nerve bundles, composed of flat, band-like fibers crossing at small angles, providing support and protection for the nerve fibers.

[0106] Figure 2 This is a schematic diagram of the single-fiber wet spinning platform of the present invention.

[0107] Depend on Figure 2 As can be seen, the spinning solution is loaded into a syringe, extruded by a micro-injection pump, and then immersed in a coagulation bath. In the coagulation bath, the fibers separate into single fibers, which are then received by a triaxial moving receiver. By changing a series of parameters, precise control can be achieved over the fiber diameter, arrangement angle, and interfiber adhesion.

[0108] Figure 3 This is a schematic diagram of the fabrication process of the biomimetic endometrial stent of the present invention.

[0109] Depend on Figure 3As can be seen, the fabrication process of the biomimetic endometrial scaffold of the present invention mainly consists of four steps: the first step is to construct oriented microfibers using single-fiber wet spinning technology; the second step is to construct oriented sacrificial fiber scaffolds using 3D printing technology; the third step is to repeat the operation of the first step, so that the oriented microfibers continue to be stacked in multiple layers until they completely cover the sacrificial fiber scaffold constructed in the second step; the fourth step is to wash the membrane scaffold with the sandwich structure of oriented microfiber-oriented sacrificial fiber-oriented microfiber to remove the water-soluble sacrificial fiber scaffold or cool it to below the phase transition temperature to remove the temperature-sensitive sacrificial fiber scaffold, thereby forming a biomimetic endometrial scaffold with a secondary guiding structure that combines oriented microchannels and oriented microfibers in situ.

[0110] Figure 4 These are stereomicroscopic and SEM images of the nerve conduit with a multi-level biomimetic nerve membrane structure in Embodiment 3 of the present invention.

[0111] Figure 4 In the image, the two leftmost images are stereomicroscopic images, and the two rightmost images are SEM images.

[0112] Depend on Figure 4 As can be seen, the secondary structures of directional microchannels and directional microfibers in the nerve conduit of the present invention are clearly visible.

[0113] Figure 5 Schematic diagrams of different biomimetic inner membrane topologies used in cell experiments.

[0114] To compare with existing biomimetic endometrial topology scaffolds, this invention constructed four types of membrane scaffolds with different topologies: a membrane scaffold without a guiding structure (i.e., a smooth plane, the same below), a membrane scaffold with only a microchannel structure, a membrane scaffold with only a microfiber structure, and a membrane scaffold with a microchannel + microfiber structure (analogous to this invention, the same below).

[0115] The specific preparation methods for these four types of membrane scaffolds are as follows:

[0116] Preparation process of smooth film scaffold: Referring to Example 3, 1.5g of PLCL with a number-average molecular weight of 80,000 was weighed and added to 10mL of dichloromethane (dichloromethane, an organic solvent with faster volatility, was chosen here to accelerate curing). The solution was stirred and dissolved overnight at room temperature to obtain a PLCL solution with a concentration of 15g / mL. The completely dissolved PLCL solution was evenly spread in a smooth 10cm diameter glass petri dish, placed in a fume hood, covered, and allowed to evaporate slowly (to prevent excessive evaporation and the generation of bubbles). After complete drying, anhydrous ethanol was added to the glass dish to moisten it, facilitating the peeling of the PLCL film from the bottom of the glass dish to form a smooth, flat PLCL film.

[0117] Microchannel membrane scaffold fabrication process: As described above, a PLCL solution with a concentration of 15 g / mL was first prepared using dichloromethane as the solvent. This PLCL solution was evenly spread on a foil surface and left to stand in a fume hood for 5-10 minutes to allow slow evaporation. During this process, the PLCL solution gradually transitioned from a highly fluid colloidal state to a semi-cured viscous state. At this point, the material was not fully cured internally and exhibited obvious plastic deformation characteristics. Next, the foil was wrapped around a roller-shaped receiver with the PLCL material layer facing outwards, using a fully cured polyvinylidene fluoride (PVDF) thread (250 μm diameter) as a sacrificial template. After fixing the outlet at a distance of 10 cm from the receiver, the taut PVDF thread was evenly wound around the partially cured PLCL film. The corresponding parameters were: receiver rod rotation speed 10 rpm, X-axis movement speed 0.2 mm / sec, and receiver distance 10 cm, to ensure complete coverage of all PLCL film layers. Subsequently, the end of the PVDF thread was fixed to the roller receiver. The device is rolled and pressed evenly on a smooth aluminum foil surface to further encourage the PVDF lines to embed into the incompletely cured PLCL film. After completion, the entire roller receiver is placed in a fume hood overnight to allow the PLCL film to fully cure. Next, the aluminum foil layer containing the PLCL and PVDF lines is peeled off the roller receiver and then immersed in anhydrous ethanol. The PVDF lines are gently separated from the PLCL film surface, and a regularly arranged microchannel structure is formed on the film surface using a linear extrusion method. Finally, the PLCL film with the microgrooved surface is peeled off the aluminum foil, thus obtaining a film with a microgrooved surface.

[0118] The microfiber membrane scaffold preparation process is as follows: As described above, a PLCL spinning solution with a concentration of 15 g / mL (dissolved in hexafluoroisopropanol) is prepared. Referring to the biomimetic endometrial scaffold preparation process in Example 3, a cylindrical receiving rod with a diameter of 100 mm is first installed on a triaxial receiver, and smooth tin foil is wrapped around the cylindrical receiving rod as a fiber receiving base. The PLCL spinning solution was drawn into a syringe, which was then attached to a micro-injection pump. The syringe needle was placed in a coagulation bath containing 200 mL of anhydrous ethanol, with the coagulation bath outlet positioned 30 cm from the receiving rod. The injection pump was set to a feed rate of 1.5 mL / h, the receiving rod rotation speed of 25 rpm, the x-axis movement speed of 3 mm / sec, and the spinning time of 60 min for wet spinning. The resulting spun fibers were collected onto the receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 15 ± 5 μm and an inter-fiber angle of 35 ± 10°, with a collected thickness of 75 ± 25 μm. A second winding process was then performed on the surface of the first mesh fiber membrane. The receiving parameters were adjusted, with the receiving rod rotation speed at 45 rpm and the x-axis movement speed at 0.2 mm / sec, to wind a second oriented fiber membrane with a fiber diameter of 15 ± 5 μm and an inter-fiber angle within 5°, with a collected thickness of 125 ± 25 μm. The fiber membrane was then peeled off the foil, yielding a membrane scaffold with only a microfiber structure.

[0119] Microchannel + microfiber membrane scaffold preparation process: As described above, in this invention, a PLCL spinning solution with a concentration of 15 g / mL (dissolved in hexafluoroisopropanol) is prepared, and referring to the biomimetic endometrial scaffold preparation process in Example 3, the specific processing steps are as follows: First, the sacrificial material is printed on a cylindrical receiver. 2.0 g of PVA with a number-average molecular weight of 50,000 is weighed and added to the printer barrel. The printer nozzle is adjusted to a distance of 1 cm from the receiver rod, the printing air pressure is set to 1.5 kPa, the barrel heating temperature is set to 180℃, the receiver rod rotation speed is set to 10 rpm, the X-axis movement speed is set to 0.2 mm / sec, and the printing time is set to 10 minutes. Using 3D printing technology, the obtained solidified PVA fibers are uniform and exhibit a spaced, parallel deposition state on the surface of the foil. Subsequently, the oriented fibers (first layer) and the mesh fibers (second layer) are collected sequentially through a wet spinning process. The specific wet spinning process is as follows: The PLCL spinning solution is drawn into a syringe and attached to a micro-injection pump. Next, the syringe needle was placed in a coagulation bath containing 200 mL of anhydrous ethanol, with the outlet of the coagulation bath 30 cm away from the receiving rod. The injection pump speed was set to 1.5 mL / h, the receiving rod rotation speed to 45 rpm, the X-axis movement speed to 0.2 mm / sec, and the spinning time to 60 minutes. Wet spinning was performed under these conditions, and the resulting spun fibers were collected on the receiver for a single winding process. During this process, the first layer of oriented fiber membrane had a fiber diameter of 15 ± 5 μm, an inter-fiber angle of less than 5°, and a collection thickness of 125 ± 25 μm. Subsequently, the receiving process parameters were adjusted: the receiving rod rotation speed was set to 25 rpm, and the X-axis movement speed was set to 3 mm / sec. At this point, a second layer of mesh fiber membrane was obtained, with a fiber diameter of 15 ± 5 μm, an inter-fiber angle of 35 ± 10°, and a collection thickness of 75 ± 25 μm. At this point, PVA lines are embedded within the oriented fiber membrane, while the mesh fiber membrane wraps around the outside of the oriented fiber membrane. The cured fiber membrane is peeled off from the foil surface and immersed in distilled water to remove the PVA template, ultimately yielding a microchannel + microfiber membrane scaffold with embedded parallel microchannels on the oriented fiber membrane.

[0120] Figure 6 Two-dimensional characterization of the non-guided structure membrane scaffold, the microchannel structure membrane scaffold, the microfiber structure membrane scaffold, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0121] Depend on Figure 6As can be seen, the smooth planar membrane scaffold is a flat two-dimensional structure; only the microchannel structure membrane scaffold has parallel-oriented micron-sized (diameter in the range of 250 μm) microchannels; only the microfiber structure membrane scaffold has parallel-oriented micron-sized (diameter in the range of 25 μm) microfibers; and the microchannel + microfiber structure membrane scaffold has both parallel-oriented micron-sized (diameter in the range of 250 μm) microchannels and parallel-oriented micron-sized (diameter in the range of 25 μm) microfibers.

[0122] Figure 7 The study examines the regulation of neuronal cell morphology by non-guided structure membrane scaffolds, microchannel structure membrane scaffolds only, microfiber structure membrane scaffolds only, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0123] The same number of PC12 cells were seeded onto the surfaces of the four membrane scaffolds with different topologies, and cultured under the same conditions for 3 days. Figure 7 As shown in Figure A, immunofluorescence staining and laser confocal microscopy imaging were performed (where DAPI channels represent cell nuclei, β-III-tubulin channels represent neurons, and F-actin channels represent the cytoskeleton). Quantitative statistical analysis of cell morphology showed that both the microchannel + microfiber structure membrane scaffold and the microfiber structure membrane scaffold alone significantly improved the cell aspect ratio. Figure 7 B) and neurite length ( Figure 7 (C). Similarly, quantitative statistics on cell orientation showed that both the microchannel + microfiber structured membrane scaffold and the microfiber structured membrane scaffold alone significantly improved cell orientation (C). Figure 7 (D).

[0124] Figure 8 Schematic diagrams of stents for catheters with only densely cross-arranged microfiber structures, catheters with only loosely cross-arranged microfiber structures, and catheters with a densely cross-arranged microfiber structure inner layer plus a loosely cross-arranged microfiber structure outer layer.

[0125] Preparation method of densely cross-arranged microfiber ducts: First, weigh 1.5g of PLCL with a number-average molecular weight of 80,000 and add it to 10mL of hexafluoroisopropanol. Stir and dissolve overnight at room temperature to obtain a PLCL solution with a concentration of 0.15g / mL. A biomimetic outer membrane scaffold is prepared using single-fiber wet spinning technology in a fume hood at room temperature. A cylindrical receiving rod with a diameter of 4mm is installed on a triaxial receiver. The above PLCL solution is drawn into a syringe, which is then attached to a microinjection pump. The syringe needle is placed in a coagulation bath containing 200mL of anhydrous ethanol. By adjusting the y-axis of the receiver, placing the coagulation bath outlet 25cm away from the receiver rod, setting the injection pump speed to 1mL / h, the receiver rod rotation speed to 500rpm, the x-axis movement speed to 1.5mm / sec, the spinning time to 10min, and the inner layer thickness to 100μm, a fiber inner layer with a fiber diameter of 7.5±2.5μm and an inter-fiber angle of 10±5° is obtained. The fiber tube wall formed by the fine fibers crossing at small angles is obtained after the glass is removed from the receiver rod, resulting in a densely cross-arranged microfiber structure conduit.

[0126] Preparation method of loosely cross-arranged microfiber conduits: As described above, the distance between the coagulation bath outlet and the receiving rod is adjusted to 2 cm, the injection pump speed is set to 8 mL / h, the receiving rod rotation speed is 100 rpm, the x-axis movement speed is 8 mm / sec, the spinning time is 20 min, and the outer layer thickness is 200 μm, resulting in a fiber outer layer with a fiber diameter of 25 ± 5 μm and an inter-fiber angle of 35 ± 5°. After the fiber tube wall formed by the large-angle cross-arrangement of coarse fibers is peeled off from the receiving rod, a loosely cross-arranged microfiber conduit is obtained.

[0127] Figure 9 Compression cycle curves (A) and compressive stress changes (B) of dense membrane structure catheters, loose membrane structure catheters, and double-layer membrane structure catheters of Example 3.

[0128] from Figure 9 It is known that the catheter with the double membrane structure of the present invention has the strongest compression resistance compared with the other two types of catheters. After 100 compressions, its radial force still exceeds 8N, which can effectively avoid the reduction of lumen volume caused by muscle compression.

[0129] Figure 10 This invention illustrates the effect of each fibrous membrane layer of the duct constituting the double membrane structure on the migration behavior of fibroblasts in Embodiment 3 of the present invention. Figure 10 A is a schematic diagram of the experimental procedure for evaluating the ability of fibroblasts to migrate through the fibrous membrane using the Transwell system; Figure 10 B represents a comparison of fibroblast migration on different scaffold structures; Figure 10 C represents the statistical results of the porosity of the inner and outer fiber membranes; Figure 10 D represents the quantitative analysis result of cell migration. (From...) Figure 10 It is known that fibrous membranes composed only of densely cross-arranged fibers have low porosity, while fibrous membranes composed only of loosely cross-arranged fibers have high porosity; adding an inner fibrous membrane to the duct can significantly block the infiltration and migration of fibroblasts.

[0130] Figure 10 In this context, Inner represents a dense membrane structure conduit, Outer represents a loose membrane structure conduit, and Inner & Outer represent the microchannel + microfiber structure membrane scaffold of Example 3.

[0131] Figure 11 This invention illustrates the effect of the double-layered fibrous membrane constituting the outer sheath of the catheter on the permeability of nutrients in Embodiment 3. Figure 11 Figure A is a schematic diagram of the experimental procedure for evaluating bovine serum albumin (BSA) permeation through a bilayer fiber membrane using the Transwell system, which is used to detect the molecular permeability of the bilayer fiber membrane. Figure 11 B represents the statistical curves showing the change of BSA concentration in the upper and lower chambers over time. Figure 11 It is evident that the bilayer fiber membrane has excellent molecular permeability, and the BSA added to the upper chamber can penetrate the bilayer fiber membrane and permeate into the lower chamber in a short time.

[0132] Figure 12 The image shows the bending resistance of the multi-level biomimetic neural membrane structure in Embodiment 3 of the present invention.

[0133] Depend on Figure 12 It can be seen that its maximum bending angle reaches 180°. Even after double bending, it can still exhibit bending resistance at the bending point, which can effectively avoid the reduction of lumen volume caused by cross-joint repair.

[0134] exist Figure 5 and Figure 8 Based on this, four biomimetic inner membranes with different topological structures were rolled along the axial direction of oriented microchannels and loaded into a biomimetic outer membrane catheter with a double-layer structure, further constructing four three-dimensional catheter scaffolds with different topological structures. These are a membrane scaffold without a guide structure (i.e., a smooth plane, the same below), a membrane scaffold with only a microchannel structure, a membrane scaffold with only a microfiber structure, and a membrane scaffold with a microchannel + microfiber structure (i.e., the product of this invention, the same below).

[0135] Figure 13 The diagram shows the structure of the membrane scaffold without a guide structure, the membrane scaffold with only a microchannel structure, the membrane scaffold with only a microfiber structure, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0136] Figure 14Stereoscopic and SEM images (A) and internal porosity statistics (B) of the non-guided structure membrane scaffold (i.e., smooth plane, the same below), the microchannel structure membrane scaffold only, the microfiber structure membrane scaffold only, and the microchannel + microfiber structure membrane scaffold of Example 3.

[0137] Depend on Figure 14 It can be seen that after assembly, the microchannel + microfiber scaffold still retains its original secondary guiding structure. At the same time, its porosity is significantly higher than the other three structures.

[0138] Catheters with different filling structures were cut into 2cm lengths and 2mm inner diameters, sterilized, and then implanted subcutaneously into the backs of rats for the following tests.

[0139] Figure 15 H&E and DAPI staining images of the non-guided structure membrane scaffold, the microchannel structure membrane scaffold, the microfiber structure membrane scaffold, and the microchannel + microfiber structure membrane scaffold of Example 3 after subcutaneous implantation in rats 7 days and 14 days later (Figure A), and the corresponding cell infiltration count statistics (Figure B).

[0140] Figure 15 Quantitative statistics on the number of infiltrating cells show that the microchannel + microfiber structure membrane scaffold of the present invention can significantly increase the number of infiltrating cells compared with the other three structures.

[0141] Figure 16 This study evaluates the efficacy of the biomimetic outer membrane scaffold of Example 3, the microchannel + microfiber structure membrane scaffold of Example 3, and autologous nerve in repairing a 2cm sciatic nerve defect model in rats. In the figures, A shows immunofluorescence imaging of the regenerated nerve tissue, B shows transmission electron microscopy imaging of the distal myelin sheath structure of the regenerated nerve, and C, D, and E show the electrophysiological function recovery assessments of the hollow catheter, the microchannel + microfiber structure membrane scaffold of Example 3, and the autologous nerve, respectively.

[0142] Figure 16 This study demonstrates the efficacy of three different types of nerve conduits—the biomimetic outer membrane scaffold (hollow conduit) of Example 3, the biomimetic outer membrane scaffold + biomimetic inner membrane scaffold (the multi-level biomimetic membrane structure proposed in this invention), and autologous nerve (positive control, the gold standard for transplantation in clinical practice)—in repairing a 2cm sciatic nerve defect model in rats. This length exceeds the critical size (approximately 1.5cm) reported in existing technologies for rat nerve injury repair, making complete functional recovery difficult to achieve through conventional repair methods.

[0143] (A) Immunofluorescence imaging of regenerated nerve tissue

[0144] Tissue sections of regenerated nerves were observed after immunofluorescence staining. DAPI labeled cell nuclei (blue), S100β labeled Schwann cells (red), and NF-160 labeled axons (green). The figures show the morphological changes in the proximal, bridge, and distal regions of the hollow conduit, microchannel + microfiber conduit, and autologous nerve repair groups, respectively. In the hollow conduit group, axonal regeneration was sparse and disordered; in the microchannel + microfiber conduit group, regenerated axons extended along the channel direction within the bridge region, exhibiting good guidance and uniform axonal distribution; while the autologous nerve group showed the continuity and maturity most closely resembling the natural nerve structure.

[0145] (B) Transmission electron microscopy (TEM) imaging of the distal myelin sheath structure of the regenerating nerve.

[0146] TEM images revealed three groups of myelin sheath structures formed in the distal nerve region. The hollow duct group showed sparse, unmyelinated, or incompletely myelinated axons; the microchannel + microfiber duct group showed numerous regularly myelinated axons, suggesting an effective nerve remyelination process; and the autologous nerve group showed the most typical thick myelin sheath structure, consistent with the normal nerve morphology.

[0147] (CE) Electrophysiological Function Recovery Assessment

[0148] Electrophysiological recordings were performed on the sciatic nerve regeneration area of ​​three groups of animals to assess the recovery of nerve function. Figure 16 The C-shaped display showed that almost no evoked potentials were detected in the bionic outer membrane stent assembly of Example 3, indicating extremely poor functional recovery; Figure 16 D is a microchannel + microfiber catheter group, which can record clear evoked potential waveforms, indicating that nerve conduction function has been partially restored; Figure 16 E represents the autoneuron, exhibiting a strong evoked potential response close to the level of normal nerve conduction.

[0149] Figure 17 This study evaluates the effectiveness of the microchannel + microfiber structure membrane scaffold and autologous nerve in a 5cm sciatic nerve defect repair model in a large animal model, as described in Example 3. A shows intraoperative observation, images taken 6 months post-operation, and intraoperative TEM imaging; B shows MRI evaluation; C shows electrophysiological function testing; and D shows gait analysis and motor function recovery.

[0150] Figure 17This study demonstrates the regeneration and functional recovery of a 5cm sciatic nerve defect repaired using the "microchannel + microfiber catheter" method proposed in this invention in a large animal model (beagle), and compares it with that of an autologous nerve graft. This defect length far exceeds the clinical threshold of 4cm. According to existing reports, such large nerve defects typically require 12-18 months or more to observe significant motor function recovery (muscle strength M3 or higher, i.e., the ability to generate voluntary movement against gravity). However, this invention was evaluated only 6 months post-surgery and has already shown significant repair effects.

[0151] (A) Intraoperative observation, 6-month postoperative images and intraoperative TEM imaging

[0152] During the surgery, the catheter was precisely implanted into the 5cm defect gap. Six months after the surgery, when the nerve area was exposed again, obvious continuous nerve bundle regeneration was observed. The structure was full and the morphology was similar to that of the autologous transplant group.

[0153] TEM imaging: Ultrastructural observation results showed that the microchannel + microfiber conduit repair group had formed nerve axons wrapped with mature myelin sheaths. The myelin sheath layers were clear and the structure was regular, similar to the autologous nerve group, suggesting that myelination and nerve reconstruction had reached a high level.

[0154] (B) MRI assessment

[0155] Six months post-surgery, MRI imaging further confirmed complete perforation of the nerve tissue within the duct. The image shows symmetrical nerve bundle images on both the duct repair side (NS) and the contralateral healthy side (OS), indicating correct nerve regeneration direction and good morphological recovery.

[0156] (C) Electrophysiological function testing

[0157] In the electrophysiological atlas, the microchannel + microfiber catheter repair group induced a clear distal compound muscle action potential (CMAP) upon proximal stimulation. Although the amplitude was slightly lower than that of the autologous nerve group, the conduction delay was significantly shortened, indicating the presence of a relatively complete functional conduction pathway. This demonstrates that the catheters effectively promoted neural reconnection and functional reconstruction.

[0158] (D) Gait analysis and motor function recovery

[0159] The walking process of beagle dogs was recorded using high frame rate video and gait analysis was performed. Six months post-surgery, the microchannel + microfiber catheter group was able to achieve continuous and stable walking, with the affected hind limb actively lifting, supporting, and propelling itself, exhibiting a complete gait cycle. Similar to the movement pattern in the autologous nerve group, the movement trajectory was continuous and the stride was coordinated, indicating a significant recovery of motor function, reaching the clinical M3 level of voluntary movement standards, far exceeding the expected recovery period.

[0160] Figure 18 A heatmap showing the comprehensive scoring of five key performance indicators for various nerve grafts that can be used to repair long-segment nerve defects.

[0161] Figure 18 This paper presents a comprehensive score heatmap of five key performance indicators for various neural grafts that can be used to repair long-segment nerve defects, as reported in existing literature up to May 10, 2025. Figure 18 In this diagram, each dot represents a neural conduit material or configuration from a specific literature review, with a scoring range of 0-5. The color gradient transitions from dark purple (low score) to bright yellow (high score), facilitating a visual comparison of the performance advantages and disadvantages of different solutions. This diagram aims to outline the key technological trends and shortcomings in the current development of neural conduits. Figure 18 A comparative analysis of current artificial nerve conduit designs across five key dimensions revealed that structural guidance and repair length are the main technical barriers to distinguishing conduit performance. Most conduits still suffer from uneven mechanical properties and limited structural functions. The microchannel + microfiber conduit proposed in this invention (“This work”) scored higher in multiple of the five dimensions, demonstrating excellent mechanical-structural-biological functional integration capabilities. It is particularly suitable for repairing long-segment nerve defects and has strong clinical translational potential.

[0162] Figure 18 In the table, the specific reference names corresponding to the horizontal axis numbers 1-55 are as follows:

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[0217] The embodiments described above are merely preferred embodiments of the present invention and are not intended to limit the scope of the present invention. Various modifications and improvements made by those skilled in the art to the technical solutions of the present invention without departing from the spirit of the present invention should fall within the protection scope defined by the claims of the present invention.

Claims

1. A nerve conduit with a multi-level biomimetic nerve membrane structure, characterized in that, It includes a biomimetic inner membrane scaffold with a secondary guiding structure and a biomimetic outer membrane scaffold with a double-layer structure; The preparation method of the biomimetic endometrial scaffold with a secondary guiding structure includes the following steps: (1) Spinning is performed using spinning solution to obtain spun fibers, and the spun fibers are collected on the receiver and wound once to obtain a first mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°; then a second winding process is performed on the surface of the first mesh fiber membrane to prepare a second oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°, thus obtaining a receiver with oriented microfiber membrane material; (2) The sacrificial layer material is loaded onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material to obtain an oriented microfiber membrane with a spirally loaded oriented sacrificial fiber scaffold. (3) Spinning is carried out using spinning solution to obtain spun fibers. Then, the spun fibers are wound three times on the surface of the oriented microfiber membrane loaded with oriented sacrificial fiber scaffold to prepare a third oriented fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 0-5°. Then, a fourth winding process is carried out to prepare a fourth mesh fiber membrane with a fiber diameter of 1-50 μm and an inter-fiber angle of 15-45°, thus obtaining a multilayer fiber membrane that wraps the oriented sacrificial fiber scaffold. (4) Cut the multilayer fiber membrane that encapsulates the directional sacrificial fiber scaffold along the receiver axis, and then remove the directional sacrificial fiber scaffold, leaving the directional microchannel formed by the sacrificial fiber in the fiber membrane, to obtain the biomimetic endometrial scaffold with a secondary guiding structure. The method for preparing the neural conduit with the multi-level biomimetic neural membrane structure includes the following steps: After the bionic endometrial scaffold with the secondary guiding structure is rolled up along the fiber direction in the directional fiber membrane, it is inserted into the middle cavity of the bionic outer membrane scaffold with the double-layer structure to obtain the nerve conduit of the multi-level bionic nerve membrane structure. The directional microchannels are arranged at an angle of 0-5° along the long axis of the nerve conduit.

2. The nerve conduit with a multi-level biomimetic nerve membrane structure according to claim 1, characterized in that, The thickness ratio of the first reticular fiber membrane to the second oriented fiber membrane is 0.5-1, and the thickness of the first reticular fiber membrane is 50-100 μm; the thickness ratio of the third oriented fiber membrane to the fourth reticular fiber membrane is 1-2, and the thickness of the fourth reticular fiber membrane is 50-100 μm.

3. The nerve conduit with a multi-level biomimetic nerve membrane structure according to claim 1, characterized in that, In step (2): the loading method is 3D printing, which specifically includes the following steps: loading the sacrificial layer material into the 3D printer, performing 3D printing, loading the printed sacrificial fibers onto the surface of the oriented microfiber membrane of the receiver with the oriented microfiber membrane material, and during 3D printing, the sacrificial fibers are spirally loaded on the surface of the receiver by moving the receiver along the axial direction and rotating the receiver to obtain an oriented microfiber membrane loaded with an oriented sacrificial fiber support.

4. The nerve conduit with a multi-level biomimetic nerve membrane structure according to claim 3, characterized in that, The sacrificial layer material is a water-soluble polymer or a temperature-sensitive polymer; the printing air pressure of the 3D printing is 0.1-50 kPa, the barrel heating temperature is 30-200 ℃, the receiver rotation speed is 1-400 rpm, the receiver axial movement speed is 1-100 mm / sec, and the printing time is 1-100 min.

5. The nerve conduit with a multi-level biomimetic nerve membrane structure according to claim 1, characterized in that, The diameter of the sacrificial fiber is 100-1000 μm; the proportion of the oriented sacrificial fiber scaffold in the multilayer fiber membrane that encapsulates the oriented sacrificial fiber scaffold is 50-70%.

6. The nerve conduit with a multi-level biomimetic nerve membrane structure according to claim 1, characterized in that, The fabrication method of the biomimetic outer membrane scaffold with the double-layer structure includes the following steps: A first spinning fiber is obtained by spinning using a first spinning solution, and the first spinning fiber is collected on a receiver for a first winding process to obtain an inner fiber layer with a fiber diameter of 0.1-10 μm and an inter-fiber angle of 5-15°. A second spinning fiber is obtained by spinning using a second spinning solution, and then the second spinning fiber is used to perform a second winding process on the surface of the inner fiber layer to wind an outer fiber layer with a fiber diameter of 10-100 μm and an inter-fiber angle of 15-45°, thus obtaining the biomimetic outer membrane scaffold with a double-layer structure.

7. The application of a nerve conduit with a multi-level biomimetic nerve membrane structure as described in any one of claims 1-6 in the field of nerve conduits.

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