Implantable artificial nerve and preparation method thereof

Abstract

The invention provides an implantable artificial nerve and a preparation method thereof, and belongs to the field of biomedical engineering and neural restoration. Comprising a signal coupling interface, the signal coupling interface is used for being connected with broken nerves, the signal coupling interface is further used for forming interface impedance lower than 500 omega.cm < 2 > under the condition that conversion impedance between biological ion current and artificial electron current is 1 kHz, and the signal coupling interface is further filled with neurotrophic factors; the conduction unit is connected with the signal coupling interface, and the conduction unit is filled with neurotrophic factors; the bionic microvessel channel is arranged in the conduction unit, and the bionic microvessel channel is used for conveying a nutrient solution or guiding revascularization. The method aims at solving the problems that an existing neural restoration technology is low in signal conversion efficiency, passive in nerve regeneration guiding mode and the like. The device has the technical effects that efficient signal conversion is realized, and a microenvironment beneficial to regeneration can be actively created.

Description

Technical Field

[0001] This invention relates to the fields of biomedical engineering and nerve repair technology, and particularly to an implantable artificial nerve, as well as a method for preparing an implantable artificial nerve. Background Technology

[0002] Nerve damage caused by accidental trauma, iatrogenic injury, or disease can lead to sensory loss, motor dysfunction, and even muscle atrophy. If not repaired promptly and effectively, it often results in permanent disability, severely impacting a patient's daily life and work ability. Re-establishing nerve signal transmission pathways through neural connection technology is of great significance in initiating bodily function recovery. Achieving neural connections helps prevent muscle atrophy, tendon disuse, and joint stiffness, maintaining muscle volume, strength, and range of motion, thereby shortening rehabilitation time and improving recovery quality. Re-established sensory pathways can also improve nerve pain and abnormal sensations, enhancing quality of life and psychological well-being.

[0003] To promote the repair and regeneration of nerves surrounding severed nerves, the medical community has explored various microsurgical techniques and repair materials. Gallium-based room-temperature liquid metals, due to their high viscosity (3.1 × 10⁻⁶), are among the most promising options. 6 The ultra-high conductivity of S / m can reduce ohmic losses in the transmission of weak nerve signals. It behaves as a flowable liquid at room temperature and exhibits excellent compliance in vivo. Gallium-based liquid metals have been used to connect severed sciatic nerves, capable of transmitting electrical stimulation signals. Existing technologies include schemes that combine the conductive pathway of the liquid metal with a neurotrophic factor-filled chamber through concentric cannulas or multi-channel conduit structures, aiming to maintain signal transmission function during nerve regeneration. However, after in-depth research, these schemes still have certain limitations: (i) The regeneration guidance mode is passive and lacks a physical microenvironment. Existing technologies mainly rely on filling neurotrophic factors to "passively wait" for nerve growth. The internal space of the duct is either filled with a scaffold or is a simple cavity, lacking a physical microstructure that can actively guide the directional growth of axons. At the same time, the release of nutrients is difficult to control and cannot be well matched with the long nerve regeneration cycle.

[0004] (ii) Poor interfacial compatibility and low signal conversion efficiency. Although conductive nanomaterials have been introduced as a transition, the physical-chemical barrier between metallic electronic conductors and biological neural ionic conductors has not been fundamentally resolved. The interfacial impedance remains high, causing attenuation and distortion of weak neural action potentials during conduction.

[0005] In summary, existing neural repair technologies suffer from problems such as low signal conversion efficiency and a relatively passive guidance mode for neural regeneration. Summary of the Invention

[0006] This invention provides an implantable artificial nerve and its preparation method, which addresses the shortcomings of existing nerve repair technologies, such as low signal conversion efficiency and passive guidance mode for nerve regeneration. The invention achieves efficient signal conversion and can actively create a favorable microenvironment for regeneration in the implantable artificial nerve.

[0007] The first aspect of this invention provides an implantable artificial nerve, comprising: The signal coupling interface is used to connect to the severed nerve and also to form a conversion impedance between the biological ion flow and the artificial electron flow. The interface impedance at a frequency of 1 kHz is less than 500 Ω·cm. 2 The signal coupling interface is also filled with neurotrophic factors. The conduction unit is connected to the signal coupling interface and is filled with neurotrophic factors. The biomimetic microvascular channel is set within the conduction unit and is used to deliver nutrient solution or guide blood vessel regeneration.

[0008] In addition, the implantable artificial nerve according to the present invention may also have the following additional technical features: In some embodiments of the present invention, the conduction unit includes: The outer encapsulation sheath is connected to the signal coupling interface, and a biomimetic microvascular channel is provided inside the outer encapsulation sheath. The biomimetic parallel signal conduction bundle has multiple bundles, each connected to a signal coupling interface, and is spaced apart within an outer encapsulation sheath.

[0009] In some embodiments of the present invention, the conduction unit further includes: There are multiple biomimetic nerve peritunes, each corresponding to a different biomimetic parallel signal conduction bundle.

[0010] In some embodiments of the present invention, the biomimetic parallel signal transmission bundle includes: The conductive unit is a multiple conductive unit, each of which is connected to a signal coupling interface. The multiple conductive units are spaced apart within the biomimetic layer of the nerve peritunic. There are multiple longitudinal regeneration channels, all of which are located within the conductive units. Each longitudinal regeneration channel is situated between two conductive units and is filled with type I collagen or GelMA hydrogel.

[0011] In some embodiments of the present invention, the diameter of the biomimetic microvascular channel is between 50 μm and 200 μm.

[0012] In some embodiments of the present invention, the conductivity of the conductive unit is greater than 10. 5 S / m.

[0013] In some embodiments of the present invention, the outer encapsulation sheath is a biocompatible flexible material or a biodegradable material.

[0014] In some embodiments of the present invention, each conductive unit includes: A conductive layer is disposed within the biomimetic layer of the nerve peritunic and is used to connect with the signal coupling interface. An insulating layer is placed inside the biomimetic layer of the nerve peritunic and covers the conductive layer.

[0015] In some embodiments of the present invention, the signal coupling interface is a conductive hydrogel interface or a liquid metal nanocomposite gel interface.

[0016] A second aspect of the present invention provides a method for preparing an implantable artificial nerve, used to prepare the implantable artificial nerve of the first aspect of the present invention, comprising the following steps: Step S100: Prepare and integrate neurotrophic factors, conduction units, and biomimetic microvascular channels; Step S200: Construct a signal coupling interface and connect the conduction unit to the signal coupling interface; Step S300: Preparation complete.

[0017] In summary, this application has the following beneficial technical effects: by setting the signal coupling interface, efficient and low-loss conversion between ion flow and electron flow is achieved, ensuring the complete transmission of weak neural signals.

[0018] By combining a signal coupling interface with a conduction unit, this artificial nerve can immediately establish an electrophysiological connection after implantation, preventing muscle atrophy. Simultaneously, the biomimetic microvascular channel allows for postoperative infusion of nutrient solution or guidance of host blood vessel growth. The neurotrophic factor setting actively guides Schwann cell migration and axonal directional growth, providing time-controlled nutritional support, transforming passive waiting into active guidance, significantly accelerating the regeneration process and improving the quality of nerve repair. Attached Figure Description

[0019] Various other advantages and benefits will become apparent to those skilled in the art upon reading the following detailed description of preferred embodiments. The accompanying drawings are for illustrative purposes only and are not intended to limit the invention. Furthermore, the same reference numerals denote the same parts throughout the drawings. In the drawings: Figure 1 A schematic diagram of the structure of a conduction unit of an implantable artificial nerve according to some embodiments of the present invention is shown.

[0020] Figure 2 A schematic diagram of the structure of an implantable artificial nerve with a biomimetic parallel signal conduction bundle combined with a biomimetic layer of the nerve peritunic, according to some embodiments of the present invention, is shown.

[0021] Figure 3 A schematic diagram of the structure of a conductive unit of an implantable artificial nerve according to some embodiments of the present invention is shown.

[0022] Figure 4 A schematic diagram of the structure of an implantable artificial nerve according to some embodiments of the present invention is shown.

[0023] Figure 5 A schematic diagram of an implantable artificial nerve without an outer encapsulation sheath according to some embodiments of the present invention is shown.

[0024] Figure label: 1. Outer encapsulation sheath; 2. Bionic parallel signal conduction bundle; 3. Bionic microvascular channel; 4. Bionic layer of nerve peritunic; 5. Neurotrophic factor; 6. Conductive unit; 7. Signal coupling interface; 8. Severed nerve. Detailed Implementation

[0025] Nerve damage caused by accidental trauma, iatrogenic injury, or disease can lead to sensory loss, motor dysfunction, and even muscle atrophy. If not repaired promptly and effectively, it often results in permanent disability, severely impacting a patient's daily life and work ability. Re-establishing nerve signal transmission pathways through neural connection technology is of great significance in initiating bodily function recovery. Achieving neural connections helps prevent muscle atrophy, tendon disuse, and joint stiffness, maintaining muscle volume, strength, and range of motion, thereby shortening rehabilitation time and improving recovery quality. Re-established sensory pathways can also improve nerve pain and abnormal sensations, enhancing quality of life and psychological well-being.

[0026] To promote the repair and regeneration of severed peripheral nerves, the medical community has explored various microsurgical techniques and repair materials. In the field of peripheral nerve defect repair, the clinically commonly used nerve conduit technique has shown unique advantages in repairing defects shorter than 2 cm. For example, the guiding structure constructed from biomaterials can effectively reduce tissue adhesion and provide a carrier for the local slow release of neurotrophic factors. However, its application is limited by strict length indications and a long axonal regeneration cycle. For more complex nerve defects, autologous nerve transplantation, as the gold standard, can maintain a good fiber bundle arrangement structure, but it inevitably faces inherent drawbacks such as donor site function loss and limited graft sources. Allogeneic nerve transplantation can overcome length limitations, but the accompanying immunosuppressive therapy brings risks of infection and tumor development. Even the most basic nerve end-to-end anastomosis is strictly limited in its application by the defect length and the condition of the severed ends. Although the adjuvant use of bioactive substances such as growth factors can promote regeneration to some extent, issues such as the route of administration, dosage control, and long-term safety remain to be resolved. The limitations of these technologies, including specific restrictions on repair length and nerve type, long functional recovery cycles, and the risk of irreversible target organ atrophy, together constitute major clinical challenges in the current field of nerve repair and highlight the urgent need for novel nerve repair technologies.

[0027] Gallium-based room-temperature liquid metals have a density of 3.1 × 10⁻⁶. 6 The ultra-high conductivity of S / m reduces ohmic losses in the transmission of weak nerve signals. It behaves as a flowable liquid at room temperature and exhibits excellent compliance in vivo. Gallium-based liquid metals have been used to reconnect severed sciatic nerves, capable of transmitting electrical stimulation signals and demonstrating a certain degree of consistency in macroscopic signal transmission.

[0028] Chinese patent publication number CN103405289B discloses a device based on liquid metal for the repair of damaged peripheral nerve function. It combines a liquid metal conductive pathway with a neurotrophic factor-filled chamber via a concentric cannula or multi-channel conduit structure, aiming to maintain signal transmission function during nerve regeneration. This technology represents a significant advancement in the field.

[0029] However, after in-depth research, the existing technology still has the following limitations that urgently need to be addressed: First, the structural biomimicry is insufficient, resulting in low signal fidelity. Existing solutions mostly employ macroscopic functional partitioning structures, such as inner tubes filled with nutrients and outer tubes providing conductivity. This structure short-circuits the ends of all nerve fibers to one or a few macroscopic conductive channels, failing to simulate the microscopic topology of thousands of parallel, insulated axons in natural nerves. This leads to the mixing of independent electrical signals from different functional nerve fibers, such as motor and sensory fibers, during conduction, resulting in poor encoding fidelity of neural information and an inability to provide precise functional instructions to distant target organs.

[0030] Second, the regeneration guidance mode is passive and lacks a physical microenvironment. Existing technologies mainly rely on filling neurotrophic factors to "passively wait" for nerve growth. The internal space of the duct is either filled with a scaffold or is a simple cavity, lacking physical microstructures that can actively guide the directional growth of axons. At the same time, the release of nutrients is difficult to control and cannot be well matched with the long nerve regeneration cycle.

[0031] Third, poor interfacial integration and low signal conversion efficiency. Although conductive nanomaterials have been introduced as a transition, the physicochemical barrier between metallic electronic conductors and biological neural ionic conductors remains fundamentally unresolved. The interfacial impedance remains high, causing attenuation and distortion of weak neural action potentials during conduction. Therefore, there is an urgent need in this field for a novel neural repair device that can simulate natural neural anatomy at the microscale, achieve efficient signal conversion, and actively create a favorable regenerative microenvironment.

[0032] Exemplary embodiments of the present disclosure will now be described in more detail with reference to the accompanying drawings. While exemplary embodiments of the present disclosure are shown in the drawings, it should be understood that the present disclosure may be implemented in various forms and should not be limited to the embodiments set forth herein. Rather, these embodiments are provided so that this disclosure will be thorough and complete, and will fully convey the scope of the disclosure to those skilled in the art.

[0033] It should be understood that the terminology used herein is for the purpose of describing particular exemplary embodiments only and is not intended to be limiting. Unless the context clearly indicates otherwise, the singular forms “a,” “an,” and “” used herein may also indicate the inclusion of the plural forms. The terms “comprising,” “including,” “containing,” and “having” are inclusive and therefore indicate the presence of the stated features, steps, operations, elements, and / or components, but do not exclude the presence or addition of one or more other features, steps, operations, elements, components, and / or combinations thereof. The method steps, processes, and operations described herein are not construed as requiring them to be performed in a particular order described or illustrated, unless the order of performance is explicitly indicated. It should also be understood that additional or alternative steps may be used.

[0034] Although terms such as first, second, third, etc., may be used in this document to describe multiple elements, components, regions, layers, and / or segments, these elements, components, regions, layers, and / or segments should not be limited by these terms. These terms may be used only to distinguish one element, component, region, layer, or segment from another. Unless the context clearly indicates otherwise, terms such as "first," "second," and other numerical terms used herein do not imply order or sequence. Therefore, the first element, component, region, layer, or segment discussed below may be referred to as the second element, component, region, layer, or segment without departing from the teachings of the exemplary embodiments.

[0035] For ease of description, spatial relative terms may be used in the text to describe the relationship of one element or feature relative to another element or feature, as shown in the figure. These relative terms include, for example, "inside," "outside," "middle," "outer," "below," "below," "above," "over," etc. Such spatial relative terms are intended to include different orientations of the device in use or operation, other than those depicted in the figure. For example, if the device in the figure is flipped, an element described as "below other elements or features" or "below other elements or features" would subsequently be oriented "above other elements or features" or "above other elements or features." Therefore, the example term "below" can include both upper and lower orientations. The device may also be rotated 90 degrees or in other orientations, and the spatial relative descriptors used in the text will be interpreted accordingly.

[0036] like Figures 1 to 5 As shown, according to an embodiment of the first aspect of the present invention, an implantable artificial nerve is proposed, comprising a signal coupling interface 7, a conduction unit, and a biomimetic microvascular channel 3. The signal coupling interface 7 is used to connect to a severed nerve 8, and the conduction unit is connected to the signal coupling interface 7. The signal coupling interface 7 is used to form a conversion impedance between a biological ion flow and an artificial electron flow, with an interface impedance of less than 500 Ω·cm at a frequency of 1 kHz. 2 The biomimetic microvascular channel 3 is set in the conduction unit. The biomimetic microvascular channel 3 is used to deliver nutrient solution or guide blood vessel regeneration. The conduction unit and the signal coupling interface 7 are also filled with neurotrophic factors 5.

[0037] In the above embodiments, it should be noted that the neurotrophic factor 5 is composed of PLGA biodegradable microspheres loaded with neurotrophic factor 5 and / or small molecule nutrients melt-blended with the insulating shell material; the PLGA microspheres are prepared by emulsion solvent evaporation method and have a diameter between 1 μm and 10 μm.

[0038] The signal coupling interface 7 is a dynamic ion-electron signal coupling interface 7: both ends of the conduction unit are provided with signal coupling interfaces 7. The signal coupling interface 7 is used to efficiently connect the severed nerve 8 and realize low impedance conversion between biological ion flow and artificial electron flow. It can selectively adopt one or more of the following material forms: First, an anisotropic conductive hydrogel interface is prepared as follows: A prepolymer solution containing acrylamide monomer, dopamine hydrochloride, conductive nanomaterials, and a photoinitiator is first prepared. The prepolymer solution is injected into a parallel electrode mold, and a DC electric field is applied to induce the directional alignment of the nanomaterials, causing the conductive nanomaterials dispersed in the prepolymer solution to align. Subsequently, ultraviolet light is used to initiate polymerization and fix the structure, forming an interface with directional conductivity and wet adhesion properties. Its structural characteristic is that the conductive filler is arranged along a direction perpendicular to the long axis of the nerve, forming a preferential conductive pathway. The conductive nanomaterials can be materials such as carbon nanotubes or graphene, and the strength of the DC electric field is between 50 V / cm and 200 V / cm.

[0039] Second, the liquid metal nanocomposite gel interface is prepared by: preparing liquid metal into nanodroplets by ultrasonic fragmentation, and then uniformly dispersing the nanodroplets into the above-mentioned hydrogel prepolymer liquid, and forming a three-dimensional dual-mode conduction network by in-situ thermal polymerization or photopolymerization; wherein, the diameter of the nanodroplets is less than 100 nm.

[0040] Third, the multi-layer gradient structure interface is prepared by using a layer-by-layer self-assembly technique or a sequential casting-polymerization process to sequentially construct a high-water-content ion-conducting layer, a conductive nanomaterial-reinforced transition layer, and an electronically conductive layer. The high-water-content ion-conducting layer can be made of alginate gel for contact with nerves. The conductive nanomaterial-reinforced transition layer can be made of PEDOT:PSS or sodium alginate composite. The electronically conductive layer can be made of high-content liquid metal composite gel for contact with the conductive unit.

[0041] By setting the signal coupling interface 7, especially the anisotropic conductive hydrogel prepared by electric field-induced assembly, a directional hybrid conductive network is formed inside, realizing efficient and low-loss conversion between ion flow and electron flow, and ensuring the complete transmission of weak nerve signals.

[0042] The technical effect achieved by the above embodiments is that, through the setting of the signal coupling interface 7, efficient and low-loss conversion between ion flow and electron flow is realized, ensuring the complete transmission of weak neural signals.

[0043] By combining the signal coupling interface 7 with the conduction unit, this artificial nerve can immediately establish an electrophysiological connection after implantation, preventing muscle atrophy. Simultaneously, the biomimetic microvascular channel 3 allows for postoperative infusion of nutrient solution or guidance of host blood vessel growth, while the neurotrophic factor 5 actively guides Schwann cell migration and axonal directional growth, providing time-controllable nutritional support. This transforms passive waiting into active guidance, significantly accelerating the regeneration process and improving the quality of nerve repair.

[0044] Optional, such as Figure 1 and Figure 4 As shown, the conduction unit includes an outer encapsulation sheath 1 and a biomimetic parallel signal conduction bundle 2. Both the outer encapsulation sheath 1 and the biomimetic parallel signal conduction bundle 2 are connected to the signal coupling interface 7. There are multiple biomimetic parallel signal conduction bundles 2, which are spaced apart inside the outer encapsulation sheath 1. A biomimetic microvascular channel 3 is disposed inside the outer encapsulation sheath 1.

[0045] In the above optional embodiments, it should be noted that the outer encapsulation sheath 1 is made of a biocompatible flexible material, such as medical silicone, polyurethane, or biodegradable materials such as PLGA, chitosan, etc. The outer encapsulation sheath 1 integrates a biomimetic microvascular channel 3, which has a diameter of 50-200 μm and is prepared by sacrificial template method or soft photolithography technology. It is used for postoperative perfusion of nutrient solution or guiding host blood vessel growth.

[0046] The beneficial effects of the above optional embodiments are as follows: by setting up the biomimetic parallel signal transmission bundle 2, the multi-axon insulating transmission topology of natural nerves is simulated at the microscale, ensuring the independent and parallel transmission of different modalities of neural information such as movement and sensation, realizing high-fidelity transmission of neural information, and providing precise functional instructions for distant target organs.

[0047] Optional, such as Figure 1 and Figure 3 As shown, the conduction unit also includes a neural peritunic bionic layer 4. There are multiple neural peritunic bionic layers 4, and each of the multiple neural peritunic bionic layers 4 covers multiple bionic parallel signal conduction bundles 2 in a one-to-one correspondence.

[0048] In the above optional embodiments, it should be noted that the nerve bundle membrane biomimetic layer 4 covers the inner conduction bundle and is composed of a type IV collagen fiber membrane or a PLCL composite nanofiber membrane prepared by electrospinning technology. The fiber diameter of the fiber membrane is 100-500nm, which plays a selective barrier role.

[0049] The beneficial effect of the above optional embodiments is that the bionic parallel signal transmission bundle 2 can be selectively protected by the bionic percutaneous membrane bionic layer 4.

[0050] Optional, such as Figure 1 , Figure 2 and Figure 5 As shown, the biomimetic parallel signal conduction bundle 2 includes conductive units 6 and longitudinal regeneration channels. There are multiple conductive units 6, each of which is connected to a signal coupling interface 7. The multiple conductive units 6 are spaced apart within the biomimetic nerve peritunic layer 4. There are multiple longitudinal regeneration channels, each of which is located within a conductive unit. Each longitudinal regeneration channel is located between two conductive units 6. The longitudinal regeneration channels are filled with type I collagen or GelMA hydrogel.

[0051] In the above optional embodiments, it should be noted that the shape of each conductive unit 6 simulates the structure of a myelin sheath axon, and the shape is the same as that of the myelin sheath axon. Specifically, it can take the following forms: Microfiber bundle configuration: Core-shell fibers are prepared using coaxial microfluidic spinning technology. The core layer is a continuous gallium-indium alloy, and the outer shell is an insulating material of TPU (thermoplastic polyurethane) or PCL. The outer diameter of a single core-shell fiber is between 10 μm and 50 μm.

[0052] By programmatically modulating the spinning parameters, the shell can be made to form segmental thickness variations, partially mimicking the Ranvier's segment structure of the myelin sheath.

[0053] Embedded channel form: Multiple parallel microchannels are fabricated in an insulating substrate such as silicone or PLA using soft lithography or micromolding techniques. The width of the parallel microchannels is between 5μm and 20μm, and the depth is between 5μm and 20μm. Then, liquid metal is injected into the center of the channel through vacuum-assisted filling to form a conductive path.

[0054] Printed conductor form: On a flexible substrate (such as a polyimide film), liquid metal ink is used to prepare multiple parallel conductors with a linewidth between 10μm and 50μm by mask spraying or direct writing printing technology. Then, SU-8 photoresist or Piriton is spin-coated for insulating encapsulation.

[0055] Each longitudinal regeneration channel is formed naturally by ordered bundling of multiple conductive units 6 in the biomimetic parallel signal transmission bundle 2, with deliberately reserved longitudinal gaps between 5μm and 20μm in width, and filled with collagen or GelMA hydrogel, providing a physical path for axon regeneration.

[0056] The longitudinal regeneration channel is also equipped with neurotrophic factor 5: Specifically, neurotrophic factors 5 such as NGF and BDNF are loaded into PLGA biodegradable microspheres prepared by emulsion solvent evaporation method and mixed into the hydrogel of the regeneration channel; at the same time, small molecule nutrients such as vitamin B12 can be directly melt-blended with the PCL material of the fiber shell to achieve long-term release.

[0057] Neurotrophic factor 5 is composed of PLGA biodegradable microspheres loaded with neurotrophic factor 5 and / or small molecule nutrients melt-blended with an insulating shell material; the PLGA microspheres are prepared by emulsion solvent evaporation method and have a diameter between 1 μm and 10 μm.

[0058] The number of conductive units 6 is greater than 100, preferably thousands.

[0059] The beneficial effects of the above optional embodiments are as follows: by constructing thousands of independent conductive units 6 that are mutually insulated at the microscale and simulating the segmental insulation structure of the axonal myelin sheath, the independence and integrity of neural information encoding are fundamentally guaranteed, and high-fidelity signal transmission is achieved.

[0060] By combining the longitudinal regeneration pathway with the setting of neurotrophic factor 5, Schwann cells can be actively guided to migrate and axons can be directed to grow, and time-controlled nutritional support can be provided. This transforms passive waiting into active guidance, significantly accelerating the regeneration process and improving the quality of repair.

[0061] Optional, such as Figure 1 As shown, the diameter of the biomimetic microvascular channel 3 is between 50 μm and 200 μm.

[0062] In the above optional embodiments, it should be noted that the microvascular channels are prepared by sacrificial template method or soft photolithography and are used to deliver nutrient solution or guide blood vessel regeneration.

[0063] Optional, such as Figure 1 and Figure 2 As shown, the conductivity of conductive unit 6 is greater than 10. 5 S / m.

[0064] Optional, such as Figure 1 and Figure 4 As shown, the outer encapsulation sheath 1 is made of a biocompatible flexible material or a biodegradable material.

[0065] Optional, such as Figures 1 to 3 As shown, each conductive unit 6 includes an insulating layer and a conductive layer. Both the conductive layer and the insulating layer are disposed within the neural permeabilized layer 4. The insulating layer covers the conductive layer, and the conductive layer is also used to connect with the signal coupling interface 7.

[0066] The advantages of the above optional embodiments are: the conductive layer can form a conductive path, and the insulating layer can prevent short circuits from occurring.

[0067] Optional, such as Figure 4 and Figure 5 As shown, the signal coupling interface 7 is a conductive hydrogel interface or a liquid metal nanocomposite gel interface.

[0068] In the above optional embodiments, it should be noted that the signal coupling interface 7 can be an anisotropic conductive hydrogel interface, a liquid metal nanocomposite gel interface, or a multilayer gradient structure interface.

[0069] The beneficial effects of the above optional embodiments are as follows: the anisotropic conductive hydrogel prepared by electric field-induced assembly forms a directional hybrid conductive network inside, realizing efficient and low-loss conversion between ion flow and electron flow, and ensuring the complete transmission of weak nerve signals.

[0070] Preferably, this artificial nerve is a standardized artificial product. The number, length, and diameter of the biomimetic parallel signal conduction bundle 2 can be flexibly designed according to the specific defect. It can be prepared in a controllable manner through mature microfabrication technologies such as coaxial microfluidics and soft lithography, without being limited by the donor. It provides a highly efficient solution for the repair of long and complex nerve defects that can be mass-produced.

[0071] According to an embodiment of a second aspect of the present invention, a method for preparing an implantable artificial nerve, for preparing an implantable artificial nerve, further includes the following steps: Step S100: Prepare and integrate neurotrophic factor 5, conduction unit and biomimetic microvascular channel 3; Step S200: Construct the signal coupling interface 7 and connect the conduction unit to the signal coupling interface 7; Step S300: Preparation complete.

[0072] In the above optional embodiments, it should be noted that, Specifically, implantable artificial nerves can be prepared using the following two methods: Method 1: A multi-layered method based on coaxial microfluidic spinning technology, integrating a nerve regeneration unit, is used to prepare the product. The specific steps are as follows: Step S100: The method for preparing and integrating neurotrophic factor 5, conduction unit, and biomimetic microvascular channel 3 includes the following steps: First, microspheres were prepared using a water-oil-water (W / O / W) double emulsion solvent evaporation method. First, 10 mg of nerve growth factor (NGF) was dissolved in 100 μL of deionized water as the inner aqueous phase. The inner aqueous phase was then added dropwise to 2 mL of dichloromethane (containing 100 mg PLGA, a lactate-glycolic acid ratio of 50:50, and a molecular weight of ~50,000). The mixture was ultrasonically treated at 300 W for 30 seconds in an ice-water bath to form a pre-emulsion. This pre-emulsion was then rapidly poured into 50 mL of a 1% (w / v) polyvinyl alcohol (PVA) aqueous solution and mechanically stirred at 400 rpm for 3 hours to allow complete evaporation of the dichloromethane and solidification of the microspheres. Finally, the microspheres were collected by centrifugation (10,000 rpm, 10 minutes), washed three times with deionized water, freeze-dried, and stored at 4 °C. The preparation of neurotrophic factor 5 was then complete.

[0073] Next, conductive units 6 and longitudinal regeneration channels of the biomimetic parallel signal transmission bundle 2 were prepared. The conductive unit 6 is a biomimetic axon fiber bundle with a segmented structure. Specifically, the core layer was prepared by heating eutectic gallium indium alloy (EGaIn) in a 60°C water bath until completely melted; the shell layer was prepared by dissolving thermoplastic polyurethane (TPU) particles in N,N-dimethylformamide (DMF) to prepare a 12% (w / v) solution, and magnetically stirring for 8 hours until completely dissolved. Subsequently, the above-mentioned PLGA microspheres were added to the TPU solution at a ratio of 5% (w / w, relative to TPU), and an additional 1% (w / w) of vitamin B12 powder was added. The mixture was then uniformly dispersed by magnetic stirring and brief sonication.

[0074] Using a coaxial needle with an inner diameter of 210 μm and an outer diameter of 410 μm, the core layer and shell layer were connected to two precision injection pumps respectively. The flow rate of the core layer was set to 1.0 mL / h, and the flow rate of the shell layer was set to 5.0 mL / h. To prepare a segmented structure, the flow rate of the shell layer was briefly increased to 8.0 mL / h for 0.5 seconds every 30 seconds using the injection pump control system, held for 2 seconds, and then returned to 5.0 mL / h. This periodic change resulted in segments of uneven thickness in the extruded fiber shell. The extruded nascent fibers were placed in a deionized water coagulation bath at room temperature, where DMF in the TPU shell rapidly precipitated and solidified. The fibers were collected by a take-up roller at 10 rpm, ultimately yielding continuous fibers with a segmented insulating structure to form conductive units 6.

[0075] Approximately 1000 continuous fibers with segmented insulating structures were arranged in parallel in a specially designed fixture to form a fiber bundle with a diameter of approximately 0.8 mm. Type I collagen and GelMA (5% w / v) were mixed at a 1:1 volume ratio, and 0.5% LAP photoinitiator was added as the regeneration channel filling solution. This mixture was then thoroughly permeated into all gaps in the fiber bundle using a vacuum-assisted perfusion method. A 405 nm ultraviolet light source (10 mW / cm²) was used. 2 Irradiation for 60 seconds causes GelMA to crosslink and solidify, thereby forming stable, collagen-rich longitudinal regeneration channels between fibers, forming a biomimetic parallel signal transduction bundle 2.

[0076] Next, the perimembranes biomimetic layer 4 was prepared and encapsulated. Specifically, type IV collagen and PLCL (70:30) were dissolved in hexafluoroisopropanol (HFIP) at a mass ratio of 7:3 to prepare an 8% (w / v) electrospinning solution. Using the aforementioned fiber bundle with a diameter of approximately 0.8 mm as the receiving device, the spinning parameters were set as follows: voltage 18 kV, feed speed 1.0 mL / h, and receiving distance 15 cm. Spinning continued for 40 minutes, forming a dense nanofiber membrane with a thickness of approximately 20 μm on the outer surface of the fiber bundle, namely the perimembranes biomimetic layer 4, thus forming an anisotropic conductive fiber bundle.

[0077] Next, the outer encapsulation sheath 1 and the biomimetic microvessels are integrated. Specifically, the anisotropic conductive fiber bundle, which has already undergone the aforementioned steps, is carefully placed in the center of a cylindrical mold as the core. A sacrificial template is arranged, with four 150μm diameter PVA sacrificial fibers evenly distributed around the core fiber bundle, parallel to it and at a certain distance from the fiber bundle. The mixed medical silicone precursor is slowly injected into the mold, ensuring that the silicone completely submerges and encapsulates the core fiber bundle in the center and the PVA sacrificial fibers distributed around it. Subsequently, it is heated and cured at 65°C. After curing, the entire device is immersed in 60°C deionized water, where the PVA fibers dissolve. Thus, four interconnected biomimetic microvessel channels 3 are formed inside the outer silicone encapsulation sheath, surrounding the core fiber bundle.

[0078] In step S200: the signal coupling interface 7 is constructed, and the conduction unit is connected to the signal coupling interface 7 to complete the fabrication. This includes the following steps: First, a prepolymer solution was prepared by dissolving acrylamide (AM, 18% w / v), N,N'-methylenebisacrylamide (MBAA, 0.2% w / v relative to AM), dopamine hydrochloride (3 mg / mL), water-soluble carbon nanotubes (0.5 mg / mL), and photoinitiator LAP (0.5% w / v) in deionized water and ultrasonically dispersing them in an ice-water bath for 30 minutes.

[0079] Next, electric field-induced assembly and in-situ polymerization were performed: the two ends of the conductive unit were placed in specially made parallel platinum electrode molds. Prepolymer liquid was injected into the molds, and then a DC electric field with a strength of 100V / cm was applied for 20 minutes.

[0080] During this period, the negatively charged carbon nanotubes align themselves along the direction of the electric field. The electric field is maintained, and the mold is irradiated with ultraviolet light for 5 minutes to allow the hydrogel to fully polymerize.

[0081] Once the device is removed, a signal coupling interface with significant anisotropic conductivity and wet adhesion capability can be obtained at both ends of the conductive unit.

[0082] Method 2: Fabrication of biomimetic artificial nerves based on soft lithography and 3D printing This method is an alternative for repairing flattened neural tissues such as the spinal cord, but its design also strictly follows the core idea of ​​multi-layered biomimetic structures and integrated nerve regeneration units, specifically including the following steps: First, a biomimetic parallel signal conduction was fabricated. A male mold with multiple parallel grooves (15 μm wide, 15 μm deep, and 30 μm spacing) was prepared on a silicon wafer using SU-82050 photoresist through standard photolithography. This structure simulates parallel-arranged axons. A polydimethylsiloxane (PDMS) matrix and curing agent were mixed at a 10:1 ratio and poured onto the male mold. After vacuum degassing, the mixture was cured at 80°C for 2 hours. After peeling, a PDMS substrate containing parallel microchannels was obtained. This substrate will serve as the core structural layer of the device.

[0083] Then, vacuum-assisted filling of liquid metal is used. The EGaIn liquid metal is heated to 80°C to reduce its viscosity and filled into the middle section of the microchannel (accounting for about 80% of the total channel length) using a vacuum-assisted method (-80 kPa). Vacuum spaces are reserved at both ends of the channel for constructing the interface.

[0084] Next, longitudinal regeneration channels were created. Using precision 3D printing technology and GelMA (10% w / v, containing 0.5% LAP) bio-ink, multiple hydrogel septa parallel to the direction of the liquid metal microchannels were printed on the upper surface of the aforementioned PDMS substrate. These septa separated adjacent conductive microchannels, and the gaps between the septa (approximately 10-15 μm wide) constituted the longitudinal regeneration channels. After UV curing, this structure was firmly integrated.

[0085] Then, by integrating neurotrophic factor 5, before printing the GelMA regeneration channel, the NGF-loaded PLGA microspheres prepared in Implementation Case 1 were mixed into the GelMA prepolymer solution at a concentration of 1% (w / v) to achieve local sustained release of the trophic factor.

[0086] Then, a neural peritunic biomimetic layer 4 was constructed. By electrospinning, a layer of type IV collagen / PLCL composite nanofiber membrane was deposited on the surface of the PDMS substrate with integrated regeneration channels to form the neural peritunic biomimetic layer 4, which provides a selective barrier.

[0087] Finally, the outer encapsulation sheath 1 and biomimetic microvessels were constructed: the above multilayer structure was placed in a custom mold, and several soluble PVA fibers (approximately 100 μm in diameter) were arranged parallel to each other on its upper surface as a sacrificial template. A thin layer of PDMS precursor (approximately 200 μm thick) was poured to completely cover all structures and then heated to cure. The device was then immersed in warm water to dissolve the PVA fibers, thereby forming parallel biomimetic microvascular channels 3 within the outer encapsulation sheath 1 of the device.

[0088] In step S200: Constructing a signal coupling interface 7 and connecting the conduction unit to the signal coupling interface 7 includes: injecting anisotropic conductive hydrogel prepolymer liquid into the microchannel regions reserved at both ends of the device that are not filled with liquid metal and at the end of the regeneration channel.

[0089] Using electric field-induced assembly and in-situ photopolymerization techniques, a low-impedance interface connecting the host nerve to the internal conductive channel was constructed. This interface also extends through the end of the regeneration channel, providing a bioactive interface for the regenerating axon.

[0090] The advantages of the above optional embodiments are as follows: the above preparation method achieves a relatively reliable preparation of artificial nerves, and the artificial nerve can solve the following technical problems: First, it addresses the problem of mixed signal pathways caused by the lack of microscopic structure in existing technologies. Existing neural conduits or bridging devices typically have one or several macroscopic cavities as their conductive channels, short-circuiting the severed ends of nerve fibers with different functions. This structure cannot simulate the physiological topology of tens of thousands of axons conducting in parallel with insulation in natural nerves. This leads to mutual interference and mixed encoding of neural signals from different modalities such as motor and sensory signals during transmission, resulting in extremely low signal fidelity to distant target organs and hindering accurate functional recovery. The proposed neuron, by constructing thousands of independently insulating conductive units at the microscopic scale and simulating the segmental insulating structure of axons bound to myelin sheaths, fundamentally ensures the independence and integrity of neural information encoding, achieving high-fidelity signal transmission.

[0091] Second, this invention addresses the problem of existing technologies passively waiting for, rather than actively guiding, nerve regeneration. Current methods often passively support nerve growth by filling the ducts with matrix or neurotrophic substances, lacking the physical microenvironment within the ducts to guide the directional extension of axons. Nerve regeneration is a lengthy process requiring precise guidance; simple material filling cannot provide a clear path for Schwann cell migration and axon tip growth, resulting in highly random and inefficient regeneration. This new nerve transforms passive regeneration into active regeneration. Through a longitudinal regeneration channel parallel to the long axis of the artificial nerve, it provides a structured, bioactive hydrogel-filled structure for the host nerve's regeneration, actively guiding axons to directionally and orderly cross the defect area and precisely dock with distal ends, thereby significantly improving the speed and quality of regeneration.

[0092] Third, this invention addresses the issues of insufficient biointegration and long-term support. Existing devices have relatively limited functionality and lack continuous support throughout the entire neural regeneration cycle. Nutrient release is uncontrollable, and their structure does not consider deep integration with host tissue. This invention constructs a dynamic support system that utilizes biodegradable, slow-release microspheres loaded with neurotrophic factor 5 to provide sequential nutrient supply at different stages of the regeneration process. Simultaneously, a microchannel system mimicking the epineurium vessels is used to reserve the possibility of subsequent integration with the host circulatory system and long-term substance exchange. Ultimately, this results in an intelligent implantable device that can immediately restore electrophysiological function, continuously support biological regeneration, and ultimately achieve perfect integration with the host nerve.

[0093] The above are merely preferred embodiments of the present invention, but the scope of protection of the present invention is not limited thereto. Any variations or substitutions that can be easily conceived by those skilled in the art within the technical scope disclosed in the present invention should be included within the scope of protection of the present invention. Therefore, the scope of protection of the present invention should be determined by the scope of the claims.

Claims

1. An implantable artificial nerve, characterized in that, include: A signal coupling interface (7) is used to connect to the severed nerve (8). The signal coupling interface (7) is also used to form a conversion impedance between the biological ion flow and the artificial electron flow. The interface impedance at a frequency of 1 kHz is less than 500 Ω·cm. 2 The signal coupling interface (7) is also filled with neurotrophic factors (5). The conduction unit is connected to the signal coupling interface (7) and is filled with neurotrophic factors (5). A biomimetic microvascular channel (3) is provided within the conduction unit and is used to deliver nutrient solution or guide vascular regeneration.

2. The implantable artificial nerve according to claim 1, characterized in that, The conductive unit includes: An outer encapsulation sheath (1) is connected to the signal coupling interface (7), and the biomimetic microvascular channel (3) is provided inside the outer encapsulation sheath (1). A biomimetic parallel signal conduction bundle (2) is provided. There are multiple biomimetic parallel signal conduction bundles (2), and all of them are connected to the signal coupling interface (7). The multiple biomimetic parallel signal conduction bundles (2) are arranged at intervals inside the outer encapsulation sheath (1).

3. The implantable artificial nerve according to claim 2, characterized in that, The conductive unit further includes: There are multiple biomimetic nerve peritunes (4), and each of the multiple biomimetic nerve peritunes (4) covers one of the multiple biomimetic parallel signal conduction bundles (2).

4. The implantable artificial nerve according to claim 3, characterized in that, The biomimetic parallel signal transmission bundle (2) includes: Conductive unit (6), the number of conductive units (6) is multiple, and multiple conductive units (6) are connected to the signal coupling interface (7). Multiple conductive units (6) are arranged at intervals in the nerve peritunic bionic layer (4). The longitudinal regeneration channel has multiple longitudinal regeneration channels, each of which is disposed within the conductive unit. Each longitudinal regeneration channel is located between two conductive units (6), and the longitudinal regeneration channel is filled with type I collagen or GelMA hydrogel.

5. The implantable artificial nerve according to any one of claims 1 to 4, characterized in that, The diameter of the biomimetic microvascular channel (3) is between 50 μm and 200 μm.

6. The implantable artificial nerve according to claim 4, characterized in that, The conductivity of the conductive unit (6) is greater than 10. 5 S / m.

7. The implantable artificial nerve according to claim 2, characterized in that, The outer encapsulation sheath (1) is a biocompatible flexible material or a biodegradable material.

8. The implantable artificial nerve according to claim 4, characterized in that, Each of the conductive units (6) includes: A conductive layer is disposed within the biomimetic layer (4) of the nerve peritunic and is used to connect with the signal coupling interface (7); An insulating layer is placed inside the biomimetic layer (4) of the nerve bundle membrane and covers the conductive layer.

9. The implantable artificial nerve according to claim 1, characterized in that, The signal coupling interface (7) is a conductive hydrogel interface or a liquid metal nanocomposite gel interface.

10. A method for preparing an implantable artificial nerve, used to prepare the implantable artificial nerve according to any one of claims 1 to 9, characterized in that, Includes the following steps: Step S100: Prepare and integrate neurotrophic factors (5), conduction units and biomimetic microvascular channels (3); Step S200: Construct a signal coupling interface (7) and connect the conduction unit to the signal coupling interface (7); Step S300: Preparation complete.

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