An implantable bioelectrode with a structure similar to a cuscuta haustorial structure and a preparation method thereof

Abstract

The application provides an implantable bioelectrode with a structure simulating a Cuscuta suction organ and a preparation method thereof, and relates to the technical field of medical devices. The implantable bioelectrode comprises an electrode substrate and a biomimetic interface coating arranged on the surface of the electrode substrate; the biomimetic interface coating comprises, from inside to outside, a conductive buffer layer, an imitated invasive microfilament layer and a tissue adhesion layer. The implantable bioelectrode utilizes the synergistic effect of the three-layer biomimetic interface to improve the connection stability. The outer adhesion film provides rapid and non-invasive in-situ fixation, the middle nanowire array realizes minimally invasive probe-type anchoring, and the inner hydrogel provides mechanical buffering and eliminates modulus mismatch. The structure effectively overcomes the interface wear and signal attenuation caused by physiological micro-motion, and realizes long-term, stable and low-impedance adhesion.

Description

Technical Field

[0001] This invention relates to the field of medical device technology, and more specifically, to an implantable bioelectrode with a dodder-like suction device structure and its preparation method. Background Technology

[0002] Implantable bioelectrodes have broad application prospects in medical fields such as neuromodulation, cardiac pacing, brain-computer interfaces, and deep brain stimulation. After implantation, the interfacial stability between the electrode and biological tissue directly determines the device's lifespan and signal quality. Ideally, the electrode interface should form a tight, stable, and low-impedance electrical connection with the surrounding tissue, while avoiding fretting noise and interfacial damage caused by relative tissue movement.

[0003] To address the issue of electrode-tissue interface stability, existing technical approaches mainly include external fixation, surface modification, and microstructure construction. For example, electrodes can be fixed to the target tissue surface using surgical sutures or bio-adhesives. Alternatively, a layer of soft materials such as hydrogels or conductive polymers can be coated onto the electrode surface to alleviate mechanical mismatch by reducing the apparent modulus of the electrode. Furthermore, there are methods that construct microneedle arrays or barbed structures on the electrode surface to enhance anchoring force through mechanical interlocking effects.

[0004] However, existing implantable electrodes generally face the problem of insufficient interfacial stability in practical applications. A significant mechanical mismatch exists between the rigid modulus of traditional metal electrodes and soft tissue, easily leading to gaps between the tissue and the electrode, resulting in substantial attenuation of bioelectrical signals. Under the influence of physiological activities, the relative movement between the electrode and tissue exacerbates mechanical wear at the interface, causing fluctuations in electrode impedance. When using existing solutions, suturing methods cause trauma, bioadhesive adhesion is insufficient and easily degraded; soft coating materials have weak adhesion to the substrate and are prone to peeling off after long-term use; and microstructures are prone to damaging tissue during implantation, triggering significant foreign body reactions and encapsulation phenomena, which further deteriorates interfacial performance.

[0005] In summary, existing implantable electrodes and their fixation methods generally fail to adapt to the physiological characteristics of tissues, exhibiting drawbacks such as poor tissue adhesion stability, easy degradation of long-term mechanical and electrical connections, and a tendency to induce tissue damage and foreign body reactions. There is an urgent need in this field for an electrode interface technology that can achieve long-term stable adhesion and low-impedance transmission without causing significant tissue damage.

[0006] In view of this, the present invention is hereby proposed. Summary of the Invention

[0007] The purpose of this invention is to provide an implantable bioelectrode with a dodder-like aspirator structure and its preparation method. The electrode achieves minimally invasive seamless anchoring by using a synergistic hydrogel buffer layer, a nanowire microfilament layer, and a catechol adhesion layer to eliminate tissue mechanical mismatch and micro-motion wear, thereby ensuring long-term stable, low-impedance, and high-quality bioelectric signal transmission.

[0008] In order to achieve the above-mentioned objectives of the present invention, the following technical solution is adopted: In a first aspect, the present invention provides an implantable bioelectrode with a dodder-like siphon structure, comprising an electrode substrate and a biomimetic interface coating disposed on the surface of the electrode substrate; The biomimetic interface coating comprises, from the inside out, a conductive buffer layer, an intrusive microfilament layer, and a tissue adhesion layer. The conductive buffer layer is a conductive polymer hydrogel composite material composed of an interpenetrating conductive polymer and a hydrophilic polymer network. The simulated intrusion microfilament layer is an array of oriented conductive polymer nanowires; The tissue adhesion layer is a polymer film containing catechol groups.

[0009] In an optional embodiment, the conductive polymer is selected from at least one of poly(3,4-ethylenedioxythiophene)-polystyrene sulfonate, polypyrrole, and polyaniline; and / or, The hydrophilic polymer includes at least one of polyvinyl alcohol, polyacrylic acid, hyaluronic acid, sodium alginate, and gelatin; and / or, The conductive polymer nanowire array is made of at least one of poly(3,4-ethylenedioxythiophene), polypyrrole, and poly(3,4-ethylenedioxythiophene)-tetracyanoquinone dimethane; and / or The material of the tissue adhesion layer is selected from at least one of polydopamine, dopamine-methacrylate copolymer, dopamine-grafted chitosan, tannic acid coordination complex, and epigallocatechin gallate coordination complex.

[0010] In an optional embodiment, the total thickness of the biomimetic interface coating is 2 μm to 20 μm; and / or, The thickness of the conductive buffer layer is 1μm~10μm; and / or, The length of the nanowires in the simulated intrusion microfilament layer is 500 nm to 5 μm; and / or, The diameter of the nanowires in the simulated intrusion microfilament layer is 50 nm to 500 nm; and / or, The thickness of the tissue adhesion layer is 10 nm to 200 nm.

[0011] In an optional embodiment, the Young's modulus of the biomimetic interface layer exhibits a decreasing gradient transition along the direction from the conductive buffer layer to the tissue adhesion layer; and / or, The Young's modulus of the conductive buffer layer is 1 MPa to 100 MPa; preferably, the Young's modulus of the conductive buffer layer is 10 MPa to 50 MPa; and / or, The Young's modulus of the simulated invasive microfilament layer is 100 kPa to 10 MPa; and / or, The Young's modulus of the tissue adhesion layer is 10 kPa to 1 MPa.

[0012] In an optional embodiment, the electrode substrate includes at least one of a metal substrate, a silicon-based microelectrode array, a carbon-based electrode, and a flexible polymer substrate; Preferably, the metal substrate comprises at least one selected from gold, platinum, iridium, and stainless steel; Preferably, the flexible polymer substrate includes at least one of polyimide, polydimethylsiloxane, and phenelzine.

[0013] Secondly, the present invention provides a method for preparing an implantable bioelectrode with a dodder-like siphon structure as described in any of the foregoing embodiments, comprising: S1, the surface of the electrode substrate is cleaned and treated with oxygen plasma; S2, The conductive polymer is mixed with the hydrophilic polymer and deposited on the surface of the treated electrode substrate by electrodeposition to form the conductive buffer layer; S3, using template-assisted electrodeposition or template-free self-assembly, the conductive polymer nanowire array is deposited or grown on the surface of the conductive buffer layer to form the intrusion-resistant microfilament layer; S4, the electrode containing the conductive buffer layer and the pseudo-invasive microfilament layer is immersed in a buffer solution containing dopamine monomer or catechol group monomer, and the tissue adhesion layer is formed on the surface by self-polymerization deposition.

[0014] Thirdly, the present invention provides a method for preparing an implantable bioelectrode with a dodder-like siphon structure as described in the foregoing embodiments. In step S1, the surface of the electrode substrate is cleaned, including ultrasonic cleaning in acetone, ethanol, and deionized water in sequence; and / or, In step S2, the electrodeposition is performed using cyclic voltammetric electrodeposition or constant potential electrodeposition; and / or, In step S2, the deposition potential of the electrodeposition is 0.5V~1.2V; and / or, In step S2, the electrodeposition time is 5 minutes to 30 minutes; and / or, In step S3, when using the template-assisted electrodeposition method, a porous alumina template or a polycarbonate track etching film is spin-coated onto the surface of the conductive buffer layer; and / or, In step S3, the pore size of the porous alumina template or the polycarbonate track etching film is 50 nm to 500 nm; and / or, In step S3, the pore spacing of the porous alumina template or the polycarbonate track etching film is 100 nm to 1 μm; and / or, In step S3, the deposition potential is 0.6V~1.0V; and / or, In step S3, the deposition time is 10 minutes to 60 minutes; and / or, In step S3, when using the template-assisted electrodeposition method, after deposition is completed, the porous alumina template or the polycarbonate track etching film is removed with a solvent; preferably, the method of removing the porous alumina template or the polycarbonate track etching film with a solvent is: dissolving the porous alumina template with sodium hydroxide solution, or dissolving the polycarbonate track etching film with dichloromethane; and / or, In step S4, the buffer solution is a weakly alkaline buffer solution with a pH value of 8.0~9.0; and / or, In step S4, self-polymerization deposition is carried out at room temperature for 12 to 48 hours.

[0015] Fourthly, the present invention provides an implantable medical device, including an implantable bioelectrode with a dodder-inspired suction device structure as described in any of the foregoing embodiments.

[0016] In an optional embodiment, the implantable medical device includes at least one of a neural recording device, a cardiac pacing device, and a deep brain stimulation device.

[0017] Fifthly, the present invention provides an implantable medical device, including the implantable medical device as described in the foregoing embodiments; Preferably, the implantable medical device further includes a control module electrically connected to the implantable bioelectrode of the dodder-inspired suction device structure or the implantable medical device; more preferably, the control module includes at least one of a signal processing unit, a pulse generation unit, and a power supply unit.

[0018] Compared with existing technologies, the implantable bioelectrode provided in this application fundamentally improves the long-term interfacial stability between the electrode and biological tissue by constructing a biomimetic interface coating with layered functional characteristics on the substrate surface. The outermost tissue adhesion layer contains catechol groups, which can form multiple intermolecular interactions with the tissue surface in a dynamic physiological environment that is moist and undergoes continuous minor deformation, achieving rapid in-situ fixation in the early stages of implantation. This mechanism provides immediate and firm initial adhesion, effectively overcoming the shortcomings of conventional bio-adhesives that are easily degraded or physical sutures that cause significant trauma.

[0019] The central, non-invasive microfilament layer comprises an array of oriented conductive polymer nanowires. Its microstructure conforms to the contours of soft tissue surfaces, partially extending and forming a probe-like physical anchor. This layer structure provides sustained interfacial adhesion while avoiding mechanical puncture and damage to surrounding tissues by coarse microneedles, thus significantly reducing the risk of foreign body reaction and gelatinous scarring after implantation.

[0020] The innermost conductive buffer layer is composed of a hydrogel composite material formed by the interpenetration of conductive polymers and hydrophilic polymer networks. It plays a crucial role in smoothly transitioning the mechanical modulus between the rigid electrode substrate and the biological tissue. This buffer layer effectively absorbs the interfacial mechanical shock stress generated by physiological micro-movements such as heartbeat and respiration, eliminating stress concentration and coating peeling caused by mechanical mismatch, and maintaining an excellent charge transport network. The synergistic effect of these three layers enables the electrode interface to maintain a tight, low-impedance physical and electrical connection in a dynamic physiological environment, completely solving the problems of high noise and signal attenuation caused by interfacial micro-movements. Attached Figure Description

[0021] To more clearly illustrate the specific embodiments of the present invention or the technical solutions in the prior art, the drawings used in the description of the specific embodiments or the prior art will be briefly introduced below. Obviously, the drawings described below are some embodiments of the present invention. For those skilled in the art, other drawings can be obtained from these drawings without creative effort.

[0022] Figure 1 A schematic diagram of the interface structure of the implantable bioelectrode with a dodder-like suction device structure provided in an embodiment of the present invention; Figure 2 This is a scanning electron microscope (SEM) image of the cross-section of the electrode-tissue composite 28 days after the biomimetic bioelectrode was implanted into the gastrocnemius muscle tissue in Example 1 of the present invention (scale bar A 20 μm, B 10 μm). Figure 3This is a comparison of GFAP and Iba-1 immunofluorescence staining and glial scar assessment of the surrounding tissue 4 weeks after the biomimetic bioelectrode and ordinary silicon-based electrode were implanted into the brain tissue in Example 2 of the present invention.

[0023] Explanation of key component symbols: 100 - implantable bioelectrode with a dodder-like suction device structure; 1 - electrode substrate; 2 - biomimetic interface coating; 21 - conductive buffer layer; 22 - imitation invasive microfilament layer; 23 - tissue adhesion layer. Detailed Implementation

[0024] The embodiments of the present invention will be described in detail below with reference to examples. However, those skilled in the art will understand that the following examples are for illustrative purposes only and should not be considered as limiting the scope of the invention. Unless otherwise specified in the examples, conventional conditions or conditions recommended by the manufacturer are followed. Reagents or instruments whose manufacturers are not specified are all commercially available conventional products.

[0025] This application provides an implantable bioelectrode with a dodder-inspired suction device structure, including an electrode substrate and a biomimetic interface coating disposed on the surface of the electrode substrate.

[0026] The biomimetic interface coating comprises, from the inside out, a conductive buffer layer, an intrusive microfilament layer, and a tissue adhesion layer.

[0027] The conductive buffer layer is a conductive polymer hydrogel composite material composed of an interpenetrating conductive polymer and a hydrophilic polymer network. The simulated intrusion microfilament layer is an array of oriented conductive polymer nanowires; The tissue adhesion layer is a polymer film containing catechol groups.

[0028] The implantable bioelectrode, with its dodder-inspired aspirator structure, is designed to conform to and utilize the physiological characteristics of biological tissues—soft, moist, and in a dynamic micro-deformation environment—in both its macroscopic and microscopic structures. For example... Figure 1 As shown, the implantable bioelectrode includes an electrode substrate and a biomimetic interface coating disposed on the surface of the electrode substrate. The biomimetic interface coating comprises, from the inside out (i.e., from near the surface of the electrode substrate towards the external biological tissue), a conductive buffer layer, a pseudo-invasive microfilament layer, and a tissue adhesion layer. The conductive buffer layer, the pseudo-invasive microfilament layer, and the tissue adhesion layer are constructed in layers in terms of spatial sequence, mechanical properties, and chemical function, jointly simulating the layered synergistic mechanism of the haustorium of the holoparasitic plant dodder during its development and anchoring process.

[0029] The electrode substrate serves as the main mechanical support and electrical conduction pathway base for the electrode, supporting the external biomimetic interface coating and enabling physical and electrical connections with external circuits or devices. The electrode substrate can be made of conductive support materials with appropriate rigidity or flexibility, depending on the specific clinical recording or stimulation requirements.

[0030] The innermost conductive buffer layer of the biomimetic interface coating is tightly attached to and bonded to the surface of the electrode substrate. This conductive buffer layer is a conductive polymer hydrogel composite material composed of an interpenetrating network of conductive polymers and hydrophilic polymers. This interpenetrating network topology provides continuous electron and ion conduction pathways through the conductive polymer network, while simultaneously locking in moisture through the three-dimensional cross-linked network of hydrophilic polymers. This composite material, composed of an interpenetrating network, exhibits excellent elasticity and mechanical absorption capacity as a buffer pad in terms of physical and mechanical properties. In a dynamic implantation environment, the conductive buffer layer acts as an electrical transition layer, maintaining extremely low interfacial contact resistance. Furthermore, it actively absorbs and dissipates dynamic mechanical impact stress between the electrode and tissue caused by breathing, heartbeat, or limb movement, dispersing interfacial shear stress. This fundamentally alleviates the risk of overall coating cracking and peeling caused by mechanical mismatch between the underlying substrate and external soft tissue.

[0031] The pseudo-invasive microfilament layer disposed on the outer surface of the conductive buffer layer is an array of oriented conductive polymer nanowires. This nanowire array exhibits a highly ordered orientation in the microscopic space, and its micromorphology is designed to mimic the physical probe structure of a dodder haustorium invading hyphae. In the implanted state, this nanowire array can conform to the microscopic topological irregularities of the biological tissue surface. Some nanowires, due to their microscopic size and moderate flexibility, can slightly extend and embed into the superficial layer of the tissue, providing a continuous and stable physical anchoring force through a microscopic topological mechanical interlocking effect without piercing or damaging individual tissue cells. Because the nanowire array exhibits an extremely high specific surface area, it constructs a highly efficient charge transport network in one dimension, thereby significantly enhancing the charge injection capability of the electrode and significantly reducing the electrochemical impedance of the interface, effectively suppressing baseline drift and micro-motion artifact noise, and improving the signal-to-noise ratio of signal recording.

[0032] The tissue adhesion layer, located at the outermost layer of the biomimetic interface coating, is a polymer film containing catechol groups. This tissue adhesion layer directly faces the biological tissue interface. The catechol groups abundant within the film, in the moist physiological fluid environment of the organism, can rapidly form multiple intermolecular hydrogen bonds and π-π stacking interactions with macromolecules such as proteins, glycoproteins, and polysaccharides on the surface of the biological tissue, generating immediate chemical adhesion in the early stages of implantation. This tissue adhesion layer provides immediate, in-situ, non-invasive fixation in the initial and early stages of electrode implantation, enabling a tight and seamless fit between the electrode surface and the tissue surface before the nanowire array has fully established a stable topological interlock with the tissue.

[0033] The three-layer structure of the biomimetic interface coating achieves a complete closed loop of "adhesion-invasion-anchoring" through the synergistic effects of interfacial chemistry, physical spatial topology, and mechanical conduction. Along the direction from the conductive buffer layer to the tissue adhesion layer, the entire biomimetic interface coating achieves a smooth transition in mechanical properties, ensuring that there are no abrupt mechanical discontinuities between the underlying substrate and the extremely soft biological tissue. This synergistic design significantly reduces the probability of foreign body reactions in local tissues due to long-term interfacial friction, ensuring the physiological safety of minimally invasive procedures while guaranteeing the extreme stability of electrical recording and stimulation during the long-term working life of the implanted microelectrode.

[0034] In summary, the implantable bioelectrode provided in this embodiment significantly improves interface stability and signal transduction quality under dynamic physiological conditions by constructing a three-layer composite biomimetic interface coating. The outermost tissue adhesion layer containing catechol groups achieves rapid and non-invasive initial in-situ fixation; the middle layer, an array of oriented conductive polymer nanowires, conforms to the tissue microstructure and provides continuous minimally invasive anchoring while avoiding severe mechanical damage and foreign body reactions; the innermost hydrogel composite material provides mechanical buffering and modulus transition, eliminating mechanical mismatch between the substrate and the tissue. These three layers work synergistically to prevent interface wear and electrode gaps caused by physiological micro-movements, while ensuring long-term stable, low-impedance, and tight adhesion between the electrode and the tissue.

[0035] In the following embodiments, in order to achieve electrochemical stability, mechanical compliance and high interfacial adhesion of the biomimetic interface coating in a humid physiological environment, the chemical materials constituting the three-layer structure of the biomimetic interface coating were specifically optimized.

[0036] In some embodiments, the conductive polymer is selected from at least one of poly(3,4-ethylenedioxythiophene)-polystyrene sulfonate, polypyrrole, and polyaniline.

[0037] In some embodiments, the hydrophilic polymer includes at least one of polyvinyl alcohol, polyacrylic acid, hyaluronic acid, sodium alginate, and gelatin.

[0038] The conductive buffer layer aims to construct a dual network between a rigid substrate and soft tissue, possessing both charge conduction and mechanical absorption capabilities. Specifically, the conductive polymer in the conductive buffer layer is selected from at least one of poly(3,4-ethylenedioxythiophene)-polystyrene sulfonate (PEDOT:PSS), polypyrrole (PPy), and polyaniline (PANI). These intrinsically conductive polymers, with their large conjugated π-bond systems, constitute the high-speed electronic conduction framework of the interface. Simultaneously, to impart excellent flexibility and tissue mechanical compatibility to the buffer layer, the hydrophilic polymer forming the interpenetrating network with the aforementioned conductive polymer is selected from at least one of polyvinyl alcohol (PVA), polyacrylic acid (PAA), hyaluronic acid, sodium alginate, and gelatin. These hydrophilic polymer chains are rich in strongly hydrophilic groups such as hydroxyl and carboxyl groups, enabling them to absorb water and swell in physiological environments, forming a highly hydrated, soft, three-dimensional network. The rigid conductive framework interspersed within the soft hydrogel network synergistically achieves effective dissipation of mechanical impact stress and stable transmission of low-impedance signals.

[0039] In some embodiments, the material of the conductive polymer nanowire array includes at least one of poly(3,4-ethylenedioxythiophene), polypyrrole, and poly(3,4-ethylenedioxythiophene)-tetracyanoquinone dimethane.

[0040] The intrusion-resistant microfilament layer disposed outside the conductive buffer layer is a nanowire array with a high aspect ratio. The material of this array is specifically selected from at least one of poly(3,4-ethylenedioxythiophene) (PEDOT), polypyrrole (PPy), and poly(3,4-ethylenedioxythiophene)-tetracyanoquinone dimethane (PEDOT:TCNQ). These materials not only possess excellent electrochemical activity and biocompatibility but can also form one-dimensional continuous charge transport channels through polymerization processes. More importantly, the microfilaments constructed from these polymer materials possess moderate flexibility at the microscale. When they conform to and embed into the superficial layers of tissue, they can undergo adaptive bending, thereby providing topological mechanical anchoring force while avoiding rigid mechanical penetration and damage to surrounding tissue cells.

[0041] In some alternative embodiments, the pseudo-invasive microfilament layer may also be replaced by an array of oriented carbon nanotubes or a metal nanowire array instead of conductive polymer nanowires. However, considering the interfacial chemical bonding force and mechanical modulus matching with the inner conductive polymer hydrogel, the conductive polymer nanowire array is the most preferred solution in this application.

[0042] In some embodiments, the material of the tissue adhesion layer is selected from at least one of polydopamine, dopamine-methacrylate copolymer, dopamine-grafted chitosan, tannic acid coordination complex, and epigallocatechin gallate coordination complex.

[0043] Located on the outermost side of the interface, the tissue adhesion layer, which directly contacts the biological tissue, has the core function of providing immediate in-situ chemical fixation during the initial implantation phase. To this end, the material of the tissue adhesion layer is selected from at least one of polydopamine (PDA), dopamine-methacrylate copolymer, dopamine-grafted chitosan, tannic acid coordination complexes, and epigallocatechin gallate (EGCG) coordination complexes. A common chemical characteristic of these polymers or complexes is that their side chains or backbones are rich in highly reactive catechol groups or gallic acid and other polyphenolic groups. In the moist fluid environment of a living organism, these polyphenolic groups can rapidly and strongly interact with macromolecules such as proteins and glycoproteins on the surface of biological tissues, forming dense multiple intermolecular hydrogen bonds, π-π stacking interactions, coordination bonds, or covalent chemical crosslinks, resulting in extremely excellent wet interfacial adhesion. This chemical adhesion mechanism, in conjunction with the physical topological anchoring of the inner layer, jointly constructs a long-term stable biomimetic fixation interface.

[0044] In this embodiment, to ensure that the biomimetic interface coating can simultaneously meet the multiple requirements of gradual change in mechanical modulus, minimally invasive tissue anchoring, and high-resolution electrical signal transmission after implantation, the microscopic spatial dimensions of the coating and its internal layered structures were precisely designed and defined.

[0045] Specifically, in some embodiments, the total thickness of the biomimetic interface coating is 2μm to 20μm; for example, it can be 2μm, 4μm, 6μm, 8μm, 10μm, 12μm, 14μm, 16μm, 18μm, 20μm, etc.

[0046] The total thickness setting balances the mechanical stability and electrical resolution of the interface. If the thickness is below the lower limit, it will lead to insufficient mechanical buffer space and a decrease in interface anchoring strength; while if the thickness is above the upper limit, the spatiotemporal resolution of the high-density array electrodes during neural recording or local electrical stimulation may be reduced due to excessive non-metallic coating.

[0047] In some embodiments, the thickness of the conductive buffer layer is 1 μm to 10 μm. For example, it can be 1 μm, 2 μm, 3 μm, 4 μm, 5 μm, 6 μm, 7 μm, 8 μm, 9 μm, 10 μm, etc.

[0048] This range endows hydrogel composites with sufficient microscopic water absorption and swelling volume, enabling them to form effective elastic damping pads on rigid substrate surfaces, fully absorbing and dissipating mechanical impact stresses generated by physiological activities.

[0049] In some embodiments, the length of the nanowires in the pseudo-invasive microfilament layer is 500 nm to 5 μm.

[0050] In some embodiments, the diameter of the nanowires in the pseudo-invasive microfilament layer is 50 nm to 500 nm.

[0051] The aforementioned pseudo-invasive microfilament layer located in the middle layer contains nanowires with lengths controlled between 500 nm and 5 μm; for example, they can be 500 nm, 800 nm, 1 μm, 1.5 μm, 2 μm, 2.5 μm, 3 μm, 3.5 μm, 4.5 μm, 5 μm, etc. The diameter of the nanowires is controlled between 50 nm and 500 nm; for example, they can be 50 nm, 100 nm, 150 nm, 200 nm, 250 nm, 300 nm, 350 nm, 400 nm, 450 nm, 500 nm, etc. Unlike existing artificial microneedle structures of tens to hundreds of micrometers that easily cause severe tissue damage, the nanowire array size of this solution is much smaller than that of conventional biological cells. When implanted into tissue, these nanowires, which combine a high aspect ratio and moderate flexibility, can effectively undergo adaptive bending, conforming to the microscopic undulations of the soft tissue surface, allowing their tips to slightly extend and embed into the tissue surface. This biomimetic topological design achieves a powerful minimally invasive mechanical interlock without puncturing cells or causing significant glial scarring and foreign body reactions.

[0052] This biomimetic topological design achieves strong minimally invasive mechanical interlocking without puncturing cells or causing significant glial scarring or foreign body reactions. Furthermore, in the implanted state, due to its specific modulus and size design, the depth of the pseudo-invasive microfilament layer embedded in the surface of biological tissue is controlled to no more than 30% to 50% of the nanowire length. This controlled shallow penetration depth ensures both sufficient microscopic topological mechanical interlocking force and prevents the nanowire tip from damaging the cell membrane integrity of individual cells due to excessive penetration, further guaranteeing the minimally invasive safety of long-term implanted tissue from a geometric perspective.

[0053] Meanwhile, the nanowire arrays at the above-mentioned size have a high specific surface area, which significantly enhances the charge injection capability of the interface and reduces the electrochemical impedance.

[0054] In some embodiments, the thickness of the tissue adhesion layer is 10 nm to 200 nm.

[0055] The thickness of the outermost tissue adhesion layer is controlled between 10 nm and 200 nm. For example, it can be 10 nm, 30 nm, 50 nm, 80 nm, 100 nm, 120 nm, 140 nm, 160 nm, 180 nm, 200 nm, etc. This nanoscale ultrathin thickness design provides a sufficient density of catechol groups to ensure rapid in-situ chemical fixation in the early stages of implantation, while preventing the coating from being too thick and altering or even obscuring the microstructure of the underlying nanowire array. This ensures the complete preservation of the topological probe effect of the microfilament layer and the highly sensitive electrical properties of the interface.

[0056] In some embodiments, the Young's modulus of the biomimetic interface layer exhibits a decreasing gradient transition along the direction from the conductive buffer layer to the tissue adhesion layer.

[0057] In this embodiment, to address the problems of interfacial stress concentration, coating peeling, and tissue damage caused by significant mechanical mismatch between the rigid electrode substrate and soft biological tissue, the biomimetic interfacial coating incorporates a biomimetic gradient design in terms of material mechanical properties. Along the direction from the conductive buffer layer to the tissue adhesion layer (i.e., from the inside out), the Young's modulus of the biomimetic interfacial coating exhibits a smoothly decreasing gradient transition. This gradient transition design creates a continuous mechanical buffer zone between the electrode substrate, with its stiffness reaching several GPa, and the soft biological tissue, effectively dispersing the interfacial shear stress caused by heartbeat, respiration, or limb movement, significantly improving the long-term peel strength and adhesion stability of the device under dynamic physiological environments.

[0058] In some embodiments, the Young's modulus of the conductive buffer layer is 1 MPa to 100 MPa; preferably, the Young's modulus of the conductive buffer layer is 10 MPa to 50 MPa.

[0059] In some embodiments, the Young's modulus of the simulated invasive microfilament layer is 100 kPa to 10 MPa.

[0060] In some embodiments, the Young's modulus of the tissue adhesion layer is 10 kPa to 1 MPa.

[0061] Specifically, the conductive buffer layer, tightly bonded to the substrate surface, has a Young's modulus controlled within the range of 1 MPa to 100 MPa; for example, it can be 1 MPa, 10 MPa, 20 MPa, 30 MPa, 50 MPa, 60 MPa, 75 MPa, 85 MPa, 95 MPa, 100 MPa, etc. Preferably, to achieve the optimal balance between mechanical absorption and coating cohesive strength, the Young's modulus of the conductive buffer layer is 10 MPa to 50 MPa; for example, it can be 10 MPa, 15 MPa, 20 MPa, 25 MPa, 30 MPa, 35 MPa, 40 MPa, 45 MPa, 48 MPa, 50 MPa, etc. This modulus range allows the buffer layer to act as a highly efficient elastic shock absorber, significantly absorbing mechanical impacts transmitted from the underlying layer.

[0062] The pseudo-invasive microfilament layer located in the middle layer has an overall effective Young's modulus of 100 kPa to 10 MPa as part of the array structure; for example, it can be 100 kPa, 500 kPa, 1 MPa, 2 MPa, 4 MPa, 5 MPa, 7 MPa, 8 MPa, 9 MPa, 10 MPa, etc. This modulus further approximates soft tissue, giving the nanowire array appropriate microscopic flexibility. During implantation and contact with tissue, the microfilaments within this modulus range tend to undergo adaptive bending when encountering cellular resistance, thereby achieving micro-interval anchoring while avoiding rigid puncture and physical damage to the cell structure.

[0063] The tissue adhesion layer directly facing the tissue has its Young's modulus precisely controlled within the range of 10 kPa to 1 MPa; for example, it can be 10 kPa, 50 kPa, 100 kPa, 200 kPa, 400 kPa, 500 kPa, 700 kPa, 800 kPa, 900 kPa, 1 MPa, etc. This extremely low modulus range achieves a high degree of matching and overlap with the natural mechanical properties of typical biological soft tissues (such as brain tissue, myocardium, etc.), ensuring that the chemical adhesion interface does not generate secondary residual stress that leads to interface delamination when it undergoes dynamic deformation with the tissue, thus achieving an ultimate transition in interfacial mechanical compliance.

[0064] While the above embodiments demonstrate optimal performance under specific parameters, those skilled in the art will understand that as long as the Young's modulus of the conductive buffer layer is within the range of 1 MPa to 100 MPa, it can provide basic elastic energy absorption and mechanical stress dispersion. When it approaches the upper limit of 100 MPa, its internal conjugated polymer network becomes denser, which is more conducive to high-fidelity transmission of high-frequency neural signals. When it approaches the lower limit of 1 MPa, it exhibits superior ultra-soft compliance, making it more suitable for stress-free bonding with extremely low modulus tissues (such as cerebral cortex tissue), while still achieving the core objectives of the present invention: preventing micro-movement detachment and reducing contact impedance. Similarly, the Young's modulus of the pseudo-invasive microfilament layer and the tissue adhesion layer can be dynamically adjusted within their respective defined ranges to meet the intrinsic requirements of multi-level modulus gradient and minimally invasive tissue bonding.

[0065] In this embodiment, in order to meet the needs of high spatial resolution and long-term stability for neural signal recording, cardiac pacing or deep brain stimulation in different clinical medical scenarios, the implantable bioelectrode with a dodder-like suction device structure provided in this application has a variety of targeted material and topological morphology options for its electrode substrate.

[0066] Specifically, in some embodiments, the electrode substrate includes at least one of a metal substrate, a silicon-based microelectrode array, a carbon-based electrode, and a flexible polymer substrate.

[0067] Furthermore, the metal substrate includes at least one of gold, platinum, iridium, and stainless steel.

[0068] In the above scheme, the metal substrate can be an inorganic metal material with excellent biocompatibility and intrinsically high conductivity, specifically including at least one of gold, platinum, iridium, and stainless steel. Such metal substrates are typically fabricated as microfilament electrodes, disk electrodes, or mesh electrodes, providing robust physical support and excellent electrochemical discharge pathways for the overall coating. When a silicon-based microelectrode array is used, the electrode substrate manifests as a microchip integrating multiple micro / nano-scale recording or stimulation sites, enabling high-density, multi-channel, precise capture of local bioelectrical signals.

[0069] Furthermore, the flexible polymer substrate includes at least one of polyimide, polydimethylsiloxane, and phenelzine.

[0070] Polyimide films or phenelzine films are often used as substrates for ultrathin flexible microelectrode arrays, while polydimethylsiloxane flexible films can serve as compliant substrates with high elastic deformation capabilities. In this embodiment, when the above-mentioned flexible polymer substrate is used, the surface of the flexible polymer substrate is usually pre-sputtered or deposited with a nanoscale metal conductive layer (such as a gold conductive layer) to maintain the electrical pathway. Its surface thickness can be, for example, 10 nm, 50 nm, 100 nm, 200 nm, 300 nm, etc.

[0071] The aforementioned electrode substrates with different stiffnesses and material properties, combined with the biomimetic interface coating disposed on their surfaces, exhibit excellent mechanical and electrical synergy. When using rigid metal, silicon-based, or carbon-based electrode substrates, the conductive buffer layer disposed on the innermost side of the substrate can smoothly wrap around its rigid surface. The three-dimensional interpenetrating network of the hydrogel composite material effectively dissipates the mechanical stress transmitted by the micro-motion of external physiological tissue, avoiding direct mechanical wear of soft tissue by the rigid surface, and ensuring the bonding stability of the outer anti-invasive microfilament layer and the tissue adhesion layer. When using flexible polymer substrates such as polyimide, polydimethylsiloxane, or phenelzine, the thin, flexible, or stretchable characteristics of the substrate itself, combined with the biomimetic interface coating with a smooth decrease in Young's modulus on the outer layer, create a performance superposition, constructing an almost completely continuous mechanical gradient transition from the flexible substrate to the ultra-soft biological tissue at the interface. This extreme consistency allows the composite coating on the surface of the flexible bionic electrode to remain structurally intact, without cracking or physical peeling, even when subjected to bending or stretching deformation of up to 50%. This demonstrates exceptional tissue compliance, significantly reducing glial astrocyte activation and chronic inflammatory response caused by long-term implantation, and achieving the ultimate technical goal of minimally invasive seamless adhesion.

[0072] This application also provides a method for preparing an implantable bioelectrode with a dodder-inspired siphon structure as described in the foregoing embodiments. This method employs a bottom-up surface physical modification and electrochemical process, sequentially including the following steps S1 to S4: Step S1: Clean the surface of the electrode substrate and perform oxygen plasma treatment.

[0073] First, grease, organic impurities, and dust adhering to the electrode substrate surface are removed through physical and chemical solvent cleaning. Then, the substrate surface is subjected to high-energy bombardment treatment using oxygen plasma. This oxygen plasma treatment not only deeply etches and removes residual organic matter from the surface, but also breaks bonds in the substrate surface and introduces a high density of oxygen-containing polar functional groups such as hydroxyl groups in situ. This treatment transforms the substrate surface into a highly hydrophilic and chemically active interface, significantly reducing the contact angle of the subsequent aqueous electrolyte. This provides the necessary chemical bonding sites for strong adhesion between the underlying coating and the substrate, forming the cornerstone of the process that ensures the coating does not peel off over a long period.

[0074] Step S2: The conductive polymer is mixed with the hydrophilic polymer and deposited on the surface of the treated electrode substrate by electrodeposition to form the conductive buffer layer.

[0075] An electrochemical polymerization and co-deposition process is carried out on the substrate surface serving as the working electrode under an applied electric field, using a mixed solution containing conductive polymer monomers or precursors and hydrophilic polymers as the electrolyte. Simultaneously, the flexible long chains of the hydrophilic polymers are physically encapsulated and interpenetrated in situ, thereby constructing a molecularly entangled interpenetrating polymer network on the substrate surface. The use of electrodeposition rather than physical coating allows this conductive buffer layer to form a dense electrochemical interface with the activated electrode substrate surface, ensuring excellent interfacial adhesion and continuity of electronic conduction.

[0076] Step S3: The conductive polymer nanowire array is deposited or grown on the surface of the conductive buffer layer using a template-assisted electrodeposition method or a template-free self-assembly method to form the intrusion-resistant microfilament layer.

[0077] When using template-assisted electrodeposition, a nanoscale porous film covering the surface of a conductive buffer layer serves as a spatially confined template. Under the influence of an electric field, the conductive polymer grows only from bottom to top within the confined space of the pores. After deposition, the template framework is removed using a solvent that does not damage the polymer, exposing the vertically oriented microfilament array. When using a template-free self-assembly method, a specific surfactant is introduced into the electrolyte. Its anisotropic micelles at the interface guide the conductive polymer to spontaneously orient along a one-dimensional direction. Both processes can precisely and controllably construct microscopic biomimetic probe topologies with high aspect ratios in a room-temperature liquid environment, avoiding the high costs of traditional microfabrication processes and the potential damage to the underlying flexible coating.

[0078] Step S4: Immerse the electrode containing the conductive buffer layer and the pseudo-invasive microfilament layer in a buffer solution containing dopamine monomer or catechol group monomer, and form the tissue adhesion layer on the surface by self-polymerization deposition.

[0079] This step involves a surface chemical modification process under mild conditions. Dopamine or monomers containing catechol groups spontaneously oxidize to generate highly reactive quinone intermediates and undergo intermolecular polymerization and cross-linking in the presence of a specific pH buffer and dissolved oxygen. This self-polymerization deposition exhibits excellent "conformal coating" characteristics. The newly formed polymer film can uniformly and extremely thinly coat the surface of each arrayed nanowire and the gaps at the bottom, without filling or disrupting the original microscopic rough topology of the nanowire array. This process not only completely preserves the physical anchoring configuration of the intrusive microfilament layer but also introduces a high density of chemically adhering active groups on the outermost surface, thus achieving a complete closed loop of the biomimetic interface.

[0080] In some embodiments, in step S1, the surface of the electrode substrate is cleaned by sequentially performing ultrasonic cleaning in acetone, ethanol, and deionized water.

[0081] Acetone is used to efficiently dissolve and remove non-polar organic greases from the substrate surface, ethanol is used as a transitional polar solvent to remove residual organic matter, and deionized water is used to wash away water-soluble ionic impurities. Combined with the physical cavitation stripping effect of ultrasound, this sequential cleaning method can provide a defect-free, molecularly clean surface for subsequent plasma activation.

[0082] In some embodiments, in step S2, the electrodeposition is performed using cyclic voltammetric electrodeposition or constant potential electrodeposition.

[0083] In some embodiments, in step S2, the deposition potential of the electrodeposition is 0.5V to 1.2V.

[0084] In some embodiments, in step S2, the electrodeposition time is 5 to 30 minutes.

[0085] In step S2, when preparing the conductive buffer layer, the electrodeposition is performed using either cyclic voltammetric electrodeposition or potentiostatic electrodeposition. Cyclic voltammetric electrodeposition promotes dense stacking of the polymer network, while potentiostatic electrodeposition provides a continuous and efficient growth rate. The deposition potential is controlled within the range of 0.5V to 1.2V; for example, it can be 0.5V, 0.6V, 0.7V, 0.8V, 0.9V, 1.0V, 1.1V, 1.15V, 1.2V, etc. This potential range ensures sufficient oxidative polymerization of the conductive polymer monomers while effectively avoiding the occurrence of side reactions during water electrolysis. The deposition time is controlled within the range of 5 minutes to 30 minutes; for example, it can be 5 minutes, 8 minutes, 10 minutes, 15 minutes, 20 minutes, 25 minutes, 28 minutes, 30 minutes, etc. This time range, combined with the potential parameters, allows for precise control of the macroscopic thickness and Young's modulus of the buffer layer.

[0086] In some embodiments, in step S3, when using the template-assisted electrodeposition method, a porous alumina template or a polycarbonate track etching film is spin-coated onto the surface of the conductive buffer layer.

[0087] In some embodiments, in step S3, the pore size of the porous alumina template or the polycarbonate track etching film is 50 nm to 500 nm. In some embodiments, in step S3, the pore spacing of the porous alumina template or the polycarbonate track etching film is 100 nm to 1 μm; In some embodiments, in step S3, the deposition potential is 0.6V~1.0V; In some implementations, the deposition time in step S3 is 10 minutes to 60 minutes; In some embodiments, in step S3, when using the template-assisted electrodeposition method, the porous alumina template or the polycarbonate track etching film is removed with a solvent after deposition is completed.

[0088] Preferably, the method for removing the porous alumina template or the polycarbonate track etching film with a solvent is as follows: dissolving the porous alumina template with sodium hydroxide solution, or dissolving the polycarbonate track etching film with dichloromethane; In step S3, when constructing the simulated intrusion microfilament layer, if the template-assisted electrodeposition method is used, a porous alumina template or a polycarbonate track etching film is spin-coated or attached to the surface of the conductive buffer layer. The pore size range of the porous alumina template or the polycarbonate track etching film is set to 50 nm to 500 nm; for example, it can be 50 nm, 100 nm, 150 nm, 200 nm, 250 nm, 300 nm, 350 nm, 400 nm, 450 nm, 500 nm, etc. The pore spacing of the template is set to 100 nm to 1 μm; for example, it can be 100 nm, 200 nm, 300 nm, 400 nm, 500 nm, 600 nm, 700 nm, 800 nm, 900 nm, 1 μm, etc. The above pore parameters limit the diameter and distribution density of the microfilament array in physical space. Within this confined space, the deposition potential for secondary electrodeposition is set to 0.6V~1.0V; for example, it can be 0.6V, 0.65V, 0.7V, 0.75V, 0.8V, 0.85V, 0.9V, 0.95V, 1.0V, etc. The deposition time is set to 10 minutes~60 minutes; for example, it can be 10 minutes, 15 minutes, 20 minutes, 30 minutes, 40 minutes, 45 minutes, 50 minutes, 55 minutes, 60 minutes, etc.

[0089] After deposition, a highly selective demolding process is performed to remove the porous alumina template or the polycarbonate track etching film using a solvent; preferably, sodium hydroxide solution is used to dissolve the amphoteric porous alumina template, or dichloromethane is used to dissolve the organic polycarbonate track etching film. This highly selective chemical etching process can completely remove the template framework while perfectly preserving the morphology of the conductive polymer microfilament array.

[0090] In some embodiments, in step S4, the buffer solution is a weakly alkaline buffer solution with a pH value of 8.0 to 9.0.

[0091] In some embodiments, in step S4, the self-polymerization deposition is carried out at room temperature for 12 to 48 hours.

[0092] In step S4, the surface chemical modification stage, the buffer solution used to soak the semi-finished electrode is prepared as a weakly alkaline buffer solution with a pH value precisely controlled between 8.0 and 9.0; for example, it can be 8.0, 8.1, 8.2, 8.4, 8.5, 8.6, 8.8, 8.9, 9.0, etc. This weakly alkaline microenvironment is a necessary catalytic condition for triggering the oxidative crosslinking of catechol groups. In this environment, self-polymerization deposition takes place at room temperature for 12 to 48 hours; for example, it can be 12 hours, 16 hours, 20 hours, 24 hours, 30 hours, 36 hours, 40 hours, 44 hours, 48 ​​hours, etc. This mild and slow liquid-phase conformal coating process not only forms a highly active chemical adhesion layer on the surface of the microfilaments but also prevents the array pores from becoming clogged due to overly vigorous reactions.

[0093] To further demonstrate the broad industrial applicability and outstanding clinical application potential of the implantable bioelectrode with the dodder-inspired aspirator structure described in this application, an implantable medical device is also provided, including the implantable bioelectrode with the dodder-inspired aspirator structure as described in the foregoing embodiments.

[0094] The implantable medical device is an integrated medical product with the biomimetic bioelectrode as the front-end tissue-machine interface. Because the biomimetic electrode can provide rapid, non-invasive in-situ chemical adhesion, physical anchoring of the microscopic topology, and excellent mechanical gradient buffering in the extremely soft, moist, and continuously dynamically micro-moving physiological environment within the body (such as blood flow pulsation, respiratory fluctuations, and limb movements), it significantly overcomes the pain points of traditional rigid metal electrodes caused by mechanical mismatch, such as interface peeling, micro-movement noise, and chronic inflammation. This substantially extends the in vivo effective working life of the implantable medical device from the system's physical foundation and improves the stability of charge transport.

[0095] In some embodiments, the implantable medical device includes at least one of a neural recording device, a cardiac pacemaker, and a deep brain stimulation device.

[0096] When the implantable medical device is a neural recording device (e.g., a brain-computer interface device applied to the motor cortex), it is primarily used to capture high-frequency neural action potential signals with extremely weak amplitudes. The nanowire array on the electrode surface, with its large effective specific surface area, significantly reduces the interfacial contact impedance of the device in the characteristic frequency band (e.g., 1 kHz); simultaneously, the extremely stable "adhesion-intrusion-anchoring" biomimetic interface eliminates relative slippage between the electrodes and neural tissue, greatly reducing baseline drift caused by motion artifacts. This enables the neural recording device to continuously and stably acquire high signal-to-noise ratio, high-fidelity neural signals during long-term implantation.

[0097] When the implantable medical device is a cardiac pacemaker, its electrode terminals need to be fixed to the myocardial tissue and withstand the high-frequency, large-amplitude periodic contraction and tearing of the myocardium. The outermost catechol groups in the electrode interface coating provide immediate and strong chemical bonding in a moist environment such as blood, while the inner hydrogel composite material exerts extreme elastic energy absorption and mechanical attenuation. This synergistic mechanism ensures that the tip of the pacemaker will not physically displace even under intense cardiac pulsation, maintaining a tight and low-impedance electrical coupling, thereby effectively maintaining an extremely low pacing threshold and extending the long-term operating life of the internal power supply of the pacemaker.

[0098] When the implantable medical device is a deep brain stimulation (DBS) device, traditional implantation methods often lead to long-term foreign body rejection and thick, insulating scarring due to the extremely soft and fragile target brain tissue. The biomimetic interface used in the device of this application utilizes a smooth decrease in Young's modulus to transition to an ultra-low modulus (e.g., 10 kPa ~ 1 MPa), combined with nanoscale probe microfilaments, to achieve truly non-invasive tissue compliance. Histological results confirm that after long-term implantation in the deep brain region, this device significantly inhibits the overactivation of astrocytes and microglia, resulting in a substantial reduction in the thickness of the fibrous capsule surrounding the electrodes. This ensures that the therapeutic charge released by the deep brain stimulator can be injected into the neural target area long-term, efficiently, and without attenuation, maintaining a long-lasting neuromodulation therapeutic effect.

[0099] This application also provides an implantable medical device, including the implantable medical device as described above.

[0100] Preferably, the implantable medical device further includes a control module electrically connected to the implantable bioelectrode of the dodder-inspired suction device structure or the implantable medical device; more preferably, the control module includes at least one of a signal processing unit, a pulse generation unit, and a power supply unit.

[0101] Specifically, the control module coordinates the control of its internal functional units according to a specific clinical diagnosis and treatment matrix. When the control module includes a signal processing unit, the signal processing unit receives weak in vivo electrophysiological signals captured by the front-end electrodes through the electrical connection path, and performs low-noise pre-amplification, hardware filtering, and analog-to-digital conversion on these signals. Because the front-end bioelectrode uses a biomimetic interface coating with Young's modulus gradient transition and nanofiber minimally invasive anchoring, it eliminates interface slippage and motion artifact noise caused by physiological micro-movements, enabling the signal processing unit to stably capture pure original waveforms with ultra-high signal-to-noise ratio, avoiding misjudgment of pathological features due to signal distortion.

[0102] When the control module includes a pulse generation unit, the pulse generation unit is programmed to generate voltage or current stimulation pulses with specific waveforms, frequencies, and charge amounts according to a preset clinical stimulation protocol. Because the oriented polymer nanowire array on the surface of the front-end bioelectrode constructs an extremely high electrochemical specific surface area, the contact impedance at the electrode interface is extremely low, and the charge injection capability is significantly enhanced. In this case, the pulse generation unit only needs to output an extremely low discharge threshold and stimulation voltage to achieve efficient regulation of target tissues (such as nerve nuclei or myocardium). This fundamentally avoids the risk of electrical burns to biological tissues caused by traditional high discharge voltages, ensuring the safety of long-term in-situ stimulation.

[0103] When the control module includes a power supply unit, the power supply unit provides a continuous and stable power supply to the signal processing unit, pulse generation unit, and main control microprocessor. The power supply unit specifically includes a medical-grade disposable lithium battery, a rechargeable battery that can be wirelessly charged through the skin, or a passive electromagnetic induction coupling energy flow. In this embodiment, due to the excellent electrochemical performance of the front-end biomimetic interface coating, the stimulation power consumption and signal-level computing power overhead of the entire device are reduced, resulting in a significant reduction in the power decay rate of the power supply unit itself, thus extending the effective service period of the medical device within the body by several times.

[0104] The implantable medical device provided in this embodiment has diverse finished device forms depending on the specific combination of clinical functions it carries and the different target organs. Specifically, it may include, but is not limited to, the following complete devices: for example, a neuromodulation system / peripheral nerve stimulator for implantation around the vagus nerve or spinal cord parenchyma, which periodically releases modulating pulses via a control module for the treatment of refractory epilepsy, chronic pain, or depression; or a cardiac pacemaker / automated external defibrillator for implantation in the myocardium, which monitors electrocardiogram waveforms via a signal processing unit and utilizes a pulse generation unit in the event of malignant arrhythmias. Release pacing or defibrillation currents; brain-computer interface devices / motor cortex recording systems, for implantation in the motor cortex of rats, non-human primates, or humans, which use high-density signal processing units to collect single-cell discharge (Spike) signals with high fidelity over a long period of time, enabling precise thought control of external robotic arms or computer cursors; and deep brain stimulation devices (brain pacemakers), for inserting needle electrodes deep into target points such as the subthalamic nucleus of the brain, with the control module implanted subcutaneously in the chest, releasing continuous high-frequency stimulation through a pulse generation unit to achieve long-term suppression of clinical symptoms of Parkinson's disease, essential tremor, or dystonia.

[0105] The aforementioned specific implantable medical devices, by highly integrating high-strength wet chemical adhesion, minimally invasive physical interlocking anchoring, and smooth mechanical gradient buffering technology with active hardware modules, construct a complete technical closed loop from the microscopic polymer interface to the macroscopic clinical equipment. While ensuring the safety of in vivo physiological minimally invasive procedures, they also achieve extreme stability of electrical interaction during the long-term working life of the entire system.

[0106] The present invention will be further illustrated below with specific embodiments. However, it should be understood that these embodiments are merely for the purpose of more detailed illustration and should not be construed as limiting the present invention in any way.

[0107] It should be noted that in the following examples and comparative examples, all experimental data are expressed as mean ± standard deviation (n≥3); statistical analysis was performed using t-tests, and p<0.05 was considered to be statistically significant.

[0108] Example 1: Fabrication and muscle tissue implantation evaluation of biomimetic electrodes based on metal microfilament substrates 1. Experimental objective: This embodiment provides a three-layer composite biomimetic bioelectrode based on a platinum-iridium alloy microfilament substrate, and focuses on investigating the interfacial electrical stability, tissue adhesion and resistance to mechanical disturbance of the electrode after implantation in muscle tissue.

[0109] 2. Experimental methods (preparation process): (1) Substrate pretreatment: A platinum-iridium alloy microwire electrode with a diameter of 200 μm was used as the electrode substrate. The substrate was ultrasonically cleaned in acetone, ethanol and deionized water for 5 minutes in sequence, and then treated with oxygen plasma for 2 minutes to activate the surface.

[0110] (2) Preparation of conductive buffer layer: In an electrolyte containing 10 mmol / L 3,4-ethylenedioxythiophene (EDOT) and 0.1 mol / L polystyrene sulfonate (PSS), electrodeposition was performed at 0.8 V (relative to silver / silver chloride reference electrode) for 15 minutes using a constant potential method to form a PEDOT:PSS conductive buffer layer with a thickness of about 3 μm (measured Young's modulus of about 25 MPa) on the electrode surface.

[0111] (3) Preparation of the intrusive microfilament layer: A porous alumina template with a thickness of about 500 nm (pore size of about 200 nm) was spin-coated onto the surface of the above conductive buffer layer. It was placed in an electrolyte containing 0.1 mol / L pyrrole and 0.1 mol / L sodium p-toluenesulfonate and deposited at a constant potential of 0.7 V for 30 minutes. After deposition, the porous alumina template was dissolved in 1 mol / L sodium hydroxide solution to obtain a vertically oriented PPy (polypyrrole) nanowire array with a length of about 2 μm and a diameter of about 200 nm (effective Young's modulus of about 2 MPa).

[0112] (4) Preparation of tissue adhesion layer: The electrode with the microfilament layer was immersed in Tris-HCl buffer (pH 8.5) containing 2 mg / mL dopamine and self-polymerized for 24 hours at room temperature to form a polydopamine (PDA) adhesion layer with a thickness of about 50 nm (Young's modulus of about 500 kPa) on the surface of the microfilament, thus obtaining a biomimetic bioelectrode.

[0113] 3. Test methods and result analysis: (1) In vivo impedance and macroscopic observation: The prepared bionic electrode and the untreated bare platinum-iridium electrode (as a control) were implanted into the gastrocnemius muscle of rats. The contact impedance between the electrode and the tissue was measured regularly from day 1 to day 28 after implantation (test frequency 1 kHz). The test results showed that the impedance of the bionic electrode remained stable in the range of 15 kΩ to 22 kΩ within 28 days, with an overall fluctuation of less than 40%; while the impedance of the bare platinum electrode in the control group increased sharply from the initial 25 kΩ to more than 180 kΩ on day 28, with a fluctuation of more than 600%. After 28 days of implantation, the electrodes were removed for macroscopic observation. There was no obvious edema or fibrous encapsulation in the tissue around the bionic electrode, and the coating on the electrode surface remained intact; obvious fibrous capsules were visible around the bare platinum electrode.

[0114] (2) In vitro adhesion test: The biomimetic electrode was bonded to fresh ex vivo pig skin tissue using an overlap shear test. The shear adhesion strength of the biomimetic electrode with the three-layer composite coating in this embodiment was measured to be 45.3±5.2 kPa, and the peel strength (90° peel) was 12.4±1.8 N / m. In contrast, the comparative electrode coated with only a single layer of polydopamine (PDA) had a shear strength of only 18.6±3.1 kPa and a peel strength of only 5.6±0.9 N / m; the comparative electrode with only a single layer of nanowires had a shear strength of only 8.2±1.5 kPa and a peel strength of only 2.1±0.4 N / m. The data show that the three-layer composite structure produced a very strong synergistic adhesion effect.

[0115] (3) Ultrasonic oscillation disturbance test: The above-mentioned biomimetic electrode and the comparative electrode with only a PDA adhesion layer were placed in an ultrasonic water bath and oscillated for 1 hour (test conditions: power 100W, frequency 40kHz). After oscillation, the impedance change rate of the biomimetic electrode was only 4.2±1.3%, and SEM (scanning electron microscopy) observation showed that the nanowire array structure remained intact without any detachment. However, the impedance of the comparative electrode with only a PDA layer increased significantly by 235% after oscillation, and a large amount of surface coating peeled off.

[0116] (4) Interface SEM microscopic observation: refer to Figure 2 Twenty-eight days after implantation, the electrode-tissue composite was removed, and after fixation, dehydration, and embedding, a cross-sectional tissue sample was prepared for SEM observation. The images clearly showed that the tip of the PPy nanowire array extended and penetrated into the superficial layer of the gastrocnemius muscle tissue at a depth of approximately 500 nm to 800 nm. The nanowires were densely adhered to the tissue, and no obvious gaps or mechanical breakage of individual cells were observed.

[0117] Example 2: Fabrication and Evaluation of Bionic Electrodes Based on Silicon-Based Arrays for Neural Modulation Recording 1. Experimental Objective: This embodiment provides a biomimetic electrode based on a silicon-based neural microelectrode array. The main purpose is to investigate its performance in high-resolution neural signal recording of the cerebral cortex, as well as the long-term inhibitory effect of the biomimetic interface on tissue foreign body reaction and chronic inflammation.

[0118] 2. Experimental methods (preparation process): The preparation method in this embodiment is basically the same as that in Example 1, with the main difference being the parameters: (1) Substrate: A silicon-based neural microelectrode array (16-channel recording sites) was used as the electrode substrate.

[0119] (2) Invasion-like microfilament layer: PEDOT (poly(3,4-ethylenedioxythiophene)) nanowire material was used, the pore size of the porous alumina template was adjusted to 100 nm, the electrodeposition time was adjusted to 20 minutes, and the length of the prepared nanowire was about 1.5 μm.

[0120] 3. Test methods and result analysis: (1) In vivo neural signal recording: The prepared biomimetic silicon-based electrode and the untreated bare silicon-based electrode (as a control) were implanted into the motor cortex of mice. Spontaneous neural action potential signals were recorded on days 7, 14, and 28 post-implantation. The results showed that the biomimetic electrode recorded extremely clear action potential waveforms at all time points, with a signal-to-noise ratio consistently above 3. In contrast, the bare silicon electrode (as a control) showed a significant decrease in signal amplitude and a signal-to-noise ratio below 2 after day 14 post-implantation. Furthermore, due to micromotion, the recorded waveforms contained a large amount of motion artifacts. Further observation of brain tissue sections indicated that the morphology of nerve cells around the biomimetic electrode remained normal, and the electrode-tissue physical interface was tightly and seamlessly adhered without any obvious gaps.

[0121] (2) Immunofluorescence staining and assessment of inflammatory response: The experimental animals were sacrificed 4 weeks (28 days) after implantation, and brain tissue was extracted for frozen sectioning and immunofluorescence staining assessment of GFAP (astrocytic marker) and Iba-1 (microglia marker).

[0122] refer to Figure 3 Staining results showed that the thickness of the GFAP-positive region of activated astrocytes around the biomimetic electrode was only 20±5 μm, and the density of Iba-1-positive microglia was 125±35 cells / mm². 2 The final gelatinous scar (fibrous capsule) is only 15±4μm thick.

[0123] In contrast, the thickness of the GFAP-positive region around a conventional bare silicon electrode is as high as 120±25 μm, and the Iba-1-positive cell density is as high as 520±80 cells / mm². 2 The thickness of the gelatinous scar reached 85±12μm. These detailed data strongly demonstrate that the gradient biomimetic interface of this invention can significantly inhibit foreign body rejection and chronic inflammation induced by microelectrode implantation.

[0124] Example 3: Biomimetic electrodes containing interpenetrating hydrogels and metal-polyphenol adhesive layers for cardiac pacing assessment 1. Experimental objective: In this embodiment, a biomimetic electrode with a PVA interpenetrating hydrogel buffer layer, a template-free self-assembled microfilament layer, and a metal-polyphenol coordination adhesive layer was prepared to investigate its in vivo signal stability and resistance to cyclic stretching as an epicardial electrode under high-frequency dynamic deformation environment.

[0125] 2. Experimental methods (preparation process): Substrate pretreatment: A gold disk electrode was used as the electrode substrate. The gold disk electrode substrate was ultrasonically cleaned in acetone, ethanol and deionized water for 5 minutes in sequence, and then treated with oxygen plasma for 2 minutes to activate the surface.

[0126] (1) Conductive buffer layer: PEDOT:PSS dispersion and 5wt% polyvinyl alcohol (PVA) solution were uniformly mixed at a volume ratio of 1:1. The mixture was electrodeposited onto the surface of a gold disk using cyclic voltammetry (scanning potential range of -0.5V to 0.8V, scan rate of 50mV / s, for a total of 20 cycles) to form a PEDOT:PSS / PVA interpenetrating hydrogel composite material.

[0127] (2) Invasive microfilament layer: A template-free self-assembly method was used. PEDOT nanowire arrays with self-oriented structures were directly induced to grow on the electrode surface by deposition at a constant potential of 0.9V for 25 minutes in an electrolyte containing 10 mmol / L LEDOT monomer and 0.1 mmol / L sodium dodecyl sulfate (surfactant).

[0128] (3) Tissue adhesion layer: Tannic acid coordination complex is used instead of dopamine. The electrodes with the first two layers prepared are immersed in phosphate buffer containing 1 mg / mL tannic acid and 1 mg / mL ferric chloride (FeCl3) and reacted at room temperature for 6 hours. Through in-situ coordination crosslinking of metal-polyphenol, an extremely thin adhesion coating is formed on the surface.

[0129] Test methods and results analysis: (4) In vivo electrocardiogram (ECG) recording: The prepared biomimetic electrode was used as an epicardial electrode and directly implanted and attached to the surface of the beating heart of a rat for in vivo ECG signal recording. The results showed that the QRS complex amplitude captured by the biomimetic electrode was extremely stable, and the attenuation rate of the signal characteristic amplitude was less than 15% after 4 hours of continuous high-frequency heartbeat recording. In contrast, the bare gold disk electrode, under the same dynamic conditions, showed a sharp decrease in the amplitude of the captured signal of more than 50% after only 30 minutes of recording, accompanied by severe baseline drift.

[0130] (5) In vitro simulated micro-dynamic cyclic tensile test: The above-mentioned biomimetic electrode and the bare gold electrode of the comparative example were fixed together on a mechanical tensile test platform. The test conditions were set as follows: 10% physical tensile strain, 1Hz cyclic frequency, and 100,000 high-intensity cyclic tensile tests were continuously performed. After 100,000 cycles, the impedance change rate of the biomimetic electrode at a frequency of 1kHz was only 12.3±3.5%. Electrochemical impedance spectroscopy analysis further showed that its interfacial double layer capacitance decreased by only 8%, and the charge injection capability remained stable at more than 90% of the initial design value. In contrast, the impedance change rate of the bare gold electrode soared to 342±56%. This fully demonstrates that the composite buffer layer of the present invention can perfectly absorb and dissipate extreme dynamic cyclic stress.

[0131] Comparative Example 4 The biomimetic electrode was prepared using essentially the same method as in Example 3, except that the Young's modulus of the three coating layers was set to a reverse gradient (15 MPa for the outermost layer, 8 MPa for the middle layer, and 1 MPa for the innermost layer). After undergoing 100,000 cyclic tensile tests using the testing method of Example 3, the interfacial impedance change rate of this reverse gradient electrode reached as high as 198 ± 35%.

[0132] Comparative Example 5 The Young's modulus of the three coating layers was set to a uniform gradient (3~5 MPa for each layer). After the same test, the impedance change rate reached 103±18%. The above results confirm that only when the biomimetic interface coating exhibits a gradient transition of decreasing Young's modulus from the inside to the outside can the interface stress concentration be effectively eliminated and the long-term mechanical stability and impedance stability of the coating be guaranteed. Any design that deviates from this gradient direction cannot achieve the expected technical effect of this invention.

[0133] The test results of Comparative Examples 4 and 5 above show that only when the biomimetic interface coating exhibits a gradient transition with decreasing Young's modulus from the inside out can the interface stress concentration be effectively avoided, ensuring the physical integrity of the coating under long-term mechanical disturbance in the body, thereby achieving the best anti-peeling and electrical contact stability.

[0134] Example 4: Fabrication and performance evaluation of biomimetic electrodes based on flexible polymer thin film substrates 1. Experimental Objective: This embodiment provides a flexible biomimetic electrode based on an ultra-soft organic polymer substrate. The main purpose is to verify the compatibility of the multi-level modulus gradient coating process of this invention with advanced flexible thin film substrates, as well as the mechanical reliability of the coating under extreme deformation.

[0135] 2. Experimental methods (preparation process): (1) Substrate treatment: A 100 μm thick polydimethylsiloxane (PDMS) flexible elastomer film was used as the electrode substrate, and a 50 nm thick gold film was pre-sputtered on its surface as the basic conductive layer. The flexible PDMS gold electrode was ultrasonically cleaned in isopropanol and deionized water for 3 minutes in sequence, and then activated by oxygen plasma at 30 W power for 1 minute.

[0136] (2) Preparation of composite coating: The method of Example 1 is basically followed, and three biomimetic structures are constructed sequentially on the flexible surface: 1) Deposition of PEDOT:PSS conductive buffer layer (electrodeposition for 10 minutes, film thickness approximately 2 μm, Young's modulus approximately 20 MPa); 2) Deposition of PPy nanowire array microfilament layer (using an alumina template with a pore size of 150 nm, electrodeposition for 25 minutes, microfilament length of about 1.8 μm, Young's modulus of about 1.5 MPa); 3) Self-polymerized polydopamine (PDA) adhesion layer (film thickness approximately 40 nm, Young's modulus approximately 400 kPa).

[0137] 3. Test methods and result analysis: (1) In vivo impedance and tissue compatibility assessment: The flexible biomimetic electrode and the uncoated PDMS bare gold electrode (as a control) were implanted into the gastrocnemius muscle of rats. After an implantation period of up to 28 days, the interfacial impedance of the flexible biomimetic electrode remained stable in the range of 12kΩ to 18kΩ, with a long-term impedance fluctuation of less than 35%. In contrast, the PDMS bare gold electrode in the control group, lacking tissue adhesion and buffering anti-peeling mechanisms, showed a rapid increase in contact impedance from the initial 30kΩ to over 150kΩ. Tissue sections extracted from the implantation site confirmed that, due to the dual compliance design of the flexible substrate combined with the biomimetic gradient, no obvious fibrous wrapping layer (capsule) was found around the flexible biomimetic electrode, and the physical morphology of the electrode achieved a near-perfect seamless fit with the tissue.

[0138] (2) Extreme mechanical bending test: The prepared flexible biomimetic electrode was subjected to extreme mechanical strain tests. The results showed that when the electrode was subjected to a severe bending strain of up to 50%, the composite biomimetic coating on its surface did not exhibit any macroscopic or microscopic cracking, shrinkage, or interlayer delamination, and the core electrical conductivity did not show any significant attenuation. This data proves that the coating technology of this invention has unparalleled process compatibility and mechanical robustness with high-deformation flexible polymer substrates.

[0139] Example 5 This embodiment provides a biomimetic electrode based on a PANI / gelatin conductive buffer layer, a PANI nanowire microfilament layer, and a dopamine-grafted chitosan adhesive layer.

[0140] Following the preparation method of Example 1, only the EDOT monomer in the conductive buffer layer was replaced with aniline monomer (0.1 mol / L), the hydrophilic polymer was replaced with gelatin (5 wt%), the electrodeposition potential was adjusted to 0.9 V, and the deposition time was 20 minutes; the intrusive microfilament layer used PANI nanowires (template pore size 200 nm, electrodeposition 0.8 V, 30 minutes); and the tissue adhesion layer used dopamine-grafted chitosan (2 mg / mL, pH 8.5, self-polymerization for 24 hours). In vitro impedance testing showed that the electrode's initial impedance at 1 kHz was approximately 18 kΩ, comparable to 15 kΩ in Example 1; the overlap shear adhesion strength was 42.3 ± 4.8 kPa, not significantly different from 45.3 ± 5.2 kPa in Example 1. This indicates that comparable interfacial stability and anti-inflammatory effects as in Example 1 can be achieved in vivo.

[0141] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention, and not to limit them. Although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art should understand that modifications can still be made to the technical solutions described in the foregoing embodiments, or equivalent substitutions can be made to some or all of the technical features therein. Such modifications or substitutions do not cause the essence of the corresponding technical solutions to deviate from the scope of the technical solutions of the embodiments of the present invention.

Claims

1. An implantable bioelectrode with a structure inspired by dodder seed siphons, characterized in that, Includes an electrode substrate and a biomimetic interface coating disposed on the surface of the electrode substrate; The biomimetic interface coating comprises, from the inside out, a conductive buffer layer, an intrusive microfilament layer, and a tissue adhesion layer. The conductive buffer layer is a conductive polymer hydrogel composite material composed of an interpenetrating conductive polymer and a hydrophilic polymer network. The simulated intrusion microfilament layer is an array of oriented conductive polymer nanowires; The tissue adhesion layer is a polymer film containing catechol groups.

2. The implantable bioelectrode with a dodder-inspired siphon structure as described in claim 1, characterized in that, The conductive polymer is selected from at least one of poly(3,4-ethylenedioxythiophene)-polystyrene sulfonate, polypyrrole, and polyaniline; and / or, The hydrophilic polymer includes at least one of polyvinyl alcohol, polyacrylic acid, hyaluronic acid, sodium alginate, and gelatin; and / or, The conductive polymer nanowire array is made of at least one of poly(3,4-ethylenedioxythiophene), polypyrrole, and poly(3,4-ethylenedioxythiophene)-tetracyanoquinone dimethane; and / or The material of the tissue adhesion layer is selected from at least one of polydopamine, dopamine-methacrylate copolymer, dopamine-grafted chitosan, tannic acid coordination complex, and epigallocatechin gallate coordination complex.

3. The implantable bioelectrode with a dodder-inspired siphon structure as described in claim 1, characterized in that, The total thickness of the biomimetic interface coating is 2μm~20μm; and / or, The thickness of the conductive buffer layer is 1μm~10μm; and / or, The length of the nanowires in the pseudo-invasive microfilament layer is 500 nm to 5 μm; and / or, The diameter of the nanowires in the simulated intrusion microfilament layer is 50 nm to 500 nm; and / or, The thickness of the tissue adhesion layer is 10 nm to 200 nm.

4. The implantable bioelectrode with a dodder-inspired siphon structure as described in claim 1, characterized in that, Along the direction from the conductive buffer layer to the tissue adhesion layer, the Young's modulus of the biomimetic interface layer exhibits a decreasing gradient transition; and / or, The Young's modulus of the conductive buffer layer is 1 MPa to 100 MPa; preferably, the Young's modulus of the conductive buffer layer is 10 MPa to 50 MPa; and / or, The Young's modulus of the simulated invasive microfilament layer is 100 kPa to 10 MPa; and / or, The Young's modulus of the tissue adhesion layer is 10 kPa to 1 MPa.

5. The implantable bioelectrode with a dodder-inspired siphon structure as described in claim 1, characterized in that, The electrode substrate includes at least one of a metal substrate, a silicon-based microelectrode array, a carbon-based electrode, and a flexible polymer substrate; Preferably, the metal substrate comprises at least one selected from gold, platinum, iridium, and stainless steel; Preferably, the flexible polymer substrate includes at least one of polyimide, polydimethylsiloxane, and phenelzine.

6. A method for preparing an implantable bioelectrode with a dodder-inspired siphon structure as described in any one of claims 1-5, characterized in that, include: S1, the surface of the electrode substrate is cleaned and treated with oxygen plasma; S2, The conductive polymer is mixed with the hydrophilic polymer and deposited on the surface of the treated electrode substrate by electrodeposition to form the conductive buffer layer; S3, using template-assisted electrodeposition or template-free self-assembly, the conductive polymer nanowire array is deposited or grown on the surface of the conductive buffer layer to form the intrusion-resistant microfilament layer; S4, the electrode containing the conductive buffer layer and the pseudo-invasive microfilament layer is immersed in a buffer solution containing dopamine monomer or catechol group monomer, and the tissue adhesion layer is formed on the surface by self-polymerization deposition.

7. The method for preparing the implantable bioelectrode with the dodder-inspired siphon structure as described in claim 6, characterized in that, In step S1, the surface of the electrode substrate is cleaned, including ultrasonic cleaning in acetone, ethanol, and deionized water in sequence; and / or, In step S2, the electrodeposition is performed using cyclic voltammetric electrodeposition or constant potential electrodeposition; and / or, In step S2, the deposition potential of the electrodeposition is 0.5V~1.2V; and / or, In step S2, the electrodeposition time is 5 minutes to 30 minutes; and / or, In step S3, when using the template-assisted electrodeposition method, a porous alumina template or a polycarbonate track etching film is spin-coated onto the surface of the conductive buffer layer. And / or, In step S3, the pore size of the porous alumina template or the polycarbonate track etching film is 50 nm to 500 nm; and / or, In step S3, the pore spacing of the porous alumina template or the polycarbonate track etching film is 100 nm to 1 μm; and / or, In step S3, the deposition potential is 0.6V~1.0V; and / or, In step S3, the deposition time is 10 minutes to 60 minutes; and / or, In step S3, when using the template-assisted electrodeposition method, after deposition is completed, the porous alumina template or the polycarbonate track etching film is removed with a solvent; preferably, the method of removing the porous alumina template or the polycarbonate track etching film with a solvent is: dissolving the porous alumina template with sodium hydroxide solution, or dissolving the polycarbonate track etching film with dichloromethane. And / or, In step S4, the buffer solution is a weakly alkaline buffer solution with a pH value of 8.0~9.0; and / or, In step S4, self-polymerization deposition is carried out at room temperature for 12 to 48 hours.

8. An implantable medical device, characterized in that, The implantable bioelectrode includes the dodder-inspired aspirator structure as described in any one of claims 1-5.

9. The implantable medical device as described in claim 8, characterized in that, The implantable medical device includes at least one of a neural recording device, a cardiac pacing device, and a deep brain stimulation device.

10. An implantable medical device, characterized in that, Including the implantable medical device as described in claim 8 or 9; Preferably, the implantable medical device further includes a control module electrically connected to the implantable bioelectrode of the dodder-inspired suction device structure or the implantable medical device; more preferably, the control module includes at least one of a signal processing unit, a pulse generation unit, and a power supply unit.

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