Bio-electronic interface system with interface function partition and application of bio-electronic interface system
By designing the interface with functional partitions, a signal coupling zone for the target tissue and an interference suppression zone for non-target tissue are constructed. This solves the problems of signal stability and interference suppression of bioelectronic interfaces in dynamic and humid environments, and achieves high signal-to-noise ratio and long-term stable bioelectric signal acquisition.
Patent Information
- Authority / Receiving Office
- CN · China
- Patent Type
- Applications(China)
- Current Assignee / Owner
- JINAN UNIVERSITY
- Filing Date
- 2026-01-21
- Publication Date
- 2026-05-12
AI Technical Summary
Existing bioelectronic interfaces struggle to achieve stable coupling of target tissue signals and effective suppression of non-target interface disturbances in dynamic, humid physiological environments, resulting in low signal-to-noise ratios and unstable long-term monitoring.
By designing functional partitions on the interface, a signal coupling zone for the target tissue and an interference suppression zone for non-target tissue are constructed. By utilizing the differences in the composition, configuration, and contact method of the interface layer, a stable, low-impedance electrode-tissue coupling interface is formed, and the mechanical disturbance of non-target tissue is reduced.
It significantly improves the signal-to-noise ratio and long-term monitoring stability of bioelectric signals, with a signal-to-noise ratio increase of approximately 2.15 times. It also reduces baseline drift and background noise, making it suitable for bioelectric signal acquisition in complex physiological environments.
Smart Images
Figure CN122004871A_ABST
Abstract
Description
Technical Field
[0001] This invention relates to the field of bioelectronics and flexible electronic devices, specifically a bioelectronic interface system with functional partitioning of the interface and its application. Background Technology
[0002] Bioelectrical signals (such as electromyography and electrocardiography) are important information carriers reflecting the electrical activity state of biological tissues and have wide applications in sports health monitoring, rehabilitation assessment, and clinical auxiliary diagnosis. As a key bridge between biological tissues and signal acquisition circuits, the performance of bioelectronic interfaces directly determines the stability, signal-to-noise ratio, and long-term monitoring reliability during signal acquisition.
[0003] In existing technologies, surface electrode systems typically rely on a metal electrode to form a tissue-electrode coupling interface with a conductive gel or flexible interface layer (e.g., US3998215A). However, there is a significant mechanical mismatch between traditional rigid electrodes and soft, physiologically dynamic tissues, which easily leads to fluctuations in contact state during movement or in bodily fluid environments, resulting in problems such as motion artifacts, baseline drift, and signal attenuation (e.g., KR20140144173A). To improve interface adhesion, existing technologies have introduced flexible interface materials such as hydrogels for bioelectrical signal acquisition. However, these solutions are mostly designed based on the premise of overall uniform interface characteristics. When the interface simultaneously contacts the target tissue and surrounding non-target tissues, external disturbances can still couple into the signal acquisition area through the interface layer, affecting signal stability (e.g., US20140206976A1).
[0004] To address the issues of gel interfaces easily absorbing water and swelling, and the degradation of adhesion performance in humid environments, existing patents have proposed obtaining differentiated adhesion behavior through asymmetric interface regulation at the material level. For example, a method for preparing asymmetric adhesive hydrogels using a reversal regulation strategy has been disclosed, and its application to surface electromyography signal acquisition has been shown. This type of technology focuses on the preparation method of hydrogel materials and the regulation of interface chemistry, and improves the moisture resistance and adhesion stability by adjusting the surface composition on both sides of the gel (e.g., CN118599052A).
[0005] However, the aforementioned existing technologies primarily focus on improving the interface material itself or its preparation process, with their starting point still centered on "obtaining a superior asymmetric gel material." In practical applications, bioelectronic interfaces are often not composed of a single interface material, but rather a multi-layered system structure encompassing conductive pathways, encapsulation layers, flexible support structures, and tissue contact interfaces. In this system, different interface regions perform different functions in terms of mechanical response, wetting behavior, and signal coupling paths. Relying solely on the asymmetric design of the material itself is insufficient to address the coexistence of stable signal coupling to the target tissue and suppression of interference from non-target tissues at the system level.
[0006] Furthermore, under dynamic motion conditions, traction, friction, and fluid disturbances generated by non-target tissues are often transmitted to the signal acquisition area through the overall interface structure, forming background noise that cannot be ignored. Existing technical solutions centered on material preparation have not yet systematically designed the synergistic relationships between the various functional interfaces within the bioelectronic interface, making it difficult to achieve directional control of the signal transmission path and effective isolation of interference sources.
[0007] Therefore, it is necessary to move beyond simply optimizing interface performance from the perspective of material preparation and instead construct a bioelectronic interface scheme that can achieve stable coupling of target tissue signals, suppression of non-target interface disturbances, and high-fidelity transmission of electrical signals within the same interface, based on the design of bioelectronic interface system structure and interface functional zoning, in order to meet the practical needs of long-term and stable acquisition of bioelectrical signals under complex physiological environments.
[0008] The information disclosed in this background section is intended only to enhance the understanding of the overall background of the invention and should not be construed as an admission or in any way implying that the information constitutes prior art known to those skilled in the art. Summary of the Invention
[0009] The technical problem to be solved by this invention is to overcome the above-mentioned technical defects and provide a bioelectronic interface system with functional partitioning of the interface and its application. This bioelectronic interface can achieve stable coupling of target tissue signals in a dynamic and humid physiological environment, while effectively suppressing mechanical disturbances and noise interference introduced by non-target interfaces, thereby significantly improving the signal-to-noise ratio and long-term monitoring stability of bioelectrophysiological signals. The main feature of the invention is that, through interface structure design rather than simply relying on the uniform performance regulation of interface materials, the bioelectronic interface can simultaneously possess strong signal coupling capability and low interference characteristics in the same device, overcoming the shortcomings of existing two-sided hydrophilic interface bioelectronic interfaces and commercial gel electrodes in terms of motion artifacts, swelling instability and signal attenuation.
[0010] To address the above problems, the technical solution of this invention is a bioelectronic interface system with functional partitioning of the interface. The preparation method of the bioelectronic interface system includes the following steps:
[0011] Step 1: Preparation of flexible conductive pathways: Liquid metal is mixed with a polymer stabilizer and subjected to ultrasonic treatment under water bath conditions to obtain conductive ink with good dispersibility; the conductive ink is patterned and deposited on the surface of a flexible polyurethane (PU) film using direct writing printing. After the solvent evaporates, mechanical treatment is used to form a continuous conductive network inside the conductive pathway, thereby obtaining a stretchable and bendable flexible conductive pathway.
[0012] Step 2, Flexible encapsulation of conductive paths: The conductive paths obtained in Step 1 are bonded or covered with a flexible encapsulation layer, so that the conductive paths are encapsulated inside the flexible structure, thereby improving their mechanical and electrical stability in tensile, bending and humid environments.
[0013] Step 3: Construction of the Interface Functional Zoning Flexible Interface Layer: A flexible interface layer is constructed on one side of the conductive path, and at least two functionally distinct interface regions are formed in this interface layer. The interface regions include a target tissue signal coupling region and a non-target tissue interference suppression region. The target tissue signal coupling region is used to form a stable, low-impedance electrode-tissue coupling interface with the target biological tissue. The non-target tissue interference suppression region is used to reduce the impact of adhesion, friction, and mechanical disturbances generated during contact with non-target tissues on signal acquisition.
[0014] Step 4: Integration of the overall bioelectronic interface system: The structures built in Steps 2 and 3 are integrated to form an integrated bioelectronic interface system, which enables the acquisition, transmission and interface stabilization of bioelectric signals.
[0015] Preferably, in step one, the liquid metal is gallium-based or a gallium-indium alloy with a gallium-indium molar ratio of 1:1 to 1:5; and the ultrasonic power is 200 to 800 W.
[0016] Preferably, in step three, the interface functional partitioning is achieved through differences in interface layer composition, interface configuration, or interface contact methods, rather than relying on uniform changes in the overall performance of a single interface material.
[0017] The advantages of this invention compared to existing technologies are:
[0018] 1. This invention provides a method for achieving interface functional partitioning through interface structure design, so that the bioelectronic interface forms a stable, low-impedance electrode-tissue coupling interface when in contact with the target tissue, while exhibiting a weaker interface interaction capability when in contact with non-target interfaces, thereby reducing mechanical interference introduced by non-target interfaces.
[0019] 2. This invention provides a bioelectronic interface structure that integrates a flexible conductive pathway with a functionally partitioned interface layer, enabling the conductive pathway to maintain stable conductivity in a dynamic physiological environment, while ensuring the structural stability and adhesion reliability of the interface layer under humid conditions.
[0020] 3. This invention provides an application of a bioelectronic interface based on the above-mentioned interface functional partition design in bioelectrophysiological signal acquisition. It can achieve high signal-to-noise ratio and low artifact bioelectrophysiological signal acquisition without complex material modification or additional adhesion layers, providing a new technical approach for bioelectrophysiological monitoring and electronic skin applications. Attached Figure Description
[0021] Figure 1 This is a schematic diagram of the multilayer structure of the bioelectronic interface in Example 1 and its overall preparation process.
[0022] Figure 2 This is a schematic diagram illustrating the principle of the bioelectronic interface formation process in Example 1;
[0023] Figure 3 This is a comparison image of scanning electron microscopy (SEM) showing the changes in surface microstructure of the conductive pathways before and after mechanical activation in Example 1.
[0024] Figure 4 This is a comparison diagram of the interface response behavior of the bioelectronic interface in Example 1 under the target tissue contact state and the non-target interface contact state;
[0025] Figure 5 The figure shows the structural and performance stability test results of the bioelectronic interface in Example 1 under humid environmental conditions;
[0026] Figure 6 This is a schematic diagram of the bio-electromyography signal waveforms collected under in vivo experimental conditions using the bioelectronic interface in Example 1, along with the corresponding experimental data.
[0027] Figure 7 This is a comparison chart of the signal-to-noise ratio of bioelectromyography signals acquired by the bioelectronic interface under different interface design conditions in Example 1; Detailed Implementation
[0028] To make the content of this invention easier to understand, the technical solutions in the embodiments of this invention will be clearly and completely described below with reference to the accompanying drawings.
[0029] like Figure 1-7 As shown, this invention proposes a bioelectronic interface system with functional partitioning of the interface. By constructing a target tissue signal coupling interface and a non-target tissue interference suppression interface within the same flexible bioelectronic interface, selective acquisition and interference isolation of bioelectrical signals are achieved. The construction method of the bioelectronic interface system includes the following steps:
[0030] Step 1: Preparation of Flexible Conductive Pathways: Liquid metal is mixed with a polymer stabilizer and subjected to ultrasonic treatment under water bath conditions to obtain conductive ink with good dispersibility. Subsequently, the conductive ink is patterned and deposited onto the surface of a flexible polyurethane (PU) film using direct-write printing. After the solvent evaporates, mechanical processing is performed to form a continuous conductive network within the conductive path, thereby obtaining a stretchable and bendable flexible conductive path. The liquid metal is gallium-based or a gallium-indium alloy with a gallium-indium molar ratio of 1:1 to 1:5; the ultrasonic power is 200–800 W.
[0031] Step 2, Flexible encapsulation of conductive pathways: The conductive pathways obtained in step (1) are bonded to or covered with a flexible encapsulation layer so that the conductive pathways are encapsulated inside the flexible structure to improve their mechanical and electrical stability in tensile, bending and humid environments.
[0032] Step 3: Construction of the Interface Functional Zoning Flexible Interface Layer: A flexible interface layer is constructed on one side of the conductive path, forming at least two functionally distinct interface regions within this layer. These include a target tissue signal coupling region, used to form a stable, low-impedance electrode-tissue coupling interface with the target biological tissue; and a non-target tissue interference suppression region, used to reduce the impact of adhesion, friction, and mechanical disturbances generated during contact with non-target tissues on signal acquisition. The interface functional zoning is achieved through differences in interface layer composition, interface configuration, or interface contact methods, rather than relying on uniform changes in the overall performance of a single interface material.
[0033] Step 4: Integration of the overall bioelectronic interface system: The structures built in Steps 2 and 3 are integrated to form an integrated bioelectronic interface system, which enables the acquisition, transmission and interface stabilization of bioelectric signals.
[0034] In the aforementioned bioelectronic interface system, the target tissue signal coupling region and the non-target tissue interference suppression region differ in their interfacial wettability and interfacial interaction mechanisms. The target tissue signal coupling region is rich in polar groups that can generate non-covalent interactions with functional groups on the surface of biological tissues, enabling stable and conformal adhesion and electrode coupling at the tissue interface. In contrast, the non-target tissue interference suppression region reduces the effective interfacial contact or interfacial adhesion strength, thereby reducing mechanical disturbances introduced by non-target tissue traction and friction, thus achieving interference isolation at the interfacial level.
[0035] Experimental results show that, in electromyography (EMG) signal acquisition tests, the interface functional partition type bioelectronic interface described in this invention can maintain a stable interface state and clear signal output even under dynamic and humid environments. Compared with a homogeneous hydrophilic flexible interface, its EMG signal signal-to-noise ratio is improved by approximately 2.15 times; compared with commercial gel electrodes, its signal-to-noise ratio is improved by approximately 2.04 times. Furthermore, due to the effective suppression of interference introduced by non-target tissues by the interface functional partition design, the bioelectronic interface of this invention exhibits minimal baseline drift and excellent signal stability during long-term continuous monitoring, making it suitable for persistent bioelectrophysiological signal monitoring.
[0036] This invention achieves stable acquisition of target tissue signals and effective isolation of interference from non-target tissues through interface functional partitioning design at the bioelectronic interface system level, without relying on complex material preparation or reverse modulation processes. Compared with traditional double-sided uniform interfaces and commercial gel electrodes, this invention has significant advantages in resistance to moisture interference, swelling, and long-term stable monitoring, and has broad application prospects in medical and health monitoring, rehabilitation assessment, and human-computer interaction.
[0037] To verify the feasibility and effectiveness of this bioelectronic interface in actual electrophysiological signal acquisition, the signal-to-noise ratio (SNR) is used as a measure of the acquisition quality of electrophysiological signals. The formula for calculating SNR is as follows:
[0038]
[0039] Where Vsignal and Vnoise represent the voltage values of the signal and noise, respectively, and N represents the total number of samples.
[0040] Example 1
[0041] Fabrication of an asymmetric interface bioelectronic interface and its application in bioelectrophysiological signal acquisition:
[0042] Step (1): The surfactant hexadecyltrimethylammonium bromide (CTAB, 9 g, 24 mmol) and the flux sodium chloride (0.94 g, 16 mmol) were added to 20 mL of deionized water and stirred continuously at 75 °C until completely dissolved. Then, the hydrophobic monomer octadecyl acrylate (1.02 g, 3 mmol) was added to the resulting solution and stirred for 1 h to obtain a transparent and homogeneous surfactant-hydrophobic monomer micelle solution.
[0043] Step (2): Dimethylaminoethyl methacrylate (1.05 g, 6.7 mmol), hydroxypropyl acrylate (0.30 g, 2.2 mmol), crosslinking agent N,N-methylenebisacrylamide (3 mg, 0.019 mmol), and nano-silica (0.026 g, 0.43 mmol) were added sequentially to a 5 mL sample vial. Then, a solution of ferric chloride hexahydrate (0.08 g, 0.29 mmol) dissolved in 0.5 mL of water was added, and the mixture was thoroughly mixed until the system was homogeneous. Finally, 1 mL of the micelle solution was added to the system, and the mixture was stirred thoroughly to obtain gel precursor solution A.
[0044] Step (3): Add polyvinylpyrrolidone (PVP, 2 g, 17.9 mmol) to 20 mL of anhydrous ethanol and stir at room temperature to obtain an adhesive solution for stable dispersion of liquid metal.
[0045] Step (4): Add the gallium-indium liquid metal (75% gallium content, 25% indium content) and the adhesive solution (100 μL) obtained in step (3) together into a 5 mL centrifuge tube. Disperse the liquid metal using a probe ultrasonic method under water bath conditions. The ultrasonic power is set to 330 W and the processing time is 2 min to obtain liquid metal conductive ink.
[0046] Step (5): The conductive ink is printed onto the surface of a polyurethane (PU) film using a direct-write printing method to form a conductive circuit of a preset shape. After the anhydrous ethanol has completely evaporated, a brush is used to apply mechanical action to the circuit to achieve mechanical sintering, so that the liquid metal microstructures are interconnected to form a continuous conductive path, and the conductive circuit is further encapsulated using a PU film.
[0047] Step (6): Place the encapsulated conductive circuit in a hydrophilic glass mold. Dissolve ammonium persulfate (0.05 g, 0.2 mmol) in 0.5 mL of deionized water to prepare an initiator solution, and pour the gel precursor solution A with the initiator added into the glass mold for in-situ polymerization. After standing for 5 min, an integral asymmetric bioelectronic interface with hydrophobic weak adhesion properties on the bottom surface and hydrophilic strong adhesion properties on the top surface is obtained.
[0048] Step (7): Connect the obtained asymmetric bioelectronic interface to the bioelectrophysiological signal monitoring device for the acquisition and testing of bioelectrophysiological signals.
[0049] Comparative Example 1
[0050] Two-sided hydrophilic interface bioelectronic interface
[0051] Except for the interface forming conditions, the remaining steps are exactly the same as in Example 1.
[0052] Specifically, repeat steps (1) to (5) of Example 1 to prepare and encapsulate the conductive circuit.
[0053] Step (6): Place the encapsulated conductive circuit in a polytetrafluoroethylene (PTFE) mold, and perform in-situ polymerization using the same initiator solution and gel precursor solution A as in Example 1. Let it stand for 5 minutes to obtain an integral two-sided hydrophilic bioelectronic interface with strong hydrophilic adhesion on both the top and bottom surfaces.
[0054] Step (7): Connect the obtained two hydrophilic bioelectronic interfaces to the monitoring device to collect bioelectrophysiological signals.
[0055] In this comparative example, the gel interface of the bioelectronic interface has hydrophilic and strongly adhesive interfaces on both sides. Compared with Example 1, it did not form an asymmetric interface structure during the gelation process. This is used to compare and illustrate the effect of asymmetric interface adhesion characteristics in the process of bioelectrophysiological signal acquisition.
[0056] Comparative Example 2
[0057] Commercial Gel Interface Bioelectronic Interface
[0058] Step (1): Commercially available hydrogel electrodes are selected as interface materials. The commercial hydrogels are obtained directly through formal commercial channels. The specific formula and preparation process are provided by the manufacturer. This invention does not involve the preparation or modification process of the commercial hydrogels.
[0059] Step (2): Using the flexible conductive circuit prepared and encapsulated in Example 1 as an electronic pathway structure, the conductive circuit is directly sandwiched between the two layers of commercial hydrogel, so that the commercial hydrogel is located on the upper and lower surfaces of the conductive circuit, forming a bioelectronic interface structure with commercial hydrogel as the interface layer.
[0060] Step (3): Connect the commercial gel interface bioelectronic interface to the bioelectrophysiological signal monitoring device and collect bioelectrophysiological signals under the same test conditions as in Example 1.
[0061] In this comparative example, the interface material of the bioelectronic interface is a commercially available hydrogel, which differs from the interface functional partitioning type interface layer used in Example 1. Commercial hydrogels are prone to swelling in humid environments, leading to a decrease in interfacial adhesion over time, thus affecting the stability of bioelectrophysiological signal acquisition. This comparative example serves to illustrate the technical advantages of the asymmetric interface bioelectronic interface described in this invention in terms of anti-swelling performance and sustained bioelectrophysiological signal acquisition.
[0062] Effect
[0063] (1) This invention constructs an asymmetric interface structure with interface functional partitions in the bioelectronic interface, which makes the target tissue contact interface and the non-target interface have significant differences in interface behavior. During the acquisition of bioelectrophysiological signals, the target tissue side interface can form a stable, low-impedance electrode-tissue coupling interface, while the non-target side interface effectively weakens the mechanical disturbances introduced by external tissue traction and friction, thereby significantly reducing the generation of motion artifacts and background noise.
[0064] (2) Thanks to the above-mentioned interface functional partitioning design, the bioelectronic interface of the present invention can maintain a stable interface state under dynamic and humid conditions, resulting in higher signal-to-noise ratio and smaller baseline drift of the acquired bio-electromyography signals. Experimental results show that the signal-to-noise ratio of the bio-electromyography signals acquired by the present invention is approximately 33.288 dB, which is significantly higher than the 15.442 dB of the two-sided hydrophilic bioelectronic interface with uniform interface properties, representing an improvement of approximately 2.15 times.
[0065] (3) Furthermore, compared with commercially available gel electrodes purchased directly, the asymmetric interface bioelectronic interface of the present invention exhibits superior interface stability in humid environments. Figure 5 Commercial gels are prone to swelling during use, leading to decreased interfacial adhesion and signal attenuation. In contrast, this invention effectively suppresses interfacial state changes through interfacial structure design, maintaining stable bioelectrophysiological signal output even during experiments with continuous monitoring for more than 2 hours. Its signal-to-noise ratio is approximately 2.04 times higher than that of commercial gel electrodes.
[0066] In summary, the asymmetric interface bioelectronic interface proposed in this invention can simultaneously achieve stable coupling of signals from the target tissue and effective suppression of interference from non-target interfaces without relying on complex material modification or additional adhesion layers, significantly improving the signal-to-noise ratio and long-term stability of bioelectrophysiological signal acquisition. Compared with existing bifacial hydrophilic interface bioelectronic interfaces and commercial gel electrodes (as shown in the comparative examples), this invention has significant advantages in signal quality and persistent monitoring capabilities under dynamic physiological environments, making it suitable for future applications such as electronic skin, health monitoring, and human-computer interaction.
[0067] The above description is only a preferred embodiment of the present invention, but the scope of protection of the present invention is not limited thereto. Any equivalent substitutions or modifications made by those skilled in the art within the scope of the technology disclosed in the present invention, based on the technical solution and inventive concept of the present invention, should be covered within the scope of protection of the present invention.
Claims
1. A bioelectronic interface system with functional partitioning of the interface, characterized in that, The method for preparing the bioelectronic interface system includes the following steps: Step 1: Preparation of flexible conductive pathways: Liquid metal is mixed with a polymer stabilizer and subjected to ultrasonic treatment under water bath conditions to obtain conductive ink with good dispersibility; the conductive ink is patterned and deposited on the surface of a flexible polyurethane (PU) film using direct writing printing. After the solvent evaporates, mechanical treatment is used to form a continuous conductive network inside the conductive pathway, thereby obtaining a stretchable and bendable flexible conductive pathway. Step 2, Flexible encapsulation of conductive paths: The conductive paths obtained in Step 1 are bonded or covered with a flexible encapsulation layer, so that the conductive paths are encapsulated inside the flexible structure, thereby improving their mechanical and electrical stability in tensile, bending and humid environments. Step 3: Construction of the Interface Functional Zoning Flexible Interface Layer: A flexible interface layer is constructed on one side of the conductive path, and at least two functionally distinct interface regions are formed in this interface layer. The interface regions include a target tissue signal coupling region and a non-target tissue interference suppression region. The target tissue signal coupling region is used to form a stable, low-impedance electrode-tissue coupling interface with the target biological tissue. The non-target tissue interference suppression region is used to reduce the impact of adhesion, friction, and mechanical disturbances generated during contact with non-target tissues on signal acquisition. Step 4: Integration of the overall bioelectronic interface system: The structures built in Steps 2 and 3 are integrated to form an integrated bioelectronic interface system, which enables the acquisition, transmission and interface stabilization of bioelectric signals.
2. The bioelectronic interface system with functional partitioning according to claim 1, characterized in that: In step one, the liquid metal is gallium-based or a gallium-indium alloy with a gallium-indium molar ratio of 1:1 to 1:5; the ultrasonic power is 200 to 800 W.
3. The bioelectronic interface system with functional partitioning according to claim 1, characterized in that: In step three, the interface functional partitioning is achieved through differences in interface layer composition, interface configuration, or interface contact method, rather than relying on uniform changes in the overall performance of a single interface material.