Preparation method of conductive core-shell nanofiber membrane applied to flexible electrode and nerve conduit

Abstract

According to the invention, a bicontinuous phase conductive core-shell nanofiber membrane is applied to construction of a flexible neural electrode and a neural conduit, firstly, a bicontinuous phase conductive polymer system is constructed, and the system realizes a three-dimensional network in which an electric phase and a mechanical phase are respectively continuous, interpenetrating but independent; and the long-term technical problems of conductive component agglomeration, conductive path interruption and the like are effectively solved. And preparing the conductive nanofiber scaffold with a core-shell structure by taking a high-molecular polymer as a core layer material and a bicontinuous-phase conductive polymer system as a shell layer material through a coaxial electrostatic spinning method. The construction mode effectively establishes a stable, continuous and uniform electron conduction path, can realize effective nerve electrical stimulation and high-fidelity electrophysiological signal recording in vivo, and can be used as a nerve conduit to regulate and repair related bio-electricity signals in a nerve repair process, so as to accelerate regeneration and repair of nerves.

Description

Technical Field

[0001] This invention relates to the field of nerve conduit technology, specifically to a method for preparing a conductive core-shell nanofiber membrane that combines the applications of flexible electrodes and nerve conduits. Background Technology

[0002] Peripheral nerve injury (PNI) is a common and serious clinical condition that significantly impacts quality of life, affecting over one million people worldwide each year. Despite advancements in clinical treatments, autologous nerve transplantation remains the gold standard for long-distance, complex PNI. However, this method has significant limitations, including a limited number of donor nerves, impaired donor site function, and postoperative complications. Therefore, developing tissue-engineered neural conduits (NGCs) with biomimetic functions and regeneration-promoting capabilities has become a core direction for alternatives to autologous transplantation. Currently, NGC development focuses on two key properties: structural guidance and electrical conductivity. On one hand, axonal regeneration is highly dependent on directional guidance; by constructing conduit scaffolds with micro / nanostructures, the microenvironment of natural neural tissue can be effectively simulated, preventing axonal aberration. On the other hand, the nerve regeneration process inherently relies on the regulation of bioelectrical signals; therefore, electrically active conductive materials are introduced into NGCs to enhance intercellular electrical signal transmission and activate nerve regeneration-related signaling pathways. Meanwhile, neural repair strategies increasingly emphasize interfaces with bioelectronic functions, enabling targeted electrical stimulation and real-time recording of neural activity, thereby accelerating regeneration and improving long-term recovery. Flexible nanofibers are particularly suitable for integration into regenerative conduits and bioelectronic interfaces. Their compliance allows for close adhesion to soft, dynamic neural tissue, reducing interfacial stress, inflammation, and fibrosis while achieving high-fidelity charge transfer and signal capture, offering significant advantages over rigid metal or silicon-based electrodes. Existing conductive scaffolds are mostly based on nanofillers (such as carbon nanotubes and graphene) or conductive polymers (such as polypyrrole, polyaniline, and PEDOT:PSS). However, high-content nanofillers are prone to aggregation and biocompatibility issues; conductive polymer coatings have limited adhesion and poor stability under physiological conditions; and blending electrospinning often leads to a decrease in fiber flexibility and mechanical properties. Therefore, a design strategy that ensures both conductivity and good processability and compliance is urgently needed. In view of this, this invention is proposed. Summary of the Invention

[0003] The purpose of this invention is to provide a method for preparing a conductive core-shell nanofiber membrane that combines the applications of flexible electrodes and nerve conduits, in order to solve the bottleneck problems of existing conductive core-shell nanofiber membranes, such as difficulty in dispersing conductive components, discontinuous construction of conductive networks, imbalance between mechanical and electrical properties, and poor biocompatibility.

[0004] To achieve one of the above objectives, the present invention provides the following technical solution:

[0005] A method for preparing a conductive core-shell nanofiber membrane that can be used as both a flexible electrode and a nerve conduit includes the following steps:

[0006] S100. Stir the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution at room temperature, filter to obtain filtrate, freeze the filtrate with liquid nitrogen and place it in a freeze dryer to obtain PEDOT:PSS nanofibers.

[0007] S200. PEDOT:PSS nanofibers and hydrophilic PU are dissolved in a 70% ethanol solution at a weight ratio of 2-4:7 to prepare a bicontinuous phase conductive polymer solution with a mass concentration of 0.5-10%.

[0008] S300. A conductive core-shell nanofiber membrane is prepared by coaxial electrospinning using a bicontinuous phase conductive polymer solution as the shell solution and a polymer solution as the core solution, wherein the polymer solution is one or more of PCL solution, P34HB solution, and PLGA solution.

[0009] S400. The conductive core-shell nanofiber membrane is cut into long strips that can be wrapped around nerves to obtain flexible nerve electrodes.

[0010] S500: A conductive core-shell nanofiber membrane is wrapped into a tube to obtain a composite nerve conduit that combines structural guidance and conductivity.

[0011] Preferably, in step S100:

[0012] The stirring time at room temperature was 6 hours; the filtration accuracy was 10 μm; the liquid nitrogen freezing time was 15 minutes; and the freeze-drying time was 72 hours.

[0013] Preferably, in step S200:

[0014] PEDOT:PSS nanofibers and hydrophilic PU are mixed in a weight ratio of 3:7; the mass concentration of the bicontinuous phase conductive polymer solution is 3%.

[0015] Preferably, in step S300:

[0016] The mass concentration of the PCL polymer solution is 12%.

[0017] Preferably, in step S300:

[0018] The conductive core-shell nanofiber membrane has a bilayer polymer fiber layer with an inner unidirectional ordered layer and an outer layer with random arrangement.

[0019] Preferably, in step S500:

[0020] The conductive core-shell nanofiber membrane was bonded using a 12% PCL polymer solution.

[0021] To achieve the second objective mentioned above, the present invention provides the following technical solution:

[0022] A conductive core-shell nanofiber membrane that can be used as both a flexible electrode and a nerve conduit is prepared by the aforementioned method.

[0023] To achieve the third objective mentioned above, the present invention provides the following technical solution:

[0024] A conductive core-shell nanofiber membrane prepared by the method described above, which can be used as both a flexible electrode and a nerve conduit, or the application of the conductive core-shell nanofiber membrane, which can be used as both a flexible electrode and a nerve conduit, in flexible nerve electrodes and nerve conduits in the treatment of peripheral nerve injuries.

[0025] Compared with existing technologies, this invention proposes to apply a bicontinuous phase structure to the construction of nerve conduits. First, a bicontinuous phase conductive polymer system (PEDOT:PSS / PU, PPP) is constructed by microphase separation of PEDOT:PSS and hydrophilic polyurethane (PU) in 70% ethanol solvent. This system achieves a three-dimensional network where the electrical and mechanical phases are continuous, interpenetrating, but independent, effectively solving long-standing technical problems such as the aggregation of conductive components and the interruption of conductive pathways. Then, using FDA-approved polycaprolactone (PCL), poly-3-hydroxybutyrate-4-hydroxybutyrate (P34HB), and polylactic-glycolic acid copolymer (PLGA) as core layer materials and the bicontinuous phase conductive polymer system as the shell material, a core-shell structured conductive nanofiber scaffold is prepared by one-step coaxial electrospinning. Each nanofiber prepared by coaxial spinning is uniformly coated with a bicontinuous phase conductive polymer layer, achieving continuous, uniform, and stable deposition of the conductive layer on the fiber surface, resulting in a stable balance of electrical and mechanical properties at the microscale. This construction method effectively establishes a continuous and uniform electronic conduction pathway, providing cells with persistent electroactive stimulation signals during neural repair while simultaneously providing the necessary mechanical support environment. Furthermore, the ordered arrangement of the inner layer of the fabricated nanofiber scaffold provides excellent topological cues for cell adhesion, migration, and axonal guidance. We also found that the excellent electrochemical properties enabled the bicontinuous conductive core-shell nanofiber scaffold to achieve effective neural stimulation and high-quality electrophysiological signal recording of the rat sciatic nerve and rhesus monkey median nerve, demonstrating its great potential as a neuro-electronic interface material.

[0026] To make the above-mentioned objects, features and advantages of the present invention more apparent and understandable, preferred embodiments are described below in detail with reference to the accompanying drawings. Attached Figure Description

[0027] To more clearly illustrate the technical solutions of the embodiments of the present invention, the accompanying drawings used in the embodiments will be briefly introduced below. It should be understood that the following drawings only show some embodiments of the present invention and should not be regarded as a limitation on the scope. For those skilled in the art, other related drawings can be obtained based on these drawings without creative effort.

[0028] Figure 1 A schematic diagram illustrating the design and characterization of the composite nerve conduit prepared in this invention, which combines structural guidance and electrical conductivity.

[0029] Figure 2 This diagram illustrates the proliferation and myelination of Schwann cells (SCs) on the composite neural conduit prepared in this invention, which combines structural guidance and electrical conductivity.

[0030] Figure 3 This is a schematic diagram illustrating how the composite nerve conduit prepared according to the present invention, which combines structural guidance and electrical conductivity, promotes nerve axon extension.

[0031] Figure 4 This is a schematic diagram of the in vivo nerve stimulation and composite muscle action potential recording of the composite nerve conduit prepared in this invention, which combines structural guidance and conductivity.

[0032] Figure 5 This is a schematic diagram of the regeneration and nerve function recovery of the composite nerve conduit prepared in this invention, which combines structural guidance and electrical conductivity.

[0033] Figure 6 This is a schematic diagram of the immunofluorescence staining of the regenerated nerve of the composite nerve conduit prepared in this invention, which combines structural guidance and conductivity.

[0034] Figure 7 This is a schematic diagram of the composite nerve conduit myelin sheath and gastrocnemius muscle regeneration prepared according to the present invention, which combines structural guidance and electrical conductivity. Detailed Implementation

[0035] The technical solutions of the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. The components of the embodiments of the present invention described and shown in the accompanying drawings can generally be arranged and designed in various different configurations. Therefore, the following detailed description of the embodiments of the present invention provided in the accompanying drawings is not intended to limit the scope of the claimed invention, but merely to illustrate selected embodiments of the invention. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without inventive effort are within the scope of protection of the present invention.

[0036] Example 1

[0037] S100. Stir the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution at room temperature for 6 hours, filter with a precision of 10 μm to obtain filtrate, freeze the filtrate with liquid nitrogen for 15 minutes and then place it in a freeze dryer for 72 hours to obtain PEDOT:PSS nanofibers.

[0038] S200. PEDOT:PSS nanofibers and hydrophilic PU are dissolved in a 70% ethanol solution at a weight ratio of 3:7 to prepare a 3% bicontinuous phase conductive polymer solution.

[0039] S300. A conductive core-shell nanofiber membrane was prepared by coaxial electrospinning using a bicontinuous phase conductive polymer solution as the shell solution and a 12% PCL polymer solution as the core solution.

[0040] S400. A conductive core-shell nanofiber membrane is wound into a tube to obtain a composite nerve conduit that combines structural guidance and conductivity. The conductive core-shell nanofiber membrane is bonded using a 12% PCL polymer solution.

[0041] Example 2

[0042] S100. Stir the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution at room temperature for 6 hours, filter with a precision of 10 μm to obtain filtrate, freeze the filtrate with liquid nitrogen for 15 minutes and then place it in a freeze dryer for 72 hours to obtain PEDOT:PSS nanofibers.

[0043] S200. PEDOT:PSS nanofibers and hydrophilic PU are dissolved in a 60% ethanol solution at a weight ratio of 2:7 to prepare a 0.5% bicontinuous phase conductive polymer solution.

[0044] S300. A conductive core-shell nanofiber membrane was prepared by coaxial electrospinning using a bicontinuous phase conductive polymer solution as the shell solution and a 10% P34HB polymer solution as the core solution.

[0045] S400. A conductive core-shell nanofiber membrane is wound into a tube to obtain a composite nerve conduit that combines structural guidance and conductivity. The conductive core-shell nanofiber membrane is bonded using a 10% PCL polymer solution.

[0046] Example 3

[0047] S100. Stir the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution at room temperature for 6 hours, filter with a precision of 10 μm to obtain filtrate, freeze the filtrate with liquid nitrogen for 15 minutes and then place it in a freeze dryer for 72 hours to obtain PEDOT:PSS nanofibers.

[0048] S200. PEDOT:PSS nanofibers and hydrophilic PU are dissolved in an 80% ethanol solution at a weight ratio of 4:7 to prepare a 10% bicontinuous phase conductive polymer solution.

[0049] S300. A conductive core-shell nanofiber membrane was prepared by coaxial electrospinning using a bicontinuous phase conductive polymer solution as the shell solution and a PLGA polymer solution with a mass concentration of 20% as the core solution.

[0050] S400. A conductive core-shell nanofiber membrane is wound into a tube to obtain a composite nerve conduit that combines structural guidance and conductivity. The conductive core-shell nanofiber membrane is bonded using a 20% PCL polymer solution.

[0051] In the above embodiments, all reagents were commercially available products. Specifically, the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution (PEDOT:PSS solution) and the hydrophilic PU were both commercially available products. Specifically, the PEDOT:PSS solution was purchased from Heraeus, model PH1000; the hydrophilic PU was purchased from AdvanSource biomaterials, model HydroMed D3.

[0052] To further illustrate the beneficial effects of the present invention, the composite nerve conduit prepared in Example 1 was studied and analyzed.

[0053] In this invention, to prepare a composite nerve conduit possessing both structural guidance and electrical conductivity, we first prepared a bicontinuous conductive polymer solution. Due to the different solubilities of hydrophilic polyurethane (PU) and PEDOT:PSS in ethanol and water, hydrophilic polyurethane exhibits phase separation in mixed solvents with ethanol concentrations less than 70 v / v%, while PEDOT:PSS shows phase separation in mixed solvents with ethanol concentrations above 70 v / v%. Notably, both hydrophilic polyurethane and PEDOT:PSS exhibit moderate phase separation in 70 v / v% ethanol solvent, without significant aggregation or precipitation, demonstrating a unique phase separation between the mechanical phase (hydrophilic PU) and the electrical phase (PEDOT:PSS). In contrast, conventional hydrophilic polymers that can be used as the mechanical phase (e.g., polyvinyl alcohol, polyacrylamide, polyacrylic acid) are soluble in water, resulting in uniform dispersion in the polymer solution without pre-forming phase separation between the mechanical and electrical phases, making it difficult to form interconnected mechanical and electrical phases. Therefore, hydrophilic PU and PEDOT:PSS were dispersed in 70 v / v% ethanol solvent. After the solvent evaporated, the phase-separated hydrophilic PU and PEDOT:PSS were compacted to form a polymer film (PPP) with mechanical and electrical phases, respectively. To simulate the physiological environment in vivo, PPP could be immersed in PBS for equilibration before use. Swelling results showed that PPP could rapidly reach swelling equilibrium within 1 hour, with a swelling rate of 223.68 ± 3.73%. Figure 1 a). After swelling equilibrium, the modulus of PPP decreased from 67.8±5.9 MPa to 1±0.4 MPa, forming a soft and stable polymer film with a modulus comparable to that of soft biological tissue. Figure 1 b).

[0054] To achieve high connectivity between the mechanical and electrical phases simultaneously, we investigated the electrical conductivity and tensile properties of PPP with different PEDOT:PSS / PU (w / w%) ratios. Figure 1 a and Figure 1 Results showed that with increasing PEDOT:PSS / PU (w / w%) ratio, the electrical conductivity of PPP increased from 0.007±0.0005 to 198.1±34.6 S / m, while the tensile properties decreased from 670% to 70%. To determine the formation of the bicontinuous phase, we used AFM tapping mode to distinguish the microstructures of the hydrophilic PU and PEDOT:PSS phases after separation. In AFM phase image analysis, the higher modulus PEDOT:PSS appeared as a brighter region, while the lower modulus hydrophilic polyurethane appeared as a darker region. AFM results showed that PPP with a low PEDOT:PSS / PU (w / w%) ratio (20 / 80) exhibited low connectivity between the electrical phases, while PPP with a high PEDOT:PSS / PU (w / w%) ratio (40 / 60) exhibited low connectivity between the mechanical phases. Figure 1 c shows that when the PEDOT:PSS / PU (w / w%) ratio is 30 / 70, the electrical phase PEDOT:PSS and the mechanical phase hydrophilic polyurethane are uniformly distributed in PPP, exhibiting a dual continuity of the mechanical and electrical phases. As a result, PPP not only achieves high conductivity but also maintains its mechanical integrity. Figure 1 a and Figure 1 b also shows that when the PEDOT:PSS / PU (w / w%) ratio is 30 / 70, PPP simultaneously achieves high conductivity of 77.1±11.6 and stretchability of 200%. The surface structure of PPP was examined using scanning electron microscopy (SEM), and the results are as follows... Figure 1 As shown in Figure d, SEM revealed an interconnected network on the PPP surface, exhibiting a nano-wrinkled topology. Notably, after immersing PPP in 90% ethanol for 24 hours to remove the mechanical phase by dissolving the hydrophilic PU in PPP, a stable, independent PEDOT:PSS film was still obtained. Energy dispersive spectroscopy (EDS) analysis confirmed that after removing the mechanical phase, the sulfur content, a characteristic element of PEDOT:PSS, increased from 3.09% to 19.73%, further confirming the existence of a bicontinuous electrical and mechanical phase in PPP. Shear-dependent viscosity profiles showed that the viscosity of the PPP polymer solution could be easily adjusted by changing the amount of solvent. Moreover, conductivity tests showed no significant change in conductivity at different PPP concentrations. This indicates that the physicochemical properties of PPP remain stable after the formation of a bicontinuous phase, demonstrating its advantage in manufacturing flexibility.

[0055] Numerous studies have shown that the introduction of excessive conductive polymers can produce severe biotoxicity in vivo. To reduce the PPP content in the electrospun fiber scaffold while maintaining the bicontinuous phase structure of PPP, we used coaxial electrospinning technology to encapsulate the bicontinuous phase structure of PPP onto the outer layer of PCL, preparing core-shell structured PCL / PPP nanofibers. To determine the optimal concentration of PPP for promoting neurorepair, we performed coaxial electrospinning with PCL at concentrations of 3%, 5%, and 7%. By controlling the speed of the receiving roller, we prepared inner-layer oriented nanofibers. Figure 1The core-shell structure consists of PCL / PPP nanofibers with an outer layer of random arrangement. The orderly arranged nanofibers provide physical topological guidance for cell migration and growth, while the random nanofibers facilitate the free exchange of nutrients and metabolic waste. Nanofiber diameter statistics show that the average diameters of PCL / PPP nanofibers are 240.8±41.2 nm, 249.7±62.2 nm, 253.8±37.3 nm, and 264.2±70.1 nm, respectively. To determine the formation of the core-shell structure of the PCL / PPP nanofibers, we observed their microstructure using transmission electron microscopy. Figure 1 The results showed that PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofibers all possessed a core-shell structure. EDS elemental analysis also revealed a uniform distribution of PEDOT:PSS and the characteristic S and N elements of hydrophilic PU on the surface of the PCL / PPP nanofibers and nanofiber membranes, indicating the successful deposition of PPP on the PCL nanofiber surface.

[0056] NGCs must meet the in vivo mechanical strength requirements for the long-term repair process of PNIs. To examine the mechanical properties of nanofiber scaffolds, we performed tensile stress-strain tests on PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. Figure 1 The f-values ​​show that compared to PCL (19.55±1.50), PCL / 3%PPP (27.25±0.83), PCL / 5%PPP (32.60±1.07), and PCL / 7%PPP (39.32±0.82) exhibit enhanced tensile strength. This provides unimpeded circulation of neurotrophic factors and nutrients, as well as stable mechanical support to prevent NGC collapse during the most critical phase of in vivo neural repair (the first two months). However, the mechanical strength, including Young's modulus, tensile strength, and fracture strain, of PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP did not change significantly, indicating the stability of the physicochemical properties of the bicontinuous phase structure. Furthermore, the introduction of PPP significantly improved the surface hydrophilicity of the nanofiber scaffold, which helps promote cell adhesion and neurite extension.

[0057] PCL / PPP conductive nanofibers with a bicontinuous phase structure provide an efficient, uniform, and continuous pathway to accelerate electron transfer, thus providing electroactive signals for neural repair. To evaluate the electrochemical performance of PCL / PPP nanofibers, we first used a four-probe method to compare the conductivity of PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. Figure 1 The results showed that the introduction of PPP effectively enhanced the conductivity of the PCL nanofiber scaffold. The conductivity of the PCL / PPP nanofiber scaffold (PCL / 3%PPP: 2.36±0.66 S / m, PCL / 5%PPP: 2.44±0.51 S / m, PCL / 7%PPP: 2.90±0.81 S / m) was significantly higher than that of the PCL (0.002±0.0004 S / m) nanofiber scaffold and significantly higher than that of biological tissue (0.3~0.7 S / m). The enhanced conductivity promoted the transmission of electrical signals in bioelectrically active tissues and accelerated the nerve healing process. We then used a three-electrode electrochemical apparatus to test the impedance properties of the PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. Figure 1 As shown in h, compared to PCL nanofiber scaffolds, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds have lower interfacial charge transfer impedance and faster electron transport rate. Figure 1 The Bode plots also showed that the electrical impedance (≈40 kΩ) of the PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds was significantly lower than that of PCL nanofibers (≈70 kΩ). This low charge transfer impedance characteristic is beneficial for enhancing intercellular communication and promoting nerve healing by regulating the endogenous bioelectrical signals generated by ion channels and ion pumps. Furthermore, we evaluated the redox properties and charge storage capacity of the PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds using cyclic voltammetry (CV) curves. Figure 1 The CV curves of all samples exhibited a near-parallelogram shape with no obvious redox peaks, indicating that the materials possess high electrochemical stability and have not undergone oxidation. Quantitative analysis showed that the charge capacities of the PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds were 3.3 × 10⁻⁵ C, 3.56 × 10⁻⁵ C, and 3.91 × 10⁻⁵ C, respectively, significantly higher than those of PCL nanofibers (1.86 × 10⁻⁵ C). High charge capacity maintains the stability and continuity of charge injection at the cell-material interface, thereby effectively enhancing cellular electrochemical signal transduction. Effective integration of electrochemical signals with tissues depends on charge injection and storage capacity. Figure 1The results showed that, compared with PCL, the charge injection capacity of PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds was significantly improved, which facilitates the efficient transfer of electrical signals from electronic conduction circuits to ion conduction tissues. Furthermore, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofibers could all light up LEDs. Simultaneously, when PCL / PPP nanofibers were rolled into NGC, they maintained their conductivity and were able to power LED bulbs. Figure 1 l).

[0058] It is noteworthy that the mechanical and electrochemical properties of the PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds showed no significant differences, indicating the stability of their physicochemical properties after the formation of the bicontinuous phase structure. This will help to maximize the conditions for cell growth and development while regulating their biocompatibility without causing toxic effects.

[0059] In this invention, to verify the applicability of the composite neural conduit with both structural guidance and conductivity properties for the adhesion and proliferation of SCs, we seeded SCs onto the surface of PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. Figure 2 As shown in b, PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds all supported the continuous proliferation of RSC96. However, after 3 days of culture, compared with PCL and PCL / 3%PPP nanofiber scaffolds, the number of cells on the surface of PCL / 5%PPP and PCL / 7%PPP nanofiber scaffolds decreased, indicating that the introduction of excessive PPP may produce cytotoxicity. PCL / 3%PPP, on the other hand, showed good cell compatibility and effectively promoted RSC96 proliferation and adhesion. This may be due to the fact that the presence of conductive components can enhance cell signal transduction and promote the adsorption of extracellular matrix proteins (such as fibronectin and laminin) by regulating the integrin β1 signaling pathway.

[0060] Myelinated neuronal stem cells (SCs) can regulate neuronal survival, axonal development, and synaptic functionalization in the peripheral nervous system by secreting various neurotrophic factors. To evaluate the enhancing effect of conductive nanofiber scaffolds on myelination of SCs, we seeded SCs on PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofibers, and monitored the morphological changes of SCs on the scaffolds at 1, 3, and 5 days of culture. Figure 2Phalloidin staining revealed that, with prolonged culture time, the SCs on the PCL / PPP nanofiber scaffold exhibited a distinctly mature bipolar morphology compared to PCL. Pseudopodia length statistics showed that after 5 days of culture, the pseudopodia length of the PCL nanofiber scaffold was 29.92±24.89 μm, which may be due to the topological guidance effect of the neatly arranged electrospun nanofibers on the cells. Under the synergistic effect of topological guidance and electroactivity, the cell extension behavior induced by PCL / 3%PPP conductive nanofibers (pseudopodia length 54.95±17.61 μm) was the most pronounced. Figure 2 c), which highly matches the characteristics of myelinated SCs reported in the literature. To investigate the regulatory role of conductive nanofiber scaffolds on the expression of genes related to myelination in SCs, we seeded RSC96 onto the surfaces of PCL, PCL / 3%PPP, PCL / 5%PPP, PCL / 7%PPP nanofiber scaffolds, and tissue culture polystyrene (TCP, control group). After 5 days of culture, the expression levels of marker genes, including peripheral myelin 22 (PMP22), early growth response protein 2 (Krox20 / EGR2), and nerve growth factor (NGF), which indicate myelination of SCs, were detected by real-time quantitative PCR. Figure 2 As shown in Figure d, compared to the TCP and PCL nanofiber scaffold groups, the PCL / 3%PPP conductive nanofiber group significantly upregulated the expression of PMP22, Krox20, and NGF genes. However, due to the introduction of excessive PPP, the expression of myelination-related genes in SCs of the PCL / 5%PPP and PCL / 7%PPP nanofiber scaffold groups did not show a significant increase and remained comparable to the PCL group. These results indicate that conductive materials can mimic the natural electrical microenvironment of neural tissue and significantly enhance the myelination capacity of SCs.

[0061] Although research on the molecular mechanisms by which conductive materials regulate cell behavior is currently limited, evidence suggests that changes in intracellular Ca²⁺ concentration can significantly affect the survival, proliferation, and differentiation of cellular spinal cord (SCs) and neurons. To investigate the effects of four nanofiber scaffolds on intracellular Ca²⁺ concentration in SCs, we used the calcium fluorescent probe Fluo-3AM to detect changes in intracellular Ca²⁺ concentration on the surfaces of PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. The results showed that, compared to the high Ca²⁺ levels of the TCP and PCL nanofiber scaffolds, the Ca²⁺ concentration in the PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP conductive nanofiber scaffolds (RSC96) remained low throughout the culture process, suggesting that the conductive interface may promote myelination of SCs by regulating Ca²⁺ homeostasis. Related literature also reports that conductive materials reduce intracellular Ca²⁺ concentration by inhibiting the CaSR-PLC pathway, thereby relieving the transcriptional inhibition of myelination genes by PKC and promoting myelination of SCs and continuous secretion of neurotrophic factors.

[0062] Dorsal root ganglion neurons (DRGs), as a classic in vitro model of primary sensory neurons in the peripheral nervous system, have become a key tool for studying axonal regeneration mechanisms due to their unique biological characteristics and regenerative potential. To investigate the effects of four nanofiber scaffolds on DRG axonal growth, we seeded DRGs onto the surface of PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP nanofiber scaffolds. After culturing for 5 days, the DRGs were stained with Tuji and DAPI. Figure 3 a). The results showed that due to the guiding effect of the aligned nanofibers, neurites in all groups grew directionally along the fiber alignment. Quantitative analysis of neurite length showed that the average neurite lengths of the PCL, PCL / 3%PPP, PCL / 5%PPP, and PCL / 7%PPP groups were 651.74±174.01 nm, 1430.80±327.67 nm, 851.33±186.08 nm, and 1101.62±248.60 nm, respectively, while the longest neurite lengths reached 979.91±28.27 nm, 1935.13±26.81 nm, 1236.08±199.21 nm, and 1511.56±7.45 nm, respectively. Figure 3 b and Figure 3 c). Experimental results show that the PCL / 3%PPP conductive nanofiber scaffold has the most significant effect on promoting DRG neurite growth, with an average neurite length 2.2 times that of the PCL nanofiber scaffold and the longest neurite length 1.97 times that of the PCL scaffold.

[0063] In this invention, to assess the promoting effect of composite neural conduits on neurite spur growth at the gene level, we selected PC12 cells as a neural differentiation model. Numerous studies have used this cell model to reveal the regulatory mechanisms of material topology and electroactivity on synaptic extension. Observing the effect of nanofiber scaffolds on PC12 cell differentiation in vitro can effectively predict their potential efficacy in neural regeneration. Based on the good biocompatibility and significant DRG-promoting axon extension effect of the PCL / 3%PPP conductive nanofiber scaffold, we selected PCL / 3%PP to study the effect of conductive nanofiber scaffolds on the expression of genes related to neurite extension and axon growth in PC12 cells. We seeded PC12 cells onto the surface of TCP, PCL, and PCL / 3%PPP nanofiber scaffolds, and performed Tuji and DAPI staining after 5 days of culture. Figure 3 As shown in Figure d, after 5 days of culture, the neurites of differentiated PC12 cells extended directionally along PCL and PCL / 3%PPP nanofibers, while exhibiting random distribution on the TCP surface. Figure 3 e and Figure 3As shown in figure f, the neurite length in the PCL / 3%PPP group was significantly increased compared to the TCP and PCL groups, with the average and longest neurite lengths being 1.65 and 1.48 times that of the PCL group, respectively. Simultaneously, the expression levels of axonal growth-related genes (GAP-43, NF-H, NF-L, Synapsin I, PSD95) were detected using real-time quantitative PCR. Figure 3 As shown in g, compared to the TCP and PCL groups, the expression levels of GAP-43, NF-H, NF-L, Synapsin I, and PSD95 genes were significantly upregulated in the PCL / 3%PPP group. These results indicate that the electroactivity of the scaffold can synergistically promote the differentiation of PC12 cells into neuron-like cells and achieve directional extension of neurites. These results confirm that the material's topology and electroactivity signals can synergistically promote neurite extension, cell differentiation, and adhesion, providing an ideal platform for peripheral nerve injury repair and creating more options and possibilities for neural tissue engineering.

[0064] Flexible electrodes play a crucial role in achieving high-quality electrophysiological recording and effective electrical stimulation. Existing implantable electrodes primarily rely on conductive materials with high Young's modulus, such as platinum (Pt), silicon, and stainless steel. Mechanical differences at the rigid material-tissue interface can lead to increased immune responses and chronic tissue inflammation. Therefore, scar tissue growing on conductive surfaces degrades the quality of electrical signal recording (e.g., signal-to-noise ratio) and stimulation efficacy. In electrochemical performance testing, bicontinuous conductive nanofiber scaffolds exhibited low interfacial impedance, high charge injection capacity, and high charge storage capacity—parameters crucial for effective electrical stimulation and neuromodulation. To demonstrate the potential of bicontinuous conductive nanofiber scaffolds as flexible electrodes, we compared their performance in neurostimulation by performing in vivo neural stimulation and compound muscle action potential (CMAP) recording with clinically used stainless steel acupuncture hooks. Figure 4 a) First, we performed electrical stimulation on the sciatic nerve of Lewis rats and recorded CMAP and ankle rotation angles. (e.g.) Figure 4 As shown in b, PCL / 3%PPP was wrapped around the sciatic nerve of Lewis rats, with a stainless steel acupuncture hook as a control. Results showed that under the same electrical stimulation conditions (frequency 3 Hz, current 2 mA), PCL / 3%PPP and the stainless steel acupuncture hook exhibited similar changes in ankle rotation angle. Figure 4 c and Figure 4 d) and CMAP Figure 4 e.g., we also conducted large animal experiments, wrapping PCL / 3%PPP and stainless steel acupuncture hooks around the median nerve of rhesus monkeys for electrical stimulation (5 mA) and CMAP recording. Figure 4h). The results also showed that PCL / 3%PPP could record electrophysiological signals comparable to those of clinically used conductive needles, and CMAP recordings were not significantly different from those of stainless steel acupuncture hook needles (h). Figure 4 These results demonstrate that PCL / 3%PPP can be used for effective stimulation and high-quality electrophysiological signal recording, and has the potential as a bioelectronic implant. Furthermore, such as Figure 4 As shown in b, we also found that compared to rigid stainless steel acupuncture hooks which create nanoindentations on nerves, flexible PCL / 3%PPP fiber membranes do not cause nerve damage.

[0065] Eight weeks after surgery, a rat model of 10 mm sciatic nerve defect showed varying degrees of improvement in regenerated nerve function. Footprint analysis reflected improvements in hindlimb motor ability. Figure 5 a) The sciatic nerve function index (SFI) showed no significant difference among the groups at 4 weeks postoperatively; however, at 8 weeks postoperatively, the PCL / 3%PPP group (-77.99±4.75) was significantly better than the PCL group (-84.01±4.07), while there was no significant difference between the PCL / 3%PPP group and the Autograft group (-70.56±2.74). Figure 5 b). Withdrawal threshold is used to assess sensory recovery. Figure 5 c). Although there were no significant differences between groups at 4 and 8 weeks, the withdrawal thresholds of PCL / 3%PPP were closer to those of Autograft. Compound muscle action potential (CMAP), as an important electrophysiological indicator of neuromuscular function, is one of the key indicators for assessing sciatic nerve function recovery. Quantitative analysis ( Figure 5 e) The CMAP amplitude of the PCL / 3%PPP group (7.40±1.49 mV) reached 67.48% of that of the Autograft group (10.97±1.37 mV), approximately 2.22 times that of the PCL group (3.33±0.68 mV), suggesting that PCL / 3%PPP has a significant advantage in functional nerve reinnervation recovery compared to PCL. At 8 weeks, H&E staining of longitudinal sections of regenerated nerves showed that nerve connections were complete in all groups. Figure 5 f). H&E cross-sectional observation of the regenerated segment and distal portion revealed that the PCL / 3%PPP regenerated nerves were denser than those in the PCL, more closely resembling Autograft (f). Figure 5 g).

[0066] In conclusion, the PCL / 3%PPP nerve conduit demonstrates promising therapeutic potential in promoting functional recovery after peripheral nerve transection. Its superior repair effect may be attributed to the conductivity and bioactivity imparted to the material by the PPP component, thereby enhancing functional recovery following peripheral nerve regeneration.

[0067] To assess the regenerative capacity of the nerve conduit, we performed immunofluorescence staining analysis on longitudinal and transverse sections of the regenerated nerve segment and distal nerve segment at 8 weeks postoperatively (Figure 6). Longitudinal immunofluorescence staining showed the presence of blood vessels (CD31 positive) in all groups. The distribution of growth cones and axons (GAP43 positive) in the PCL / 3% PPP group was comparable to that in the autologous transplantation group, while the fluorescent area in the PCL group was significantly smaller. Figure 6 a) Schwann cells (SCs) and nerve fibers extend continuously throughout the regenerated nerve (Fig. 6b). The extent of nerve regeneration and tissue remodeling was assessed by quantitative analysis of key markers of angiogenesis, axonal growth, and Schwann cell activity. Newly formed blood vessels deliver oxygen and nutrients to the injury site, thus supporting nerve regeneration. Angiogenesis was assessed by detecting the endothelial cell marker CD31 (Fig. 6d). Data showed that CD31 positive expression in the PCL / 3% PPP group reached 74.72% and 76.95% in the regenerated and distal segments, respectively, which were 1.29 and 1.46 times higher than in the PCL group (Fig. 6h).

[0068] Growth cones are dynamic structures located at the ends of neuronal axons that guide axonal extension toward their targets. Axonal regeneration can be assessed by detecting the growth cone biomarker GAP43. Figure 6 c). The results showed that in the regenerated segment, the GAP43 expression level in the PCL / 3% PPP group reached 71.99% of that in the autologous transplantation group, which was 2.15 times that of the PCL group; in the distal segment, the GAP43 expression level in the PCL / 3% PPP group reached 65.20% of that in the autologous transplantation group, which was 2.88 times that of the PCL group. Figure 6 These results indicate that PCL / 3% PPP is more effective than PCL in promoting axonal and angiogenesis. Schwann cells play an important role in nerve regeneration and are a key indicator for assessing nerve repair. S100 is the most characteristic marker of mature Schwann cells (g). Figure 6 e). In the PCL / 3% PPP group, S100 expression in the regenerated segment reached 63.52% of the autologous transplantation group, and in the distal segment, it reached 35.73%, which were 2.90 times and 2.02 times higher than those in the PCL group, respectively. Figure 6 i). The formation of mature axons is assessed by detecting the neurofilament marker NF200. In the regenerating segment ( Figure 6 e), NF200 expression in the PCL / 3% PPP group reached 55.99% of that in the autologous transplantation group, which was 4.51 times that of the PCL group; in the distal segment ( Figure 6f), its expression reached 33.60% in the autologous transplantation group, which was 2.83 times that of the PCL group ( Figure 6 These results indicate that PCL / 3% PPP catheters effectively promote Schwann cell proliferation and migration, as well as nerve fiber formation.

[0069] The PCL / 3% PPP catheter exhibits good biocompatibility and conductivity, and promotes the formation of axons, blood vessels, Schwann cells, and mature nerve fibers more effectively than the PCL catheter, with performance approaching that of the autologous transplant group. These results indicate that the PCL / 3% PPP catheter holds great potential in peripheral nerve repair.

[0070] To assess myelin regeneration and muscle atrophy reversal after nerve repair, we stained the regenerated nerve tissue with Luxol Fast Blue (LFB) at 8 weeks post-surgery (Fig. 7a). In the PCL / 3% PPP group, the myelin density in the mid-segment and distal segments reached 68.96% and 83.55% of that in the autologous transplantation group, respectively, which were 1.68 and 1.53 times that of the PCL group (Fig. 7b). The myelin thickness in the mid-segment and distal segments of this group reached 94.96% and 86.14% of that in the autologous transplantation group, respectively, which were 1.01 and 1.30 times that of the PCL group (Fig. 7c). Regarding axonal diameter, the mid-segment and distal segments of the PCL / 3% PPP group reached 98.00% and 95.92% of that in the autologous transplantation group, respectively, which were 1.09 and 1.18 times that of the PCL group (Fig. 7d). These results indicate that this material has a positive effect on promoting myelin formation. To assess the impact of neurological function recovery on target muscle preservation, we analyzed gastrocnemius muscle weight and Masson's trichrome staining (Fig. 7e). The wet weight of the gastrocnemius muscle in the PCL / 3% PPP group reached 47.94% of that in the autologous transplant group, which was 1.23 times that of the PCL group (Fig. 7f); the cross-sectional area of ​​muscle fibers reached 63.55% of that in the autologous transplant group, which was 1.60 times that of the PCL group (Fig. 7g). Although these two indicators were still significantly different from those in the autologous transplant group, they were significantly better than those in the PCL group. These results indicate that this material has a positive effect on myelination and muscle preservation.

[0071] PCL / 3% PPP catheters can not only promote myelin regeneration and axonal maturation of regenerating nerves, but also indirectly reduce target muscle atrophy by enhancing nerve function recovery.

[0072] The above description is merely a preferred embodiment of the present invention and is not intended to limit the invention. Various modifications and variations can be made to the invention by those skilled in the art. Any modifications, equivalent substitutions, improvements, etc., made within the spirit and principles of the invention should be included within the scope of protection of the invention. It should be noted that similar reference numerals and letters in the following figures denote similar items; therefore, once an item is defined in one figure, it does not need to be further defined and explained in subsequent figures.

Claims

1. A method for preparing a conductive core-shell nanofiber membrane that combines the applications of flexible electrodes and nerve conduits, characterized in that, Includes the following steps: S100. Stir the poly(3,4-ethylenedioxythiophene)-poly(styrene sulfonate) solution at room temperature, filter to obtain filtrate, freeze the filtrate with liquid nitrogen and place it in a freeze dryer to obtain PEDOT:PSS nanofibers. S200. PEDOT:PSS nanofibers and hydrophilic PU are dissolved in an ethanol solution at a weight ratio of 2-4:7 to prepare a bicontinuous phase conductive polymer solution with a mass concentration of 0.5-10%. S300. A conductive core-shell nanofiber membrane is prepared by coaxial electrospinning using a bicontinuous phase conductive polymer solution as the shell solution and a polymer solution as the core solution, wherein the polymer solution is one or more of PCL solution, P34HB solution, and PLGA solution.

2. The method for preparing a conductive core-shell nanofiber membrane for flexible electrode and nerve conduit applications according to claim 1, characterized in that, In step S100: The stirring time at room temperature was 6 hours; the filtration accuracy was 0.45 μm; the liquid nitrogen freezing time was 15 minutes; and the freeze-drying time was 72 hours.

3. The method for preparing a conductive core-shell nanofiber membrane with both flexible electrode and nerve conduit applications according to claim 1, characterized in that, In step S200: PEDOT:PSS nanofibers and hydrophilic PU are mixed in a weight ratio of 3:7; the mass concentration of the bicontinuous phase conductive polymer solution is 3%.

4. The method for preparing a conductive core-shell nanofiber membrane with both flexible electrode and nerve conduit applications according to claim 1, characterized in that, In step S300: The mass concentration of the polymer solution is 12%.

5. The method for preparing a conductive core-shell nanofiber membrane with both flexible electrode and nerve conduit applications according to claim 4, characterized in that, In step S300: The conductive core-shell nanofiber membrane has a bilayer polymer fiber layer with an inner unidirectional ordered layer and an outer layer with random arrangement.

6. A conductive core-shell nanofiber membrane that combines the applications of flexible electrodes and nerve conduits, characterized in that, Prepared by any one of the methods of claims 1-5.

7. A conductive core-shell nanofiber membrane that combines the applications of flexible electrodes and nerve conduits, obtained by any one of the methods of claims 1-5, or the application of the conductive core-shell nanofiber membrane that combines structural guidance and conductivity as described in claim 7 in flexible nerve electrodes and nerve conduits in the treatment of peripheral nerve injuries.

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