A biomimetic cutaneous ionic electrochemical pressure sensor array and a method of making the same

Abstract

The application discloses a bionic skin texture ionic electrochemical pressure sensor array and a preparation method thereof, relates to the technical field of flexible tactile sensing, and the sensor array comprises an upper electrode, a lower electrode and an ionic sensitive layer located between the upper electrode and the lower electrode; the surface of the ionic sensitive layer is provided with a bionic skin texture microstructure, and the inside of the ionic sensitive layer is provided with a porous structure; and the upper electrode and the lower electrode adopt different electrode materials. The bionic skin texture microstructure is arranged on the surface of the ionic sensitive layer which is compounded by thermoplastic polyurethane (TPU) and ionic liquid, the inside of the ionic sensitive layer is provided with a porous structure, the ionic sensitive layer is embedded into a pore array of a polydimethylsiloxane (PDMS) isolation layer to form an independent pressure sensitive unit, and an asymmetric electrochemical interface is constructed between the upper electrode and the lower electrode, so that the purposes of improving the static pressure continuous detection stability, inhibiting the crosstalk between array elements, enhancing the heterogeneous interface combination reliability and improving the low pressure sensitivity are achieved.

Description

Technical Field

[0001] This invention relates to the field of flexible tactile sensing technology, specifically to a biomimetic skin texture ion electrochemical pressure sensor array and its fabrication method. Background Technology

[0002] With the rapid development of wearable electronics, embody intelligent systems, and electronic skin, flexible pressure sensors, as core components for human-computer interaction and environmental perception, need to achieve stable and reliable pressure detection under conditions of complex curved surface attachment, dynamic deformation adaptation, and multi-point contact. Existing flexible pressure sensors are mainly classified into resistive, capacitive, piezoelectric, and triboelectric types according to their working principles. Among them, self-powered sensors mostly rely on triboelectric or piezoelectric mechanisms, utilizing contact electrification or the flow of polarized charges to generate signals. These sensors respond quickly and require no external power supply in dynamic pressure detection scenarios. However, under continuous static pressure, due to charge dissipation or polarization saturation, the output signal is difficult to maintain stability, resulting in significantly insufficient static pressure holding detection capability. Meanwhile, flexible pressure sensors based on ion conduction mechanisms, due to their unique interfacial double-layer effect and ion migration characteristics, can generate stable electrochemical signals under continuous pressure and have received widespread attention in recent years.

[0003] Existing ion sensors generally employ a structure with a continuous ion-sensitive layer between the upper and lower electrodes. External pressure alters the actual contact area between the electrodes and the sensitive layer, enhancing interfacial ion migration and electrochemical processes to achieve voltage or current output. To further improve sensitivity, researchers have fabricated microstructure arrays on the surface of the sensitive layer using a template method to amplify the rate of change in contact area in the low-pressure region. However, in such schemes, the sensitive layer and the elastic substrate are often physically bonded or simply adhered. Under repeated bending, shearing, and long-term cyclic loading, the microstructures are prone to detachment or wear, resulting in insufficient interfacial bonding reliability. Furthermore, in array applications, the continuous sensitive layer structure lacks effective physical isolation between array elements, leading to significant crosstalk issues in electric field and ion conduction, thus limiting multi-point pressure resolution. Additionally, existing sensitive layers are often dense structures, relying solely on surface deformation to transmit signals under pressure, resulting in insufficient bulk ion response and hindering further improvements in low-pressure sensitivity and linear detection range.

[0004] To address the challenges of interface stability and element independence in ion-type sensor arrays, researchers have undertaken valuable explorations. For example, the team led by Guo Chuanfei at Southern University of Science and Technology reported an ion-capacitive skin fabrication strategy that embeds microstructured ionogels into a porous array of polydimethylsiloxane (PDMS) elastomers. This strategy uses a composite ionogel of polyvinyl alcohol (PVA) and phosphoric acid (H3PO4) as the sensing unit. An intermediate PDMS isolation layer is formed by laser drilling. After injecting the precursor, the surface microstructure is formed by sandpaper template. Covalent bonding between the ionogel and the pore walls is achieved by treating with 3-aminopropyltriethoxysilane (APTES) and glutaraldehyde (GA). High sensitivity, high mechanical stability, and low crosstalk sensing performance are achieved in a 28×28 array. The device operates stably for over 100,000 cycles under combined 44 kPa shear and 200 kPa compression.

[0005] However, this approach uses polyvinyl alcohol (PVA) hydrogel as the sensitive material, which limits its water resistance and long-term stability. Furthermore, the interface adhesion strategy only targets the single combination of PVA and polydimethylsiloxane (PDMS), without addressing the heterogeneous interface system of thermoplastic polyurethane (TPU) and PDMS. In contrast, TPU has significant advantages in mechanical strength, processing compatibility, and integration with flexible circuit boards, and there is currently no mature adhesion solution for it.

[0006] On the other hand, existing fatigue-resistant hydrogel interfacial adhesion strategies mainly rely on physical interactions. For example, using polyvinyl alcohol (PVA) hydrogel as a model, densely oriented nanocrystalline domains are formed through freeze-thaw cycles and dry annealing. These domains are then anchored firmly to rigid substrates such as glass through numerous hydrogen bonds, achieving an interfacial fatigue threshold of up to 800 Jm. -2 However, this mechanism is based on the unique crystallization behavior of hydrogels, making it difficult to directly migrate to hydrophobic ion-conducting polymer systems such as thermoplastic polyurethane (TPU), and its adhesion effect on low surface energy elastomers such as polydimethylsiloxane (PDMS) is limited.

[0007] The aforementioned technological status indicates that flexible ion pressure sensor arrays face multiple technical bottlenecks in their path towards practical application, including the difficulty in integrating the microstructure of the sensitive layer surface with the internal ion transport channels, crosstalk caused by the lack of effective physical isolation between array elements, and insufficient long-term reliability of heterogeneous interfaces such as thermoplastic polyurethane (TPU) and polydimethylsiloxane (PDMS). Therefore, a systematic solution is needed that can simultaneously address surface morphology control, bulk structure optimization, independent isolation of array elements, and multi-layer interface reinforcement to promote the reliable application of flexible ion pressure sensor arrays under complex operating conditions. Summary of the Invention

[0008] The purpose of this invention is to provide a biomimetic skin texture ion electrochemical pressure sensor array and its preparation method. This is achieved by setting a biomimetic skin texture microstructure on the surface of an ion-sensitive layer composed of thermoplastic polyurethane (TPU) and an ion liquid, and by setting a porous structure inside the layer. This ion-sensitive layer is then embedded within the porous array of a polydimethylsiloxane (PDMS) isolation layer to form independent pressure-sensitive units. Simultaneously, an asymmetric electrochemical interface is constructed between the upper and lower electrodes. This addresses the technical problems of existing flexible ion pressure sensor arrays, such as insufficient stability in continuous static pressure detection, severe crosstalk between array elements, poor reliability of heterogeneous interface bonding, and limited low-pressure sensitivity.

[0009] To address the aforementioned technical problems, in a first aspect, embodiments of the present invention provide a biomimetic skin texture ion electrochemical pressure sensor array, comprising an upper electrode, a lower electrode, and an ion-sensitive layer located between the upper electrode and the lower electrode; The ion-sensitive layer comprises a material composed of thermoplastic polyurethane (TPU) and an ionic liquid. The surface of the ion-sensitive layer is provided with a biomimetic skin texture microstructure, and its interior is provided with a porous structure. The ion-sensitive layer is embedded in the pore array of the polydimethylsiloxane (PDMS) isolation layer, and the pore array of the PDMS isolation layer corresponds to the electrode pattern array of the upper electrode to form multiple independent pressure-sensitive units. The upper electrode and the lower electrode are made of different electrode materials, and an electrochemical interface is formed between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, respectively. When the ion-sensitive layer is compressed and deformed under external pressure, the biomimetic skin texture microstructure and the porous structure increase the contact area between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, shortening the ion migration path and forming a potential difference inside the pressure sensor array.

[0010] As an optional implementation, the upper electrode comprises a gold-plated electrode on a polyimide (PI) substrate, and the lower electrode comprises a composite electrode of a polyethylene terephthalate (PET) substrate and copper foil.

[0011] As an alternative implementation, the ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer are covalently bonded and / or chemically bonded through surface functionalization.

[0012] As an optional implementation, the pore array of the polydimethylsiloxane PDMS isolation layer has a unit diameter of 3-5 mm, a pore spacing of 1.5 mm-2.5 mm, and an array size of 3×3-10×10.

[0013] Secondly, embodiments of the present invention provide a method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array, comprising the following steps: S1: Mix the ionic liquid with thermoplastic polyurethane (TPU) in a certain proportion, add a mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF), heat and stir to dissolve and let stand to defoam, to obtain a thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution. S2: The polydimethylsiloxane PDMS main agent and curing agent are mixed in a weight ratio, degassed, and then cast and cured to form an array pore structure on the cured polydimethylsiloxane PDMS film. After being activated by oxygen plasma, the surface is chemically functionalized to obtain the polydimethylsiloxane PDMS isolation layer. S3: A lower electrode is formed by combining a polyethylene terephthalate (PET) substrate with copper foil. The PET surface of the lower electrode is subjected to oxygen plasma treatment and then immersed in a 3-glycidyl ether oxypropyltrimethoxysilane (GLYMO) solution for surface functionalization. The functionalized polydimethylsiloxane (PDMS) isolation layer is aligned, bonded, and pressed with the lower electrode to form a pre-assembled structure. S4: The thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution is injected into the cavity of the pre-assembled structure and moisture-induced phase separation is performed to allow the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature is raised to evaporate the solvent, and an embedded biomimetic skin texture ion-sensitive layer is obtained. S5: The surface of the upper electrode is treated with oxygen plasma and coated with adhesive, aligned and attached to the polydimethylsiloxane PDMS isolation layer and cured to complete the array assembly.

[0014] As an optional implementation, the ionic liquid in S1 comprises 1-ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imine ([EMIM]). + [TFSI] - The ionic liquid accounts for 15 to 25 wt% of the total mass of thermoplastic polyurethane (TPU) and ionic liquid.

[0015] Specifically, in this embodiment of the invention, the dosage range is used to balance ion transport capability, TPU skeleton mechanical support, and interfacial electrochemical stability. When the ionic liquid content is below 15 wt%, the number of free ions in the sensitive layer is insufficient, the continuous ion transport channels are inadequate, the bulk ion resistance and interfacial charge transfer impedance are large, the voltage change under pressure is small, and the low-pressure sensitivity is insufficient. When the ionic liquid content is above 25 wt%, the plasticizing effect of the ionic liquid on TPU is too strong, the modulus of the sensitive layer decreases too much, and ion liquid leakage and near-saturation of interfacial reaction are likely to occur in the low-pressure stage. When the pressure increases subsequently, the space for new ion migration is limited, the response range becomes narrower, and signal drift is more likely to occur during cycling. Therefore, 15–25 wt% is a range that balances ion conductivity, structural recovery, and output stability, with 20 wt% being the optimal ratio for this system.

[0016] As an optional embodiment, the mass ratio of N,N-dimethylformamide (DMF) to tetrahydrofuran (THF) is 2–4:1, and the mass ratio of the mixed solvent to thermoplastic polyurethane (TPU) is 4–6:1. In this embodiment of the invention, the mass ratio is used to control the solvent evaporation rate and the degree of aggregation during the wet phase separation process. Due to aggregation, the growth of droplets forming porous structures inside is delayed in the upper layer because the continuous polymer-rich phase reaches its glass transition concentration later than the bulk membrane, making droplet coarsening more likely. As the solvent continues to diffuse and evaporate, TPU self-assembles on the surface to form a skin-textured microstructure, which is specifically manifested in controlling the microstructure on the gel surface and the size of the internal pores. If the DMF ratio is too high, surface microstructures cannot be formed, resulting in a smooth surface; if the THF ratio is too high, evaporation is too rapid, leading to insufficient phase separation.

[0017] The heating and stirring temperature in S1 is 70-90℃, the stirring time is 4-8h, and the settling and defoaming time is 4-8h.

[0018] Specifically, the heating and stirring temperature is set to 70–90°C, primarily to promote the full expansion of TPU segments and accelerate their dissolution in the DMF / THF mixed solvent, while ensuring uniform dispersion of the ionic liquid in the polymer phase. Below 70°C, the TPU dissolves slowly, and undissolved polymer particles or locally high-viscosity regions are likely to exist in the precursor. Above 90°C, THF volatilization accelerates significantly, the mixed solvent ratio easily deviates from the set range, and the precursor concentration and viscosity change.

[0019] As an optional implementation, the process conditions for the moisture-induced phase separation treatment in S4 include: temperature 30-40°C, relative humidity 70-90%RH, and treatment time 6-12h.

[0020] As an optional implementation, the weight ratio of the polydimethylsiloxane PDMS main agent to the curing agent in S2 is 8-12:1, the curing temperature is 60-80°C, and the curing time is 2-4 hours. The concentration of the 3-glycidyl etheroxypropyltrimethoxysilane GLYMO solution in S3 is 3-7% V / V (less than 3% results in a smaller degree of modification and weakened binding strength, while greater than 7% wastes raw materials), and the soaking time is 15-25 minutes.

[0021] Specifically, the GLYMO concentration range is used to control the silane bridging density between PET and PDMS. After oxygen plasma treatment, oxygen-containing active sites such as hydroxyl groups are introduced into the PET surface. The silanol groups generated after GLYMO hydrolysis can condense with the PET surface and form Si–O–Si connections with the PDMS side, thereby improving the PDMS / PET interfacial bonding strength. When the GLYMO concentration is less than 3% V / V, the effective silane coverage on the PET surface is insufficient, the interfacial bridging point density is low, and localized debonding and insufficient peel strength are prone to occur after lamination. When the GLYMO concentration is greater than 7% V / V, excess GLYMO is more likely to undergo self-condensation in the solution or on the surface, forming an uneven silane aggregate layer. This aggregate layer will not continue to effectively improve the interfacial strength and may also increase interfacial brittleness, residues, and processing costs.

[0022] As an optional implementation, the pressing temperature in S3 is 50-70°C, the pressing load is 300-500g, and the pressing time is 4-8h.

[0023] Compared with the prior art, the embodiments of the present invention have the following advantages and beneficial effects: 1. This invention provides a biomimetic skin-textured ion electrochemical pressure sensor array. The sensor array includes an upper electrode, a lower electrode, and an ion-sensitive layer located between the upper and lower electrodes. The ion-sensitive layer is composed of a composite material of thermoplastic polyurethane (TPU) and an ionic liquid. Its surface is provided with a biomimetic skin-textured microstructure, and its interior is provided with a porous structure. When external pressure is applied to the sensor array, the pressure is first transmitted to the upper electrode and then to the ion-sensitive layer. Under the pressure, the ion-sensitive layer undergoes compression deformation. The biomimetic skin-textured microstructure on its surface and the porous structure inside work together to make the upper electrode and the ion-sensitive layer... The actual contact area between layers and between the lower electrode and the ion-sensitive layer is significantly increased. At the same time, the pore walls of the internal porous structure are deformed by pressure, shortening the ion migration path. At this time, the ions inside the ion-sensitive layer are redistributed, and the interfacial ion migration, electron transfer and charge exchange processes between the upper electrode and the ion-sensitive layer and between the lower electrode and the ion-sensitive layer are enhanced. Since the upper electrode and the lower electrode use different electrode materials, different interfacial electrochemical states are formed between the two electrode materials and the ion-sensitive layer. As a result, a potential difference is formed inside the pressure sensor array. This potential difference can be output by an external circuit to form a detectable voltage signal.

[0024] Specifically, when the external pressure increases, the compression of the biomimetic skin-like microstructure and porous structure increases, resulting in more effective contact between the electrode and the ion-sensitive layer. This further enhances the interfacial charge exchange process, causing the output voltage to change accordingly. When the external pressure decreases or is released, the thermoplastic polyurethane (TPU) matrix recovers through its own elasticity, and the biomimetic skin-like microstructure and porous structure gradually return to their original morphology. The interfacial contact state and ion distribution state also recover, and the output voltage returns to the corresponding baseline level. By detecting the magnitude and change of the output voltage, the magnitude of the external pressure, the loading process, and the release process can be determined, enabling the detection of both static and dynamic pressure. Simultaneously, the ion-sensitive layer is embedded in polydimethylsiloxane P... In the pore array of the DMS isolation layer, the pore array of the polydimethylsiloxane (PDMS) isolation layer corresponds to the electrode pattern array of the upper electrode, forming multiple relatively independent pressure-sensitive units. This embedded configuration allows each sensing unit to have a relatively independent force-bearing region and ion conduction boundary, effectively blocking the ion migration and electric field coupling paths between adjacent array elements, thereby reducing crosstalk between array elements and ensuring the independence and accuracy of the signals of each array element during multi-point pressure detection. In addition, the elastic dissipation characteristics of the PDMS isolation layer can buffer external shear and compressive stress, protect the ion-sensitive layer and its interface with the electrode, and improve the mechanical stability and long-term output consistency of the device under bending and cyclic conditions.

[0025] 2. In this embodiment of the invention, a gold-plated electrode on a polyimide (PI) substrate is used as the upper electrode, and a composite electrode of a polyethylene terephthalate (PET) substrate and copper foil is used as the lower electrode. Because the gold and copper electrodes have different electrochemical properties, they form different interfacial electrochemical states when in contact with the thermoplastic polyurethane (TPU) / ionic liquid ion-sensitive layer. When the ion-sensitive layer is compressed and deformed under external pressure, the interfacial contact states between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, change differently. This results in asynchronous changes in the electrochemical potential at the interface between the two electrodes and the electrolyte, thereby generating a potential difference between the upper and lower electrodes and outputting a voltage signal. Simultaneously, the polyimide (PI) substrate provides high-temperature and bending-resistant mechanical support for the gold-plated electrode, and the PET substrate provides flexible support for the copper foil electrode, ensuring the integrity of the electrode structure and the stability of the electrochemical interface under complex curved surfaces and dynamic deformation conditions.

[0026] 3. In this embodiment of the invention, a covalent bond or chemical bond is formed between the ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer through surface functionalization. Specifically, after the surface of the PDMS isolation layer is activated by oxygen plasma, a 3-aminopropyltriethoxysilane (APTES) silane layer is grafted onto it. A thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution is in-situ cast onto the 3-aminopropyltriethoxysilane (APTES)-functionalized PDMS surface and then cured by heating. This embodiment of the invention constructs a chemical bond between the thermoplastic polyurethane (TPU) and the PDMS through surface functionalization. Bridging is achieved through covalent bonding and / or chemical bonding of the thermoplastic polyurethane (TPU) / polydimethylsiloxane (PDMS) heterostructure interface (the amino groups on the PDMS surface modified by APTES can form multi-site hydrogen bonds with the C=O (carbonyl) and N–H (nitrogen-hydrogen bonds) in the hard segment urethane of TPU, and the C–O–C (ether bonds) in the soft segment polyether, thereby improving the bonding strength). This interfacial bonding mechanism enables the ion-sensitive layer to resist interfacial shear and peel stress under cyclic compression and bending conditions, preventing array element boundary failure and sensitive layer slippage, thus ensuring the structural integrity and output signal consistency of the array device during long-term cyclic use.

[0027] Specifically, the multilayer interface reinforcement includes three interfaces: TPU / PDMS, PDMS / PET, and PDMS / PI. For the TPU / PDMS interface, oxygen plasma first introduces active groups such as silanol groups onto the PDMS surface. After APTES hydrolysis, it condenses with the PDMS surface to form a silane coupling layer with amino-terminated groups. During in-situ casting and curing of the TPU / ionic liquid precursor, the exposed –NH2 and partially protonated –NH groups of APTES are... 3+It can form multi-site hydrogen bonds with the C=O and N–H groups in the hard segment urethane of TPU and the C–O–C groups in the soft segment polyether; [EMIM] + [TFSI] - Near the interface, ion pairs, hydrogen bonds, and ion-dipole couplings can form with amine sites, creating a "dynamic hydrogen bond network + ion coupling network" at the interface. This network can break and reform during peeling and cyclic compression, thereby dissipating applied energy and improving interfacial toughness. For the PDMS / PET interface, GLYMO-modified PET and PDMS form PET–O–Si and Si–O–Si bridging networks through hydrolysis-condensation. For the PDMS / PI interface, E43, after curing, forms a continuous Si–O–Si crosslinked network and interacts strongly with the polar groups on the PI surface. This multi-interface synergistic enhancement makes the device less prone to interlayer slippage, delamination, and signal drift under bending, compression, and peeling loads.

[0028] 4. The preparation method provided in this embodiment of the invention involves injecting a thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution into a cavity pre-assembled from a polydimethylsiloxane (PDMS) isolation layer and a lower electrode. Moisture-induced phase separation is then performed for 6–12 hours at a temperature of 30–40 °C and a relative humidity of 60–80% RH. The moisture in the environment induces phase separation between the TPU and the ionic liquid and solvent system, causing the TPU to self-assemble on the surface to form a biomimetic skin-textured microstructure and internally to form a porous structure. Subsequently, the temperature is raised to allow the solvent to fully evaporate, resulting in an embedded biomimetic skin-textured ion-sensitive layer. This in-situ construction method allows the skin-textured microstructure and porous structure to be formed synchronously and integrally within the cavity, eliminating the need for subsequent template imprinting or additional assembly processes. This avoids interface defects and delamination risks associated with traditional bonding processes. Simultaneously, the surface skin-textured microstructure generates a significant change in contact area in the low-pressure region, while the internal porous structure shortens the ion migration path and enhances the bulk ion response under pressure. Both synergistically improve the sensitivity and linear detection range of the sensor array in the low-pressure range. Attached Figure Description

[0029] To more clearly illustrate the technical solutions of the exemplary embodiments of the present invention, the accompanying drawings used in the embodiments will be briefly described below. It should be understood that the following drawings only show some embodiments of the present invention and should not be considered as a limitation of the scope. For those skilled in the art, other related drawings can be obtained based on these drawings without creative effort. In the drawings: Figure 1 This is a schematic diagram of the preparation process of the interface bonding experiment between ionic liquid TPU and PDMS in an embodiment of the present invention; Figure 2 This is a schematic diagram of the ion-sensitive layer assembly process in an embodiment of the present invention; Figure 3 These are scanning images of the surface morphology and internal structure of the sensitive layer as a function of humidity in an embodiment of the present invention. Figure 3 a-1 in the figure corresponds to the surface morphology of the sensitive layer at 60% RH. Figure 3 a-2 in the figure corresponds to the surface morphology of the sensitive layer at 70% RH. Figure 3 a-3 in the figure corresponds to the surface morphology of the sensitive layer at 80% RH. Figure 3 b-1 in the figure corresponds to the pore structure of the sensitive layer cross section at 60% RH. Figure 3 b-2 in the figure corresponds to the pore structure of the sensitive layer cross section at 70% RH. Figure 3 b-3 in the figure corresponds to the pore structure of the sensitive layer section at 80% RH. Figure 3 c-1 in the image corresponds to a magnified view of the pore size at 60% RH. Figure 3 c-2 in the image corresponds to a magnified view of the pore size at 70% RH. Figure 3 c-3 in the image corresponds to a magnified view of the pore size at 80% RH. Figure 4 This is a schematic diagram of the lower electrode assembly process in an embodiment of the present invention; Figure 5 This is a schematic diagram of the peel force test in an embodiment of the present invention, wherein... Figure 5 In this context, 'a' corresponds to the test result. Figure 5 The 'b' in the text corresponds to the test method; Figure 6 This is a graph showing the sensitivity detection results of the sensor array prepared in Example 1 of the present invention; Figure 7 This is a schematic diagram of the static pressure response sensitivity results of the sensor array prepared in Example 1 of the present invention; Figure 8 This is a graph showing the continuous pressure monitoring results of the sensor array prepared in Example 1 of the present invention; Figure 9 This is a graph showing the continuous pressing cycle stability of the sensor array prepared in Example 1 of the present invention; Figure 10 A comparison of crosstalk between neighboring array elements when a single array element is pressed, showing the difference between array sensors with and without isolation structures. Figure 11 A comparison of the drift characteristics of devices with strong interface bonding and devices without strong interface bonding; Figure 12 For XPS maps, where Figure 12 In the image, 'a' corresponds to the XPS spectrum on the PDMS side after stripping. Figure 12 In the image, b corresponds to the XPS spectrum of the TPU side after stripping. Detailed Implementation

[0030] To make the objectives, technical solutions, and advantages of the embodiments of the present invention clearer, the technical solutions in the embodiments of the present invention will be clearly and completely described below. Obviously, the described embodiments are some embodiments of the present invention, but not all embodiments.

[0031] Therefore, the detailed description of the embodiments of the present invention provided below is not intended to limit the scope of the claimed invention, but merely to illustrate selected embodiments of the invention. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without inventive effort are within the scope of protection of the present invention.

[0032] This invention provides a method for fabricating a biomimetic skin-textured ion electrochemical pressure sensor array. A biomimetic skin-textured microstructure is formed on the surface of a TPU / ionic liquid sensitive layer through moisture-induced phase separation, and an interconnected porous structure is formed internally to enhance interfacial deformation and ion migration response in the low-pressure region. Independent array element boundaries are formed by embedding the sensitive layer into a PDMS isolation layer to reduce crosstalk between array elements. Surface functionalization and multilayer interface bonding design improve the bonding strength between TPU, PDMS, PET, and PI, thereby enhancing the stability of the array device under bending and cyclic operating conditions. Specifically, it includes the following: (1) Preparation of thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution 1-Ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([EMIM]) + [TFSI] - Ionic liquid and thermoplastic polyurethane (TPU) are mixed in a certain proportion, wherein the mass fraction of ionic liquid is 15-25 wt%. A mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF) is added, with a mass ratio of N,N-dimethylformamide (DMF) to tetrahydrofuran (THF) of 2-4:1. The mass ratio of the mixed solvent to TPU is 4-6:1. The mixture is magnetically stirred at 70-90°C for 4-8 hours until completely dissolved. After standing for 4-8 hours to defoam, a homogeneous precursor solution is obtained.

[0033] (2) Preparation of polydimethylsiloxane PDMS isolation layer Please combine Figure 1 , Figure 2As shown, polydimethylsiloxane (PDMS) main agent and curing agent (such as Dow Corning DC184 product from Dow Corning) are mixed at a weight ratio of 8–12:1. After vacuum degassing, the mixture is poured into a mold and cured at 60–80°C for 2–4 hours to obtain a PDMS film. An array of pores is formed on the cured PDMS film using a cutting device. The unit diameter is 3–5 mm, the pore spacing is 1.5–2.5 mm, and the array size is 3×3–10×10. After cleaning, the PDMS film with the pore array is placed in a plasma cleaner. The activated polydimethylsiloxane (PDMS) was subjected to oxygen plasma activation treatment at an oxygen pressure of 80–120 Pa, a power of 120–180 W, and a treatment time of 90–150 s. The activated PDMS was then immersed in a 3–7% V / V solution of 3-aminopropyltriethoxysilane (APTES) for 15–25 min. After the 3-aminopropyltriethoxysilane (APTES) formed a stable silane layer on the surface of the activated PDMS through a hydrolysis-condensation reaction, the PDMS was removed and dried with nitrogen to obtain a surface-functionalized PDMS isolation layer.

[0034] (3) Preparation of the lower electrode A polyethylene terephthalate (PET) substrate was composited with copper foil to form a PET-copper lower electrode. The PET surface was then subjected to oxygen plasma treatment at an oxygen pressure of 80–120 Pa, a power of 120–180 W, and a treatment time of 90–150 s. The surface was then immersed in a 3–7% V / V solution of 3-glycidyl etheroxypropyltrimethoxysilane (GLYMO) for 15–25 min to functionalize the surface. The functionalized polydimethylsiloxane (PDMS) insulating layer was aligned and bonded to the functionalized PET-copper lower electrode. A pressure load of 300–500 g was applied at 50–70 °C for 4–8 h to form a pre-assembled structure.

[0035] (4) Constructing a biomimetic skin texture ion-sensitive layer The precursor solution prepared above is injected into the cavity of the pre-assembled structure (filled). The component is placed in a constant temperature and humidity environment for moisture-induced phase separation treatment at a temperature of 30-40°C and a relative humidity of 70-90%RH for 6-12 hours. This allows the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature is raised to 50-70°C to allow the solvent to fully evaporate, resulting in an embedded biomimetic skin texture ion-sensitive layer. This ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer are covalently bonded and / or chemically bonded through surface functionalization.

[0036] (5) Assemble the upper electrode: Reference Figure 4 As shown, a gold-plated electrode on a polyimide (PI) substrate is used as the top electrode. The top electrode array elements are circular electrode patterns with a diameter of 3–5 mm and a spacing of 1.5–2.5 mm, corresponding to the hole array of the polydimethylsiloxane (PDMS) isolation layer. The surface of the polyimide (PI)-gold top electrode is treated with oxygen plasma and coated with an adhesive (such as E43). The oxygen pressure is 80–120 Pa, the power is 120–180 W, and the treatment time is 90–150 s. Then, it is aligned and attached to the PDMS isolation layer and cured at room temperature for 10–14 h to complete the array assembly and obtain a biomimetic skin texture ion electrochemical pressure sensor array.

[0037] The electrical signal of the biomimetic skin texture ion electrochemical pressure sensor array provided in this embodiment of the invention is generated by the principle of ion electrochemistry and has a pressing structure with upper and lower electrodes. The reason for adopting the pressing structure with upper and lower electrodes is that this structure allows external pressure to be directly transmitted to the ion-sensitive layer and maintains a relatively stable electrical signal output under continuous pressure, thereby realizing the detection of static and dynamic pressure.

[0038] The basic pressure signal acquisition process is as follows: The sensor is installed on the object or contact surface to be measured, with the pressure-bearing surface of the sensor facing the direction of external pressure, ensuring that the external pressure acts perpendicularly on the upper electrode of the sensor. The upper electrode of the sensor is a gold-plated electrode on a polyimide (PI) substrate, and the lower electrode is a composite electrode of a polyethylene terephthalate (PET) substrate and copper foil. In between is a thermoplastic polyurethane (TPU) / ionic liquid biomimetic skin-textured ion-sensitive layer formed by moisture-induced phase separation. This biomimetic skin-textured ion-sensitive layer is embedded in a pore array within a polydimethylsiloxane (PDMS) isolation layer, forming a sandwich-type pressure-sensitive unit together with the upper and lower electrodes. This structure allows each pressure-sensitive unit to have a relatively independent force-bearing area, thereby reducing crosstalk between array units.

[0039] When external pressure is applied to the upper surface of the sensor, the pressure is first transmitted to the polyimide (PI)-gold upper electrode, and then to the thermoplastic polyurethane (TPU) / ionic liquid biomimetic skin-textured ion-sensitive layer. Under pressure, this ion-sensitive layer undergoes compression deformation. The biomimetic skin-textured microstructure on its surface and the porous structure inside increase the actual contact area between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, while simultaneously shortening the ion migration path. At this time, ions within the ion-sensitive layer redistribute, and the interfacial ion migration, electron transfer, and charge exchange processes between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, are enhanced. Since the upper and lower electrodes are made of gold and copper respectively, different interfacial electrochemical states are formed between the two electrode materials and the ion-sensitive layer, thus creating a potential difference within the device. This potential difference is output through an external circuit to form a detectable voltage signal.

[0040] When external pressure increases, the compression of the biomimetic skin-textured ion-sensitive layer increases, resulting in more effective contact between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer. This further enhances the interfacial charge exchange process, causing a change in the output voltage. When the external pressure decreases or is released, the thermoplastic polyurethane (TPU) matrix recovers through its own elasticity, and the biomimetic skin-textured ion-sensitive layer gradually returns to its original morphology. The interfacial contact state and ion distribution state also recover, and the output voltage returns to the corresponding baseline level. By detecting the magnitude and change of the output voltage, the magnitude of the external pressure, the loading process, and the release process can be determined.

[0041] It should be noted that, in the embodiments of the present invention, the fabrication processes of the polyethylene terephthalate (PET)-copper lower electrode and the polyimide (PI) substrate gold-plated upper electrode both adopt conventional processes in the prior art or commercially available products, and are not limited here. Exemplary examples are provided below: The preparation method of polyethylene terephthalate (PET) copper lower electrode is as follows: select a PET flexible substrate with a thickness of 50 μm, and coat the surface of it with a copper foil with a thickness of 18–35 μm.

[0042] The gold-plated top electrode on the polyimide (PI) substrate is a commercially available flexible PI electrode manufactured by OEMs such as JLCPCB, JLCPCB, and Huaqiu.

[0043] To better demonstrate the significant effects of the embodiments of the present invention, specific experimental examples will be set up below to verify the effects.

[0044] Example 1: This embodiment of the invention provides a method for fabricating a biomimetic dermal texture ion electrochemical pressure sensor array, comprising the following: (1) Preparation of thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution 1-Ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([EMIM]) + [TFSI] - Ionic liquid and thermoplastic polyurethane (TPU) were mixed in a certain proportion, wherein the ionic liquid had a mass fraction of 20 wt%. A mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF) was added, with a mass ratio of 3:1. The mass ratio of the mixed solvent to TPU was 5:1. The mixture was magnetically stirred at 80°C for 6 hours until completely dissolved. After standing for 6 hours to remove bubbles, a homogeneous precursor solution was obtained.

[0045] (2) Preparation of polydimethylsiloxane PDMS isolation layer Please combine Figure 1 , Figure 2As shown, polydimethylsiloxane PDMS main agent and curing agent (Dow Corning DC184 product from Dow Corning) were mixed at a weight ratio of 10:1. After vacuum degassing, the mixture was poured into a mold and cured at 70°C for 3 hours to obtain a polydimethylsiloxane PDMS film. An array of pores was formed on the cured polydimethylsiloxane PDMS film using a cutting device. The unit diameter was 4 mm, the pore spacing was 2 mm, and the array size was 5×5. After cleaning, the polydimethylsiloxane PDMS film with the pore array was subjected to plasma cleaning. The device is activated by oxygen plasma at a pressure of 100 Pa and a power of 150 W for 120 s. The activated polydimethylsiloxane PDMS is then immersed in a 5% V / V solution of 3-aminopropyltriethoxysilane APTES for 20 min. After the 3-aminopropyltriethoxysilane APTES forms a stable silane layer on the surface of the activated polydimethylsiloxane PDMS through a hydrolysis-condensation reaction, the PDMS is removed and dried with nitrogen to obtain a surface-functionalized polydimethylsiloxane PDMS isolation layer.

[0046] (3) Preparation of the lower electrode A polyethylene terephthalate (PET) substrate was composited with copper foil to form a PET-copper lower electrode. The PET surface was then subjected to oxygen plasma treatment at an oxygen pressure of 100 Pa, a power of 150 W, and a treatment time of 120 s. The surface was then immersed in a 5% V / V solution of 3-glycidyl etheroxypropyltrimethoxysilane (GLYMO) for 20 min to functionalize the surface. The functionalized polydimethylsiloxane (PDMS) insulating layer was aligned and bonded to the functionalized PET-copper lower electrode. A 400 g compression load was applied at 60 °C for 6 h to form a pre-assembled structure.

[0047] (4) Constructing a biomimetic skin texture ion-sensitive layer The precursor solution prepared above was injected into the cavity of the pre-assembled structure. The component was placed in a constant temperature and humidity environment for moisture-induced phase separation treatment at a temperature of 35°C and a relative humidity of 70%RH for 9 hours. This allowed the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature was raised to 60°C to allow the solvent to fully evaporate, resulting in an embedded biomimetic skin texture ion-sensitive layer. This ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer were covalently bonded and / or chemically bonded through surface functionalization.

[0048] (5) Assemble the upper electrode: Reference Figure 4As shown, a gold-plated electrode on a polyimide (PI) substrate is used as the top electrode. The top electrode array consists of circular electrode patterns with a diameter of 4 mm and a spacing of 2 mm, corresponding to the hole array of the polydimethylsiloxane (PDMS) isolation layer. The surface of the polyimide (PI)-gold top electrode is treated with oxygen plasma (oxygen pressure of 100 Pa, power of 150 W, treatment time of 120 s) and coated with E43. Then, it is aligned and attached to the PDMS isolation layer and cured at room temperature for 12 h to complete the array assembly, thus obtaining a biomimetic skin texture ion electrochemical pressure sensor array.

[0049] Example 2: This embodiment of the invention provides a method for fabricating a biomimetic dermal texture ion electrochemical pressure sensor array, comprising the following: (1) Preparation of thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution 1-Ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([EMIM]) + [TFSI] - Ionic liquid and thermoplastic polyurethane (TPU) were mixed in a certain proportion, wherein the ionic liquid had a mass fraction of 5 wt%. A mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF) was added, with a mass ratio of 2:1. The mass ratio of the mixed solvent to TPU was 4:1. The mixture was magnetically stirred at 70°C for 8 hours until completely dissolved. After standing for 8 hours to remove bubbles, a homogeneous precursor solution was obtained.

[0050] (2) Preparation of polydimethylsiloxane PDMS isolation layer Please combine Figure 1 , Figure 2 As shown, polydimethylsiloxane (PDMS) main agent and curing agent (Dow Corning DC184 compatible product from Dow Corning) were mixed at a weight ratio of 8:1. After vacuum degassing, the mixture was poured into a mold and cured at 60°C for 4 hours to obtain a PDMS film. An array of pores was formed on the cured PDMS film using a cutting device. The unit diameter was 3 mm, the pore spacing was 1.5 mm, and the array size was 10 × 10. After cleaning the PDMS film with the pore array, it was placed in a plasma ion exchange system. The cleaning machine is used for oxygen plasma activation treatment at an oxygen pressure of 80 Pa, a power of 120 W, and a treatment time of 150 s. The activated polydimethylsiloxane PDMS is then immersed in a 3% V / V solution of 3-aminopropyltriethoxysilane APTES for 25 min. After the 3-aminopropyltriethoxysilane APTES forms a stable silane layer on the surface of the activated polydimethylsiloxane PDMS through a hydrolysis-condensation reaction, the PDMS is removed and dried with nitrogen to obtain a surface-functionalized polydimethylsiloxane PDMS isolation layer.

[0051] (3) Preparation of the lower electrode A polyethylene terephthalate (PET) substrate was composited with copper foil to form a PET-copper lower electrode. The PET surface was then subjected to oxygen plasma treatment at an oxygen pressure of 80 Pa, a power of 120 W, and a treatment time of 150 s. The surface was then immersed in a 3% V / V solution of 3-glycidyl etheroxypropyltrimethoxysilane (GLYMO) for 25 min to functionalize the surface. The functionalized polydimethylsiloxane (PDMS) insulating layer was aligned and bonded to the functionalized PET-copper lower electrode. A 300 g compression load was applied at 50 °C for 8 h to form a pre-assembled structure.

[0052] (4) Constructing a biomimetic skin texture ion-sensitive layer The precursor solution prepared above was injected into the cavity of the pre-assembled structure. The component was placed in a constant temperature and humidity environment for moisture-induced phase separation treatment at 30°C and 70%RH for 12 hours. This allowed the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature was raised to 50°C to allow the solvent to fully evaporate, resulting in an embedded biomimetic skin texture ion-sensitive layer. This ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer were covalently bonded and / or chemically bonded through surface functionalization.

[0053] (5) Assemble the upper electrode: Reference Figure 4 As shown, a gold-plated electrode on a polyimide (PI) substrate is used as the top electrode. The top electrode array consists of circular electrode patterns with a diameter of 3 mm and a spacing of 1.5 mm, corresponding to the hole array of the polydimethylsiloxane (PDMS) isolation layer. The surface of the polyimide (PI)-gold top electrode is treated with oxygen plasma (oxygen pressure of 80 Pa, power of 120 W, treatment time of 150 s) and coated with E43. Then, it is aligned and attached to the PDMS isolation layer and cured at room temperature for 14 h to complete the array assembly, thus obtaining a biomimetic skin texture ion electrochemical pressure sensor array.

[0054] Example 3: This embodiment of the invention provides a method for fabricating a biomimetic dermal texture ion electrochemical pressure sensor array, comprising the following: (1) Preparation of thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution 1-Ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([EMIM]) + [TFSI] -Ionic liquid and thermoplastic polyurethane (TPU) were mixed in a certain proportion, wherein the ionic liquid had a mass fraction of 50 wt%. A mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF) was added, with a mass ratio of 4:1. The mass ratio of the mixed solvent to TPU was 6:1. The mixture was magnetically stirred at 90°C for 4 hours until completely dissolved. After standing for 4 hours to remove bubbles, a homogeneous precursor solution was obtained.

[0055] (2) Preparation of polydimethylsiloxane PDMS isolation layer Please combine Figure 1 , Figure 2 As shown, polydimethylsiloxane PDMS main agent and curing agent (Dow Corning DC184 product from Dow Corning) were mixed at a weight ratio of 12:1. After vacuum degassing, the mixture was poured into a mold and cured at 80°C for 2 hours to obtain a polydimethylsiloxane PDMS film. An array of pores was formed on the cured polydimethylsiloxane PDMS film using a cutting device. The unit diameter was 5 mm, the pore spacing was 2.5 mm, and the array size was 3×3. After cleaning the polydimethylsiloxane PDMS film with the pore array, it was placed in a plasma ion exchange system. The cleaning machine is used for oxygen plasma activation treatment at an oxygen pressure of 120 Pa, a power of 180 W, and a treatment time of 90 s. The activated polydimethylsiloxane PDMS is then immersed in a 7% V / V solution of 3-aminopropyltriethoxysilane APTES for 15 min. After the 3-aminopropyltriethoxysilane APTES forms a stable silane layer on the surface of the activated polydimethylsiloxane PDMS through a hydrolysis-condensation reaction, the PDMS is removed and dried with nitrogen to obtain a surface-functionalized polydimethylsiloxane PDMS isolation layer.

[0056] (3) Preparation of the lower electrode A polyethylene terephthalate (PET) substrate was composited with copper foil to form a PET-copper lower electrode. The PET surface was then subjected to oxygen plasma treatment at an oxygen pressure of 120 Pa, a power of 180 W, and a treatment time of 90 s. The surface was then immersed in a 7% V / V solution of 3-glycidyl etheroxypropyltrimethoxysilane (GLYMO) for 15 min to functionalize the surface. The functionalized polydimethylsiloxane (PDMS) insulating layer was aligned and bonded to the functionalized PET-copper lower electrode. A 500 g compression load was applied at 70 °C for 4 h to form a pre-assembled structure.

[0057] (4) Constructing a biomimetic skin texture ion-sensitive layer The precursor solution prepared above was injected into the cavity of the pre-assembled structure. The component was placed in a constant temperature and humidity environment for moisture-induced phase separation treatment at a temperature of 40°C and a relative humidity of 80%RH for 6 hours. This allowed the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature was raised to 70°C to allow the solvent to fully evaporate, resulting in an embedded biomimetic skin texture ion-sensitive layer. This ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer were covalently bonded and / or chemically bonded through surface functionalization.

[0058] (5) Assemble the upper electrode: Reference Figure 4 As shown, a gold-plated electrode on a polyimide (PI) substrate is used as the top electrode. The top electrode array consists of circular electrode patterns with a diameter of 5 mm and a spacing of 2.5 mm, corresponding to the hole array of the polydimethylsiloxane (PDMS) isolation layer. The surface of the polyimide (PI)-gold top electrode is treated with oxygen plasma (oxygen pressure of 120 Pa, power of 180 W, treatment time of 90 s) and coated with E43. Then, it is aligned and attached to the PDMS isolation layer and cured at room temperature for 10 h to complete the array assembly, thus obtaining a biomimetic skin texture ion electrochemical pressure sensor array.

[0059] Example 4: This embodiment of the invention provides a method for preparing a biomimetic skin texture ion electrochemical pressure sensor array. The difference from Example 1 is that in step (4) of constructing the biomimetic skin texture ion sensitive layer, the relative humidity is 60%RH, while the other steps remain unchanged.

[0060] Example 5: This embodiment of the invention provides a method for preparing a biomimetic skin texture ion electrochemical pressure sensor array. The difference from Example 1 is that in step (4) of constructing the biomimetic skin texture ion sensitive layer, the relative humidity is 80%RH, while the other steps remain unchanged.

[0061] Example 6: This embodiment of the invention provides a method for preparing a biomimetic skin texture ion electrochemical pressure sensor array. The difference from Example 1 is that in step (4) of constructing the biomimetic skin texture ion sensitive layer, the temperature of the moisture-induced phase separation treatment is 25°C, while the other steps remain unchanged.

[0062] Example 7: This embodiment of the invention provides a method for preparing a biomimetic skin texture ion electrochemical pressure sensor array. The difference from Example 1 is that in step (4) of constructing the biomimetic skin texture ion sensitive layer, the temperature of the moisture-induced phase separation treatment is 42°C, while the other steps remain unchanged.

[0063] Comparative Example 1: A method for fabricating an ion electrochemical pressure sensor array is provided. The difference from Example 1 is that the independent pore array structure in the PDMS isolation layer is eliminated, and a single continuous TPU / ionic liquid biomimetic skin-textured sensitive layer is sandwiched between the upper and lower electrodes. The fabrication method is as follows: the same TPU / ionic liquid precursor solution as in Example 1 is integrally cast onto the surface of the PET-copper lower electrode. Moisture-induced phase separation and drying are performed according to the same temperature, humidity, and time as in Example 1, followed by covering with the PI-gold upper electrode to complete the assembly. In this comparative example, each element shares the same continuous ion-sensitive layer and does not have independent force boundaries and ion transport boundaries formed by the PDMS isolation wall; the remaining steps and parameters remain unchanged.

[0064] Comparative Example 2: A method for fabricating an ion electrochemical pressure sensor array is provided, differing from Example 1 in that strong interfacial bonding treatment is not performed. Specifically, the PDMS isolation layer is not functionalized with APTES, the PET-copper lower electrode is not functionalized with GLYMO, and the PI-gold upper electrode is not bonded to the PDMS with an E43 adhesive layer or only with ordinary physical bonding. The TPU / ionic liquid precursor is still injected into the PDMS cavity and subjected to moisture-induced phase separation and drying according to the conditions of Example 1. This comparative example retains the isolation pore array, but the layers are mainly connected by physical contact or weak adhesion, while the remaining steps and parameters remain unchanged.

[0065] Comparative Example 3: A method for preparing an ion electrochemical pressure sensor array is provided. The difference from Example 1 is that moisture-induced phase separation is not performed. Specifically, after injecting the TPU / ionic liquid precursor solution into the pores of the pre-assembled structure, instead of placing it in a constant temperature and humidity environment for moisture-induced phase separation, it is directly thermally evaporated and dried at 60°C to form a film, resulting in a dense ion-sensitive layer with a relatively smooth surface and insufficient internal pores. The remaining steps remain unchanged.

[0066] The performance of the ion electrochemical pressure sensor arrays prepared in Examples 1-7 and Comparative Examples 1-3 was tested, and the results are shown in Tables 1 and 2 below: Table 1

[0067] Table 2

[0068] As shown in Tables 1 and 2, the embodiments of the present invention provide a biomimetic skin texture microstructure on the surface of an ion-sensitive layer composed of thermoplastic polyurethane (TPU) and an ion liquid, and a porous structure inside. This ion-sensitive layer is then embedded in the pore array of a polydimethylsiloxane (PDMS) isolation layer to form an independent pressure-sensitive unit. At the same time, an asymmetric electrochemical interface is constructed between the upper and lower electrodes. This achieves the purpose of improving the stability of continuous static pressure detection, suppressing crosstalk between array elements, enhancing the reliability of heterogeneous interface bonding, and improving low-pressure sensitivity.

[0069] Please combine Figure 3 As shown in the comparison of Examples 1, 2, 3, 4, and 5, it can be seen that the moisture-induced phase separation process affects the surface morphology and internal structure of the TPU / ionic liquid sensitive layer. When the relative humidity is 70-90%RH, moisture can enter the wet film at a suitable rate and undergo solvent / non-solvent exchange with DMF / THF, forming a more uniform biomimetic skin texture structure on the surface of the sensitive layer and a more stable interconnected pore structure inside. In Example 4, the relative humidity is 60%RH, resulting in insufficient moisture input, a lower degree of phase separation, shallower annular wrinkles on the surface, and poor regularity and connectivity of the internal pores. Under pressure, the effective contact area and bulk ion migration space increase only slightly, so the sensitivity is low at both low and high pressures. In Example 5, the relative humidity is 80%RH, resulting in excessive moisture input, an excessively fast solvent / non-solvent exchange rate, and excessive surface roughening. The pore walls tend to thin or collapse locally, reducing structural consistency and thus decreasing pressure response repeatability and sensitivity. Therefore, 70-80% RH can enable the surface texture microstructure and internal porous structure to form synergistically, thereby increasing the interfacial deformation amplitude under pressure and enhancing the ion migration response.

[0070] Please combine Figure 5 As shown, this invention significantly improves the peel strength of the TPU / PDMS interface. The peel strength is obtained using a T-shaped peel test. During the test, the two flexible films to be tested are fixed to the moving and fixed fixtures of an electronic universal testing machine (Shimadzu AGS-X), forming a "T" shape. Peeling is then performed at a constant speed, and the average peel force during the stable peeling process is recorded. The peel strength is calculated using G=F / W, where G is the peel strength in N / m; F is the average peel force of the stable peeling section in N; and W is the width of the peeled sample interface in meters. This test method follows the basic principles of GB / T 2791-1995 "Test Method for T-Peel Strength of Adhesives - Flexible Materials to Flexible Materials".

[0071] Significantly improved peel strength: By employing APTES functionalization combined with ion-dipole and multi-site hydrogen bonding, the peel strength of the TPU-PDMS interface has been significantly improved compared to untreated interfaces (corresponding to...). Figure 5The value of 3.829 N / m at TP#1 in the middle, and after oxygen plasma treatment only (corresponding to Figure 5 The TP#2) in the present invention is 8.52 N / m, which is increased to 245.14 N / m in Embodiment 1 of the present invention.

[0072] Failure modes can be detected by Figure 12 The determination was made by observing the cross-section after exfoliation and by X-ray photoelectron spectroscopy (XPS). During XPS testing, the exfoliated PDMS side and TPU side were fixed on the sample stage, and surface elemental spectra were acquired and the F and Si elemental signals were compared. No new F peak appeared in the XPS spectrum of the exfoliated PDMS side, indicating the presence of [TFSI]. - The TPU / ionic liquid phase did not show significant residue on the PDMS side; after peeling, the XPS spectrum of the TPU side showed a significant Si signal, indicating the presence of PDMS residue on the TPU surface. In Example 1 of this invention, the interface failure changed from traditional interface debonding to PDMS bulk cohesive failure, indicating that the interface bonding strength was higher than the PDMS bulk failure strength. This is because the APTES functionalized interface forms a multi-site hydrogen bond, ion pair, and ion-dipole coupling network. This network can undergo dynamic reconstruction under external force, thereby dissipating peeling energy and improving fatigue resistance.

[0073] Please refer to Figure 6 As shown, the pressure sensing performance was tested using an electronic universal testing machine (Shimadzu AGS-X). During testing, the sensor was placed on a flat, insulated platform with a surface area of ​​1 cm². 2 A planar pressure head is used to vertically press the sensor; the loading speed is set to 20 mm / min; simultaneously, the output voltage between the upper and lower electrodes is recorded using an external data acquisition system. The pressure is obtained by dividing the loading force by the pressure head area. Sensitivity is calculated using S=ΔV / ΔP, where ΔV is the change in output voltage and ΔP is the change in pressure. The sensor array prepared in Example 1 exhibits a stable response within the range of 0.03–250 kPa, with a maximum sensitivity of 9.138 mV / kPa in the low-pressure region. (Refer to...) Figure 7 As shown, a stable output with a peak value of approximately 14 mV can still be obtained at 30 Pa. (Refer to...) Figure 8 As shown, the device's output signal remains relatively stable under a constant voltage of 100 kPa for 200 s. (Refer to...) Figure 9 As shown, the sensor maintains stable output even after more than 2,000 load / unload cycles, indicating that the biomimetic skin texture sensitive layer and strong interface bonding structure can improve the stability of the device during cyclic use.

[0074] Please combine Figure 10As shown, array crosstalk testing was performed using an electronic universal testing machine (Shimadzu AGS-X). During testing, the sensor array was fixed on an insulating platform, and a planar indenter matching the size of a single electrode was used to vertically press the central target element at a speed of 20 mm / min, while simultaneously recording the output voltages of the target element and adjacent elements. The crosstalk rate was calculated as the ratio of the peak output of adjacent elements to the peak output of the target element. Comparing Example 1 and Comparative Example 1, after introducing the PDMS isolation structure in Example 1 of this invention, the crosstalk between adjacent units was reduced from 54.4% to 0.05%. This result indicates that the PDMS isolation wall can limit the deformation of the sensitive layer and the ion transport range, thereby improving the accuracy of array pressure imaging and facilitating array imaging and multi-point localization.

[0075] Please combine Figure 11 As shown, comparing Example 1 and Comparative Example 2, it can be seen that Comparative Example 2 did not undergo strong interface bonding treatment. After bending or repeated pressing, the device is more prone to interface micro-slippage and localized debonding, leading to changes in the contact area and stress transmission path. Therefore, the output signal is more prone to baseline drift. In Example 1 of this invention, after constructing multiple strong interfaces using APTES, GLYMO, and E43, the connections between functional layers are more stable, and the output signal remains relatively stable after repeated deformation.

[0076] Referring to Tables 1 and 2, a comparison of Examples 6 and 7 with Example 1 shows that the temperature of the moisture-induced phase separation treatment affects solvent evaporation, non-solvent diffusion, and phase separation rate. In Example 6, when the treatment temperature is below 30°C, the solvent evaporation rate is slow, the phase separation time after water enters the wet film is prolonged, droplets are more prone to excessive aggregation, and the size uniformity of surface texture and internal pores decreases, leading to uneven interface deformation under pressure and reduced output repeatability. In Example 7, when the treatment temperature is above 40°C, solvents such as THF evaporate too quickly, the wet film surface is prone to premature solidification, the water / solvent exchange time is insufficient, nucleation phase separation is incomplete, the surface texture is weak or the pore structure is discontinuous, and the effective contact area in the low-pressure zone changes insufficiently. Therefore, a moisture-induced phase separation treatment temperature in the range of 30–40°C is more conducive to the formation of stable biomimetic textures and interconnected pores.

[0077] A comparison of Comparative Example 3 and Example 1 shows that Comparative Example 3 did not undergo moisture-induced phase separation treatment, but instead directly formed a film through thermal evaporation. This method lacks a water-based, non-solvent-induced phase separation process, making it difficult to form a biomimetic skin-like microstructure on the surface of the sensitive layer, and also lacking uniformly interconnected pores inside. When the dense sensitive layer is under pressure, it mainly relies on overall elastic compression, resulting in small changes in the contact area in the low-pressure region. There is insufficient space for bulk ion migration and extrusion, therefore, the low-pressure sensitivity and effective response range are both smaller than those of Example 1.

[0078] The specific embodiments described above further illustrate the purpose, technical solution, and beneficial effects of the present invention. It should be understood that the above description is only a specific embodiment of the present invention and is not intended to limit the scope of protection of the present invention. Any modifications, equivalent substitutions, improvements, etc., made within the spirit and principles of the present invention should be included within the scope of protection of the present invention.

Claims

1. A biomimetic dermal texture ion electrochemical pressure sensor array, characterized in that, It includes an upper electrode, a lower electrode, and an ion-sensitive layer located between the upper electrode and the lower electrode; The ion-sensitive layer comprises a material composed of thermoplastic polyurethane (TPU) and an ionic liquid. The surface of the ion-sensitive layer is provided with a biomimetic skin texture microstructure, and its interior is provided with a porous structure. The ion-sensitive layer is embedded in the pore array of the polydimethylsiloxane (PDMS) isolation layer, and the pore array of the PDMS isolation layer corresponds to the electrode pattern array of the upper electrode to form multiple independent pressure-sensitive units. The upper electrode and the lower electrode are made of different electrode materials, and an electrochemical interface is formed between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, respectively. When the ion-sensitive layer is compressed and deformed under external pressure, the biomimetic skin texture microstructure and the porous structure increase the contact area between the upper electrode and the ion-sensitive layer, and between the lower electrode and the ion-sensitive layer, shortening the ion migration path and forming a potential difference inside the pressure sensor array. The upper electrode includes a gold-plated electrode on a polyimide (PI) substrate, and the lower electrode includes a composite electrode of a polyethylene terephthalate (PET) substrate and a copper foil. The surface of the polydimethylsiloxane (PDMS) isolation layer is grafted with a 3-aminopropyltriethoxysilane (APTES) silane layer after being activated by oxygen plasma.

2. The biomimetic dermal texture ion electrochemical pressure sensor array according to claim 1, characterized in that, The ion-sensitive layer and the polydimethylsiloxane (PDMS) isolation layer are covalently bonded and / or chemically bonded through surface functionalization.

3. The biomimetic dermal texture ion electrochemical pressure sensor array according to claim 1, characterized in that, In the pore array of the polydimethylsiloxane (PDMS) isolation layer, the unit diameter is 3–5 mm, the pore spacing is 1.5 mm–2.5 mm, and the array size is 3×3–10×10.

4. A method for fabricating a biomimetic dermal texture ion electrochemical pressure sensor array as described in any one of claims 1 to 3, characterized in that, Includes the following steps: S1: Mix the ionic liquid with thermoplastic polyurethane (TPU) in a certain proportion, add a mixed solvent of N,N-dimethylformamide (DMF) and tetrahydrofuran (THF), heat and stir to dissolve and let stand to defoam, to obtain a thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution. S2: The polydimethylsiloxane PDMS main agent and curing agent are mixed in a weight ratio, degassed, and then cast and cured to form an array pore structure on the cured polydimethylsiloxane PDMS film. After being activated by oxygen plasma, the surface is chemically functionalized to obtain the polydimethylsiloxane PDMS isolation layer. S3: A lower electrode is formed by combining a polyethylene terephthalate (PET) substrate with copper foil. The PET surface of the lower electrode is subjected to oxygen plasma treatment and then immersed in a 3-glycidyl ether oxypropyltrimethoxysilane (GLYMO) solution for surface functionalization. The functionalized polydimethylsiloxane (PDMS) isolation layer is aligned, bonded, and pressed with the lower electrode to form a pre-assembled structure. S4: The thermoplastic polyurethane (TPU) ion-sensitive layer precursor solution is injected into the cavity of the pre-assembled structure and moisture-induced phase separation is performed to allow the thermoplastic polyurethane (TPU) to self-assemble on the surface to form a biomimetic skin texture microstructure and form a porous structure inside. Then, the temperature is raised to evaporate the solvent, and an embedded biomimetic skin texture ion-sensitive layer is obtained. S5: The surface of the upper electrode is treated with oxygen plasma and coated with adhesive, aligned and attached to the polydimethylsiloxane PDMS isolation layer and cured to complete the array assembly.

5. The method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array according to claim 4, characterized in that, The ionic liquid in S1 comprises 1-ethyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide, and the ionic liquid accounts for 15 to 25 wt% of the total mass of the thermoplastic polyurethane (TPU) and the ionic liquid.

6. The method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array according to claim 5, characterized in that, The mass ratio of N,N-dimethylformamide (DMF) to tetrahydrofuran (THF) is 2–4:1, and the mass ratio of the mixed solvent to thermoplastic polyurethane (TPU) is 4–6:

1. The heating and stirring temperature in S1 is 70-90℃, the stirring time is 4-8h, and the settling and defoaming time is 4-8h.

7. The method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array according to claim 4, characterized in that, The process conditions for the moisture-induced phase separation treatment in S4 include: temperature 30-40℃, relative humidity 70-90%RH, and treatment time 6-12h.

8. The method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array according to claim 4, characterized in that, The weight ratio of the polydimethylsiloxane PDMS main agent to the curing agent in S2 is 8-12:1, the curing temperature is 60-80℃, and the curing time is 2-4h. The concentration of the 3-glycidyl etheroxypropyltrimethoxysilane GLYMO solution in S3 is 3-7% V / V, and the soaking time is 15 min-25 min.

9. The method for fabricating a biomimetic skin texture ion electrochemical pressure sensor array according to claim 4, characterized in that, The pressing temperature in S3 is 50-70°C, the pressing load is 300-500g, and the pressing time is 4-8h.

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