A biomimetic tactile pain-sensing dual-modal flexible pressure sensor and its fabrication method

Abstract

This invention discloses a biomimetic tactile-pain-sensitive dual-modal flexible pressure sensor and its fabrication method. The sensor comprises an upper electrode, a tactile-pain-sensitive piezoelectric sensing component, and a lower electrode. The electrode employs an elastic substrate layer / modified functional layer / conductive network. The tactile-pain-sensitive piezoelectric sensing component comprises a magnesium-doped zinc oxide nanomaterial composite layer designed based on a concentration gradient and an elastic polymer composite layer, as well as a common ground electrode. The tactile sensing layer achieves high sensitivity by blending high-concentration Mg-ZnO nanorod filler with PDMS; the pain sensing layer uses low-concentration Mg-ZnO nanorod filler, resulting in lower sensitivity. Independent signal pathways are formed through the common ground electrode, enabling dual-modal sensing of tactile perception and pain alarm. The sensor fabricated in this invention possesses advantages such as high sensitivity, wide detection range, and strong resistance to signal crosstalk, and can be applied to biomimetic sensing fields such as intelligent robots and wearable devices.

Description

Technical Field

[0001] This invention belongs to the field of flexible pressure sensor technology, specifically relating to a biomimetic tactile pain-sensing dual-modal flexible pressure sensor and its preparation method. Background Technology

[0002] Tactile perception is one of the core technologies of bionic robots, intelligent prostheses, and human-computer interaction systems. Its core lies in simulating the multimodal perception ability of human skin to external pressure, vibration, and pain signals through sensors (ZHAN Z, YANGY, ZUO W, et al. Recent advances and challenges of tactile sensing forrobotics: from fundamentals to applications [J]. Materials Today Physics, 2025, 54(101740.)). Among the existing technologies, piezoelectric tactile sensors have attracted much attention due to their self-powered characteristics and high sensitivity. Their core principle is to convert mechanical stimulation into electrical signals through piezoelectric materials (such as PZT, PVDF, ZnO, etc.) to achieve tactile sensing (PANDIT P, CHOUGALE MY, DUBAL D, et al. Amplifying Touch Using 3D ZnOTetrapods for Tactile and Haptic Intelligence [J]. Small, 2025, 21(16):2408414.). Although traditional piezoelectric tactile sensors have made some progress in single-modal detection, they still have significant shortcomings in biomimetic dual-modal (tactile and pain) collaborative sensing, wide pressure range sensitivity adaptation, and the reliability of flexible structures.

[0003] For the circuit wiring problem of large-area array tactile sensors, each unit of a piezoelectric tactile sensor has an individual lead (MAITY S, SINGH RK, GADHEWAL M, TIWARI S P. Highly biodegradable piezoelectric flexible wearable tactile sensors with amino acid crystals: aparadigm shift towards smart transient electronics [J]. Chemical Engineering Journal, 2025, 512(162531).). For an x*y sensor array, at least x*y or 2*x*y leads are required (ZHU, MIAOMIAO, LOU, MENGNA, ABDALLA, IBRAHIM, et al. Highly shape adaptive fiber based electronic skin for sensitive joint motion monitoring and tactile sensing [J]. Nano Energy, 2020, 69: 104429.) Using a wiring addressing structure with cross- and vertical electrodes can effectively reduce the number of leads, but for piezoelectric tactile sensor arrays, significant signal crosstalk is easily generated between array elements, which seriously affects the accuracy of the sensor (LIN W, WANG B, PENG G, et al. Skin-Inspired Piezoelectric Tactile Sensor Array with Crosstalk-Free Row+Column Electrodes for Spatiotemporally Distinguishing Diverse Stimuli [J]. Adv Sci (Weinh), 2021, 8(3): 2002817.). Summary of the Invention

[0004] The purpose of this invention is to overcome the above-mentioned shortcomings of the existing technology and provide a biomimetic tactile and pain-sensing dual-modal flexible pressure sensor and its preparation method. By designing a gradient piezoelectric material and using a common-ground electrode structure, the invention solves the technical problems of existing flexible sensors being unable to simultaneously achieve tactile and pain-sensing dual-modal perception, having severe signal crosstalk, and experiencing sensitivity attenuation under high voltage.

[0005] A biomimetic tactile-pain-sensitive dual-modal flexible pressure sensor includes an upper electrode, a lower electrode, and a tactile-pain-sensitive piezoelectric sensing component located between the upper and lower electrodes. The tactile-pain-sensitive piezoelectric sensing component includes a tactile sensing layer, a pain sensing layer, and a common ground electrode located between the tactile sensing layer and the pain sensing layer. The tactile sensing layer senses tactile pressure, and the pain sensing layer senses pain pressure. The tactile-pain-sensitive piezoelectric sensing component generates a voltage through the piezoelectric effect, which is then transmitted to an external processing circuit via the upper and lower electrodes. The tactile sensing layer and the pain sensing layer are constructed using an elastic polymer mixed with magnesium-doped zinc oxide (Mg-ZnO) nanorods of varying masses, achieving dual-modal sensing of tactile pressure and pain pressure through a piezoelectric gradient response. The tactile pressure is less than the pain pressure.

[0006] Furthermore, the tactile sensing layer is composed of 25-40 wt% Mg-ZnO nanorods and an elastic polymer, and the pain sensing layer is composed of 5-20 wt% Mg-ZnO nanorods and an elastic polymer, wherein the magnesium doping amount in the Mg-ZnO nanorods is 0.1-2 at% atomic percentage.

[0007] Furthermore, the thickness of both the tactile sensing layer and the pain sensing layer is 50-200 μm, and the elastic polymer is selected from PDMS, TPU or PU.

[0008] Furthermore, both the upper and lower electrodes comprise an elastic substrate layer, a modified functional layer, and a conductive network from bottom to top; the conductive network is used to electrically connect the tactile-pain piezoelectric sensing component and the external wire in the tactile-pain pressure sensor.

[0009] Furthermore, the elastic substrate layer is selected from polydimethylsiloxane (PDMS), thermoplastic polyurethane (TPU), or polyurethane (PU), with a thickness of 50-200 μm and a Young's modulus of 0.5-5 MPa; the modified functional layer is composed of a cross-linked polyvinyl alcohol (C-PVA) and lithium chloride (LiCl) composite, wherein LiCl accounts for 2-15 wt% of the mass of the cross-linked polyvinyl alcohol, and the cross-linking agent is glutaraldehyde; the conductive network material is one or more of silver nanowires (AgNWs), carbon nanotubes (CNTs), or two-dimensional transition metal carbon / nitrides (Mxene), with a sheet resistance ≤50 Ω / sq.

[0010] Furthermore, the conductive network is formed into an island-bridge mesh structure by spin coating, photolithography, or screen printing, which is a stretchable network composed of functional units (islands) and connecting structures (bridges). The modified functional layer achieves complete encapsulation of the conductive network by spin coating at 2000-7000 rpm.

[0011] Furthermore, when the conductive network uses silver nanowires (AgNWs), it is spin-coated at 800-3000 rpm with a solution of 5-50 mg / mL concentration, in which the solvent is ethanol, isopropanol, water, or a combination of two of them, and then annealed at 80-120℃ to form a patterned mesh.

[0012] Furthermore, the common ground electrode is formed between the tactile sensing layer and the pain sensing layer by vapor deposition or magnetron sputtering; the material of the common ground electrode is gold (Au), silver (Ag), or aluminum (Al), and the vacuum degree of the vapor deposition is ≤8×10⁻⁶. -4 Pa, deposition rate 0.5-2 Å / s; common electrode thickness 50-200 nm, sheet resistance ≤5 Ω / sq, surface roughness Ra≤20 nm.

[0013] Furthermore, the common ground electrode of the touch-pain piezoelectric sensing component is formed between the touch sensing layer and the pain sensing layer by vapor deposition or magnetron sputtering, and the touch sensing layer and the pain sensing layer are automatically bonded together by the PDMS curing effect.

[0014] Furthermore, the flexible pressure sensor forms an independent signal path through the common ground electrode. The voltage signal output by the tactile sensing layer is acquired through the upper electrode-common ground electrode, and the voltage signal output by the pain sensing layer is acquired through the common ground electrode-lower electrode. The pressure is determined by comparing the voltage amplitudes. When the pressure is greater than or equal to the pressure threshold of the tactile sensing layer, the tactile sensing layer maintains a saturated output voltage signal V1 (V1 is the saturation voltage of the tactile sensing layer), and the pain sensing layer outputs a voltage signal greater than V2 (V2 is the trigger threshold voltage of the pain sensing layer). In this case, it is determined to be pain pressure, and an alarm signal is triggered. When the pressure is less than the pressure threshold of the tactile sensing layer, the tactile sensing layer outputs a voltage signal less than V1, and the pain sensing layer outputs a voltage signal less than V2. In this case, it is determined to be tactile pressure.

[0015] The present invention also provides a biomimetic tactile pain-sensing bimodal flexible pressure sensor array, which includes a plurality of the biomimetic tactile pain-sensing bimodal flexible pressure sensors arranged in rows and columns, wherein the plurality of upper electrodes located in the same row are connected in series, and the upper electrodes located in different rows are not connected to each other; the plurality of lower electrodes located in the same column are connected in series, and the lower electrodes located in different columns are not connected, and the connection is an electrical connection.

[0016] The present invention also provides a method for fabricating the aforementioned biomimetic tactile pain-sensing dual-modal flexible pressure sensor, comprising:

[0017] (1) Prepare the upper electrode and the lower electrode;

[0018] (2) Preparation of touch-pain piezoelectric sensing component: a pain sensing layer is spin-coated on the lower electrode and cured to form a pain sensing layer; a common ground electrode is vapor-deposited on the pain sensing layer; a touch sensing layer is spin-coated on the common ground electrode and cured to form a touch sensing layer. At this time, the structure of lower electrode-pain sensing layer-common ground electrode-touch sensing layer is formed from bottom to top.

[0019] (3) The upper electrode and the lower electrode-pain sensing layer-common ground electrode-touch sensing layer structure are packaged by hot pressing to form the entire device.

[0020] Furthermore, the method specifically includes:

[0021] (1) Preparation of the upper and lower electrodes:

[0022] a. Apply an organic or aqueous solution of 5-50 mg / mL silver nanowires (AgNWs), carbon nanotubes (CNTs) or MXene to a glass substrate by spin-coating, photolithography or screen printing at 800-2000 rpm to form an island-bridge structure, and then anneal at 80-120 °C.

[0023] b. Modified functional layer embedding: A mixed aqueous solution of cross-linked polyvinyl alcohol (C-PVA) and lithium chloride (LiCl) is spin-coated at 2000-7000 rpm to fill the conductive network on the glass substrate;

[0024] c. Elastic substrate layer coating: PDMS, TPU, or PU solution is spin-coated onto the modified functional layer at 800-2000 rpm. (PDMS is usually a liquid prepolymer system composed of a base adhesive (main agent) and a curing agent (crosslinking agent), which is liquid after mixing. TPU can be dissolved in polar solvents (such as DMF, THF, acetone) to form a solution, and a solid film is formed after the solvent evaporates. PU is dissolved in solvents (such as ethyl acetate, DMF) to form a solution, and a film is formed after the solvent evaporates. Flexible electrodes are obtained by mechanical exfoliation.)

[0025] (2) Fabrication of touch-pain piezoelectric sensing components:

[0026] a. Pain-sensing layer: 5-20 wt% of Mg-ZnO nanorods are added to the PDMS prepolymer, and then the mixture is spin-coated, screen-printed or printed onto the lower electrode. After curing, a pain-sensing layer of 50-200 μm is formed, which together with the lower electrode promotes the lower electrode-pain-sensing layer.

[0027] b. Common-ground electrode deposition: On the pain-sensing layer, under a vacuum degree ≤8×10 -4 Au, Ag, or Al are vapor-deposited under Pa at a deposition rate of 0.5-2 Å / s to prepare a common ground electrode, forming a lower electrode-pain-sensing layer-common ground electrode;

[0028] c. Tactile sensing layer: 25-40wt% of Mg-ZnO nanorods are added to the PDMS prepolymer, and then the mixture is spin-coated, screen-printed or printed onto the common ground electrode and cured to form a 50-200μm tactile sensing layer, forming a lower electrode-pain sensing layer-common ground electrode-tactile sensing layer structure, i.e., lower electrode-touch-pain piezoelectric sensing component.

[0029] (3) Device packaging:

[0030] The upper and lower electrodes of the tactile pain-sensing piezoelectric sensing component are bonded by hot pressing (80-120℃, 0.1-0.5MPa, holding pressure for 30-60 seconds), with an alignment deviation ≤100μm and an edge overlap ≥90%.

[0031] This invention proposes a biomimetic tactile pain-sensing dual-modal flexible pressure sensor and its fabrication method, which breaks through the existing technical bottlenecks through material innovation and structural design.

[0032] Compared with existing technologies, it has the following technical advantages:

[0033] (1) Threshold differentiation between tactile and pain sensations is achieved by controlling the concentration gradient of Mg-ZnO nanorods. The tactile sensing layer and the pain sensing layer are made of elastic polymer mixed with magnesium-doped zinc oxide (Mg-ZnO) nanorods of different masses. The tactile and pain sensing dual-modal perception is achieved through the piezoelectric effect gradient response. The specific pressure threshold can be controlled by adjusting the content of Mg-ZnO nanorods.

[0034] (2) The tactile sensing layer (e.g., <50kPa) uses 25-40wt% Mg-ZnO nanorods blended with PDMS to obtain a high-sensitivity response; the pain sensing layer (e.g., ≥50kPa) uses 5-20wt% Mg-ZnO nanorods blended with PDMS to obtain a lower-sensitivity response. That is, a high concentration of ZnO nanorods corresponds to a high-sensitivity response, and a low concentration corresponds to a low-sensitivity response. An independent signal path is formed through a common ground electrode to realize dual-modal sensing of tactile perception and pain alarm.

[0035] (3) Construct a highly flexible and stable composite electrode and improve the fatigue resistance of the conductive network by alternately stacking elastic substrate and modified functional layer.

[0036] (4) The double-layer piezoelectric structure of the sensor enables the piezoelectric sensing layer to be constructed through the row and column signal encoding mechanism of the double-layer piezoelectric sensitive layer when constructing the sensor array. The row position of the point of action is identified by the output signal of the tactile sensing layer, and the column position is identified by the output signal of the pain sensing layer. The intersection of the two is the pressure point. Furthermore, the high pressure threshold of the pain sensing layer (e.g., ≥50kPa) can further distinguish the pain signal.

[0037] (5) Introduce common ground electrodes to suppress interlayer crosstalk and ensure the signal-to-noise ratio and spatial resolution of array signals.

[0038] (6) This sensor has the advantages of high sensitivity, wide detection range and strong anti-signal crosstalk capability, and can be applied to biomimetic sensing fields such as intelligent robots and wearable devices. Attached Figure Description

[0039] Figure 1 This is a schematic diagram of the structure of the biomimetic tactile pain-sensing dual-modal pressure sensor in the embodiment.

[0040] Figure 2 This is a schematic diagram of the flexible stacked electrode assembly in the embodiment.

[0041] Figure 3 This is a scanning electron microscope image of the Mg-ZnO nanorods in the example.

[0042] Figure 4 This is a diagram showing the tactile sensing voltage output of a single biomimetic tactile analgesia dual-modal pressure sensor under a pressure of 30 kPa in the embodiment.

[0043] Figure 5 This is a diagram showing the pain perception voltage output of a single biomimetic tactile pain-sensing dual-modal pressure sensor under a pressure of 60 kPa in this embodiment.

[0044] Figure 6 This is a comparative schematic diagram of the biomimetic tactile pain-sensing dual-modal pressure sensor array from different perspectives in the embodiment.

[0045] Figure 7 This is a diagram showing the touch-pain voltage output of the biomimetic touch-pain dual-modal pressure sensor array in the embodiment. Detailed Implementation

[0046] To further understand the present invention, preferred embodiments are described below in conjunction with examples. However, it should be understood that these descriptions are merely for further illustrating the features and advantages of the present invention, and not for limiting the scope of the claims. Those skilled in the art can refer to the content of the present invention and appropriately improve the process parameters to achieve the desired results. It is particularly important to note that all similar substitutions and modifications are obvious to those skilled in the art and are considered to be included within the scope of the present invention. The methods and applications of the present invention have been described through preferred embodiments. Those skilled in the art can obviously modify or appropriately change and combine the methods and applications described in the present invention without departing from the content, mechanism, spirit, and scope of the present invention to realize and apply the technology of the present invention. While it is believed that those skilled in the art fully understand the following terms, the following definitions are set forth to help illustrate the subject matter disclosed in the present invention.

[0047] As used herein, the term "comprising" is synonymous with "including," "containing," or "characterized in," and is inclusive of endpoints or open-ended, and does not exclude additional unstated elements or method steps. "Comprising" is a technical term used in the language of claims, meaning that the stated element is present, but other elements may be added and still form a construction or method within the scope of the claims.

[0048] like Figure 1 The diagram shown is a structural schematic of a biomimetic tactile-pain-sensitive dual-modal flexible pressure sensor according to this embodiment. It sequentially includes an upper electrode 101, a tactile-pain-sensitive piezoelectric sensing component 300, and a lower electrode 102. The flexible stacked electrode component used in this embodiment includes an upper electrode 101 and a lower electrode 102. Figure 2 As shown, both the upper electrode 101 and the lower electrode 102 include an elastic base layer 201, a modified functional layer 202, and a conductive network 203. The conductive network 203 is embedded within the modified functional layer 202. The touch-pain piezoelectric sensing component 300 includes a touch sensing layer 301, a pain sensing layer 302, and a common ground electrode 303. The elastic base layer 201 enables high deformation of the sensor and wrinkle-free adhesion to irregular surfaces. The modified functional layer 202 is used to enhance the collection of induced piezoelectric charges and improve the output characteristics of the piezoelectric device. The conductive network 203 is used to electrically connect the touch-pain piezoelectric sensing component in the touch-pain pressure sensor to the external wires.

[0049] In this embodiment, the tactile sensing layer 301 is designed to sense minute tactile pressures (<50kPa), generating piezoelectric induction charges through the piezoelectric effect and conducting them to external processing circuitry via electrodes. The pain sensing layer 302 is designed to sense larger pain pressures (≥50kPa), generating piezoelectric induction charges through the piezoelectric effect and conducting them to external processing circuitry via electrodes. The common ground electrode 303 is used to reduce signal crosstalk and improve device durability.

[0050] In one embodiment, the touch-pain piezoelectric sensing component forms an independent signal path through the common ground electrode. The dual-modal sensing is determined by comparing voltage amplitudes. When the pressure is greater than or equal to the pressure threshold of the touch sensing layer, the touch sensing layer maintains a saturated output voltage signal V1, and the pain sensing layer outputs a voltage signal greater than V2, which is then determined to be pain pressure. When the pressure is less than the pressure threshold of the touch sensing layer, the touch sensing layer outputs a voltage signal less than V1, and the pain sensing layer outputs a voltage signal less than V2, which is then determined to be tactile pressure.

[0051] Analysis of the tactile pain sensing mechanism of the biomimetic tactile pain-sensing dual-modal flexible pressure sensor in the embodiments of the present invention:

[0052] A biomimetic tactile-pain-sensitive dual-modal flexible pressure sensor achieves tactile-pain dual-modal sensing through the piezoelectric effect and concentration gradient design of Mg-ZnO / PDMS composite materials. In this embodiment, the tactile sensing layer (Mg-ZnO accounts for 25-40 wt% of the elastic material) forms a dense piezoelectric network due to the high concentration of Mg-ZnO nanorods, generating a highly sensitive linear response in the low pressure range (1-50 kPa). The output voltage increases to 3V with increasing pressure. When the pressure is ≥50 kPa, the output stabilizes at 3V due to the deformation saturation of the Mg-ZnO nanorods. The pressure thresholds for different components can be obtained by testing using existing techniques, and the V1 value can be obtained. Correspondingly, when the pressure just exceeds the pressure threshold of the tactile sensing layer, the output voltage of the pain sensing layer can be measured and taken as V2. The pain sensing layer (Mg-ZnO accounts for 5-20 wt% of the elastic material) has a weaker piezoelectric effect due to its low filler concentration. When the pressure is ≥50 kPa, the large deformation of the PDMS matrix forces the sparsely distributed nanorods to compress violently, and the piezoelectric output rises to (>1V). At this point, the tactile sensing layer outputs 3V due to saturation, while the pain sensing layer's voltage exceeds the 1V threshold. The two layers form independent signal paths through a common ground electrode connected in series. The common ground electrode serves as a common potential reference point, enabling the voltages of the tactile sensing layer (upper electrode - common ground electrode) and the pain sensing layer (common ground electrode - lower electrode) to be collected synchronously without interference.

[0053] In this embodiment, by comparing the voltage amplitudes of the two layers (tactile voltage amplitude of 3V and pain voltage amplitude of 1V), it can be determined whether the pressure exceeds the threshold, thus realizing dual-modal output of tactile protection and pain alarm. The following further describes a method for fabricating a biomimetic tactile and pain-sensitive dual-modal flexible pressure sensor according to this embodiment.

[0054] In this embodiment, the elastic substrate layer 201 is made of polydimethylsiloxane (PDMS), thermoplastic polyurethane (TPU) or polyurethane (PU), with a thickness of 50-200μm and a Young's modulus of 0.5-5MPa.

[0055] The modified functional layer 202 is composed of a cross-linked polyvinyl alcohol (C-PVA) and lithium chloride (LiCl) composite, wherein the mass fraction of LiCl is 2%-15wt% (based on the mass of PVA), and the cross-linking agent is glutaraldehyde;

[0056] The conductive network 203 is composed of one or more of silver nanowires (AgNWs), carbon nanotubes (CNTs), or two-dimensional transition metal carbides, nitrides, or two-dimensional transition metal carbon / nitrides (MXene), and is patterned by spin coating, photolithography, or screen printing to form an island-bridge mesh structure with a sheet resistance ≤50Ω / sq.

[0057] The tactile sensing layer 301 described in this embodiment is composed of an elastic polymer and magnesium-doped zinc oxide (Mg-ZnO) nanorods (magnesium doping amount 0.1-2 at%) accounting for 25-40 wt% of the elastic polymer, with a thickness of 50-200 μm. The elastic polymer is selected from PDMS, TPU or PU.

[0058] The pain-sensing layer 302 described in this embodiment is composed of an elastic polymer and magnesium-doped zinc oxide (Mg-ZnO) nanorods (magnesium doping amount 0.1-2 at%) accounting for 5-20 wt% of the elastic polymer, with a thickness of 50-200 μm. The elastic polymer is selected from PDMS, TPU or PU.

[0059] The common ground electrode 303 is deposited between the tactile sensing layer 301 and the pain sensing layer 302 by vapor deposition. It is made of gold, silver, or aluminum, with a thickness of 50-200 nm. The vapor deposition process parameters are: vacuum degree ≤ 8 × 10⁻⁶. -4 Pa, deposition rate 0.5-2 Å / s, surface roughness Ra≤20nm, sheet resistance≤5Ω / sq.

[0060] Example 1

[0061] A method for fabricating a biomimetic tactile pain-sensing dual-modal flexible pressure sensor includes the following steps:

[0062] S1. Fabrication of upper electrode 101 and lower electrode 102:

[0063] The glass substrate was cleaned sequentially with deionized water, isopropanol, and acetone, and then dried in an oven at 80°C. The surface of the glass substrate was then hydroxylated using oxygen plasma (O2Plasma) at a power of 50W for 1 min. 10 µL of octadecyltrichlorosilane (OTS) solution was mixed into 10 mL of n-heptane, and the glass substrate was immersed in the mixture for 1 min for modification. Afterward, it was removed and dried in an oven at 60°C for later use.

[0064] a. When the conductive network 203 is made of silver nanowires, a silver nanowire ethanol solution with a concentration of 10 mg / mL is spin-coated onto the glass substrate at a rotation speed of 2000 rpm for 20 s. After spin-coating, the solution is annealed at 100°C for 5 min and then naturally cooled to room temperature to form a silver nanowire conductive network with an island-bridge structure. The island-bridge structure is a stretchable network composed of functional units (islands) and connecting structures (bridges).

[0065] b. Dissolve 10 wt% PVA (based on the mass of water) in deionized water. After complete dissolution, add 5 wt% LiCl based on the mass of PVA and 0.5 wt% glutaraldehyde based on the mass of PVA as a crosslinking agent. React at 80°C for 2 hours. Pour the above mixed solution onto the conductive network 203. Spin coat at 7000 rpm for 30 seconds, cure at 100°C for 1 minute to form the modified functional layer 202.

[0066] c. Cover the modified functional layer 202 with an elastic base layer 201. When the material of the elastic polymer layer 101 is polydimethylsiloxane, pour the PDMS prepolymer onto the modified functional layer 202, and then spin-coat it. The spin-coating speed is 2000 rpm, the spin-coating time is 30 s, the curing temperature is 100℃, and the curing time is 5 min. As another embodiment, the material of the elastic polymer layer 101 can also be any one of thermoplastic polyurethane (TPU) or polyurethane (PU).

[0067] d. The glass substrate is separated from the flexible stacked electrode by mechanical peeling to obtain the upper electrode 101 and the lower electrode 102.

[0068] S2. Fabrication of touch-pain piezoelectric sensing component 300:

[0069] a. Add Mg-ZnO nanorods (existing Mg-doped ZnO nanorods can be used, with a magnesium doping amount of 0.5 at%) to the PDMS prepolymer. The Mg-ZnO nanorods account for 20 wt% of the PDMS. The morphology of the Mg-ZnO nanorods used in this embodiment is shown in the figure below. Figure 3 As shown, the mixture is thoroughly stirred to prepare a pain-sensing layer composite prepolymer.

[0070] b. The above-mentioned pain-sensing layer composite prepolymer is prepared on the lower electrode 102 by spin coating, with a curing temperature of 100°C and a curing time of 5 min to form the lower electrode-pain-sensing layer.

[0071] c. A common ground electrode 303, made of gold (Au), with a thickness of 200 nm, is fabricated on the aforementioned lower electrode-pain sensing layer by vapor deposition. The vapor deposition process parameters are: vacuum degree ≤ 8 × 10⁻⁶. -4 Pa, deposition rate 0.5-2 Å / s, surface roughness Ra≤20nm, sheet resistance≤5Ω / sq.

[0072] d. Add Mg-ZnO nanorods (magnesium doping amount 0.5at%) to the PDMS prepolymer, with the Mg-ZnO nanorods accounting for 40wt% of the PDMS mass. Stir the mixture thoroughly to prepare a tactile sensing layer composite prepolymer.

[0073] e. The above-mentioned tactile sensing layer composite prepolymer is prepared on the common ground electrode 303 by spin coating, with a curing temperature of 100°C and a curing time of 5 min, to form the lower electrode-tactile pain sensing piezoelectric sensing component.

[0074] S3. Device Packaging:

[0075] The upper electrode 101 and the lower electrode-touch-pain piezoelectric sensing component are bonded together by hot pressing. The hot pressing temperature is 120℃, the pressure is 0.5MPa, and the holding time is 45 seconds to form the entire device. The alignment deviation between the upper electrode 101 and the lower electrode 102 is ≤100um, and the edge overlap is ≥90%.

[0076] The final structure diagram of the biomimetic tactile pain-sensing dual-modal pressure sensor is as follows: Figure 1 As shown, when a pressure of 30 kPa is applied to the surface of the sensor, the sensor signal output is as follows: Figure 4 As shown, the tactile sensing layer signal is connected to the upper electrode and the common ground electrode, and the pain sensing layer signal is connected to the common ground electrode and the lower electrode. Since the pressure is <50kPa, the voltage output of the tactile sensing layer is 2.1V and the output of the pain sensing layer is 0.3V, which does not reach the pain threshold. At this time, the pressure state is determined to be the tactile sensing state.

[0077] Example 2

[0078] The method for fabricating a biomimetic tactile pain-sensing dual-modal flexible pressure sensor described in this embodiment includes the following steps:

[0079] S1. Fabrication of upper electrode 101 and lower electrode 102:

[0080] The glass substrate was cleaned sequentially with deionized water, isopropanol, and acetone, and then dried in an oven at 80°C. The surface of the glass substrate was hydroxylated using O2 Plasma at 50W for 1 minute. 10 µL of octadecyltrichlorosilane (OTS) solution was mixed into 10 mL of n-heptane, and the glass substrate was immersed in the mixture for 1 minute for modification. Afterward, it was removed and dried in an oven at 60°C for later use.

[0081] a. When the conductive network 203 is made of carbon nanotube conductive material, a carbon nanotube aqueous solution with a concentration of 50 mg / mL is spin-coated onto the glass substrate at a rotation speed of 800 rpm for 20 s. After spin-coating, the substrate is annealed at 80°C for 5 min and then naturally cooled to room temperature to form a carbon nanotube conductive network.

[0082] b. Dissolve 20 wt% PVA (by weight of water) in deionized water. After complete dissolution, add 7 wt% LiCl (by weight of PVA) and 1 wt% glutaraldehyde (by weight of PVA) as a crosslinking agent. React at 80°C for 2 hours. Pour the above mixed solution onto the conductive network 203. Spin coat at 2000 rpm for 30 seconds. Curing temperature is 100°C and curing time is 1 minute to form the modified functional layer 202.

[0083] c. Cover the modified functional layer 202 with an elastic base layer 201. When the material of the elastic polymer layer 101 is thermoplastic polyurethane (TPU), pour the TPU prepolymer onto the modified functional layer 202, and then spin-coat it. The spin-coating speed is 1500 rpm, the spin-coating time is 30 s, the curing temperature is 100°C, and the curing time is 5 min. As another embodiment, the material of the elastic polymer layer 101 can also be either polydimethylsiloxane (PDMS) or polyurethane (PU).

[0084] d. The glass substrate is separated from the flexible stacked electrode by mechanical peeling to obtain the upper electrode 101 and the lower electrode 102.

[0085] S2. Fabrication of touch-pain piezoelectric sensing component 300:

[0086] a. Mg-ZnO nanorods (magnesium doping amount 1 at%) were added to the TPU prepolymer, with the Mg-ZnO nanorods accounting for 5 wt% of the TPU mass. The morphology of the Mg-ZnO nanorods used in this embodiment is shown in the figure below. Figure 3 As shown, the mixture is thoroughly stirred to prepare a pain-sensing layer composite prepolymer.

[0087] b. The above-mentioned pain-sensing layer composite prepolymer is prepared on the lower electrode 102 by spin coating, with a curing temperature of 100°C and a curing time of 5 min to form the lower electrode-pain-sensing layer.

[0088] c. A common ground electrode 303, made of aluminum (Al) and 100 nm thick, is fabricated on the aforementioned lower electrode-pain sensing layer by vapor deposition. The vapor deposition process parameters are: vacuum degree ≤ 8 × 10⁻⁶. -4 Pa, deposition rate 1.5 Å / s, surface roughness Ra≤20 nm, sheet resistance≤5 Ω / sq.

[0089] d. Add Mg-ZnO nanorods (magnesium doping amount 1at%) to the TPU prepolymer, with the Mg-ZnO nanorods accounting for 25wt% of the TPU mass. Stir the mixture thoroughly to prepare a tactile sensing layer composite prepolymer.

[0090] e. The above-mentioned tactile sensing layer composite prepolymer is prepared on the common ground electrode 303 by spin coating, with a curing temperature of 100°C and a curing time of 5 min, to form a lower electrode-touch-pain piezoelectric sensing component.

[0091] S3. Device Packaging:

[0092] The upper electrode 101 and the lower electrode-touch-pain piezoelectric sensing component are bonded together by hot pressing. The hot pressing temperature is 120℃, the pressure is 0.5MPa, and the holding time is 60 seconds to form the entire device. The alignment deviation between the upper electrode 101 and the lower electrode 102 is ≤100um, and the edge overlap is ≥90%.

[0093] The final structure diagram of the biomimetic tactile pain-sensing dual-modal pressure sensor is as follows: Figure 1 As shown, when a pressure of 60 kPa is applied to the surface of the sensor, the sensor signal output is as follows: Figure 5 As shown, the tactile sensing layer signal is connected to the upper electrode and the common ground electrode, and the pain sensing layer signal is connected to the common ground electrode and the lower electrode. Since the pressure is >50kPa, the tactile sensing layer voltage output is 3V and the pain sensing layer output is 1.3V, reaching the pain threshold. At this time, the pressure state is determined to be the pain sensing state.

[0094] Example 3

[0095] This embodiment can also fabricate a biomimetic tactile pain-sensing dual-modal flexible pressure sensor array, specifically including the following steps:

[0096] S1. Fabrication of upper electrode 101 and lower electrode 102:

[0097] The glass substrate was cleaned sequentially with deionized water, isopropanol, and acetone, and then dried in an oven at 80°C. The surface of the glass substrate was hydroxylated using O2 Plasma at 50W for 1 minute. 10 µL of octadecyltrichlorosilane (OTS) solution was mixed into 10 mL of n-heptane, and the glass substrate was immersed in the mixture for 1 minute for modification. Afterward, it was removed and dried in an oven at 60°C for later use.

[0098] a. When the conductive network 203 is made of two-dimensional transition metal carbide / nitride (MXene), an MXene aqueous solution with a concentration of 30 mg / mL is spin-coated onto the glass substrate at a rotation speed of 1200 rpm for 20 s. After spin-coating, the substrate is annealed at 80°C for 5 min, allowed to cool naturally to room temperature, and then patterned using photolithography. The resulting MXene conductive network is shown below. Figure 6The upper electrode 101 and lower electrode 102 are shown in the diagram. Those skilled in the art can modify or prepare the patterned conductive network disclosed in this invention without inventive effort, based on the above-described spin coating, photolithography, and screen printing processes.

[0099] b. Dissolve 15 wt% PVA (by weight of water) in deionized water. After complete dissolution, add 15 wt% LiCl (by weight of PVA) and 0.7 wt% glutaraldehyde (by weight of PVA) as a crosslinking agent. React at 80°C for 2 hours. Pour the above mixed solution onto the conductive network 203. Spin coat at 5000 rpm for 30 seconds. Curing temperature is 100°C and curing time is 1 minute to form the modified functional layer 202.

[0100] c. Cover the modified functional layer 202 with an elastic base layer 201. When the elastic polymeric layer 101 is made of polyurethane (PU), pour the PU prepolymer onto the modified functional layer 202, and then spin-coat it. The spin-coating speed is 800 rpm, the spin-coating time is 30 s, the curing temperature is 100℃, and the curing time is 5 min. According to existing materials in the art, the elastic polymeric layer 101 can also be made of either polydimethylsiloxane (PDMS) or thermoplastic polyurethane (TPU).

[0101] d. The glass substrate is separated from the flexible stacked electrode by mechanical peeling to obtain the upper electrode 101 and the lower electrode 102.

[0102] S2. Fabrication of touch-pain piezoelectric sensing component 300:

[0103] a. Mg-ZnO nanorods (magnesium doping amount 1.5 at%) are added to the PU prepolymer, with the Mg-ZnO nanorods accounting for 15 wt% of the PU mass. The morphology of the existing Mg-ZnO nanorods used in this embodiment is shown in the figure below. Figure 3 As shown, the mixture is thoroughly stirred to prepare a pain-sensing layer composite prepolymer.

[0104] b. The above-mentioned pain-sensing layer composite prepolymer is prepared on the lower electrode 102 by screen printing, with a curing temperature of 100°C and a curing time of 5 min to form the lower electrode-pain-sensing layer.

[0105] c. A common ground electrode 303, made of silver (Ag) and 150 nm thick, is fabricated on the aforementioned lower electrode-pain sensing layer by vapor deposition. The vapor deposition process parameters are: vacuum degree ≤ 8 × 10⁻⁶. -4 Pa, deposition rate 1.5 Å / s, surface roughness Ra≤20 nm, sheet resistance≤5 Ω / sq.

[0106] d. Add Mg-ZnO nanorods (magnesium doping amount 1.5 at%) to the PU prepolymer, with the Mg-ZnO nanorods accounting for 35 wt% of the PU mass. Stir the mixture thoroughly to prepare a tactile sensing layer composite prepolymer.

[0107] e. The above-mentioned tactile sensing layer composite prepolymer is prepared on the common ground electrode 303 by screen printing, with a curing temperature of 100°C and a curing time of 5 min, to form a lower electrode-touch-pain piezoelectric sensing component.

[0108] S3. Device Packaging:

[0109] The upper electrode 101 and the lower electrode-touch-pain piezoelectric sensing component are bonded together by hot pressing. The hot pressing temperature is 120℃, the pressure is 0.2MPa, and the holding time is 60 seconds to form the entire device. The alignment deviation between the upper electrode 101 and the lower electrode 102 is ≤100um, and the edge overlap is ≥90%.

[0110] As one embodiment, the final structure of the biomimetic tactile analgesia bimodal pressure sensor array includes multiple biomimetic tactile analgesia bimodal flexible pressure sensors distributed in an array, such as... Figure 6 In this configuration, multiple upper electrodes located in the same row are connected in series to form a channel, and upper electrodes located in different rows are not connected to each other; multiple lower electrodes located in the same column are connected in series to form a channel, and lower electrodes located in different columns are not connected, wherein the connection is an electrical connection. Figure 6 , Figure 7 In the top-down view, the dashed lines represent connections between electrodes in the same column. When... Figure 7 Apply a pressure of 10 kPa at the intersection of points 3 and 6. Figure 7 As shown in the upper left figure, the voltage output of this sensor array is as follows: Figure 7 As shown in the voltage output curve of the lower middle section, channels 3 and 6 generate voltage outputs. From this, we can deduce that the force-bearing area is the intersection of channels 3 and 6. Furthermore, at this point, the voltage output of channel 6 in the pain-sensing layer is <1V, therefore, the intersection of channels 3 and 6 represents tactile perception. When... Figure 7 Apply a pressure of 50 kPa at the intersection of points 1-7. Figure 7 As shown in the upper right figure, the voltage output of this sensor array is as follows: Figure 7 As shown in the lower voltage output curve, channels 1 and 7 generate piezoelectric outputs. This indicates the force-bearing area is the intersection of channels 1 and 7. Furthermore, at this point, the voltage output of channel 7 in the pain-sensing layer is >1V, therefore the intersection of channels 1 and 7 represents pain perception. It is evident that the piezoelectric sensing layer is constructed using a row and column signal encoding mechanism with a dual-layer piezoelectric sensitive layer. The tactile sensing layer outputs signals to identify the row position of the point of application, while the pain-sensing layer outputs signals to identify the column position. The intersection of these two signals is the pressure application point. Moreover, the high-voltage threshold (e.g., ≥50kPa) of the pain-sensing layer can further distinguish pain signals.

[0111] The above embodiments are preferred embodiments of the present invention, but the embodiments of the present invention are not limited to the above embodiments. Any changes, modifications, substitutions, combinations, or simplifications made without departing from the spirit and principle of the present invention shall be considered equivalent substitutions and shall be included within the protection scope of the present invention.

Claims

1. A bionic allodynia dual-mode flexible pressure sensor, characterized in that, The device comprises an upper electrode (101), a lower electrode (102), and a touch-pain perception piezoelectric sensing component (300) between the upper electrode (101) and the lower electrode (102); the touch-pain perception piezoelectric sensing component (300) comprises a touch perception layer (301), a pain perception layer (302), and a common electrode (303) between the touch perception layer (301) and the pain perception layer (302); the touch perception layer and the pain perception layer are realized by mixing different mass percentages of magnesium-doped zinc oxide (Mg-ZnO) nanorods into an elastic polymer, wherein the content of the Mg-ZnO nanorods in the touch perception layer is greater than that in the pain perception layer; the voltage signal output by the touch perception layer due to the piezoelectric effect is collected through the upper electrode and the common electrode, and the voltage signal output by the pain perception layer due to the piezoelectric effect is collected through the common electrode and the lower electrode; the collected voltage signals are led out of the upper electrode and the lower electrode to an external processing circuit, realizing bimodal perception of touch pressure and pain pressure; in the material of the touch perception layer (301), the Mg-ZnO nanorods account for 25-40 wt% of the mass of the elastic polymer; in the material of the pain perception layer (302), the Mg-ZnO nanorods account for 5-20 wt% of the mass of the elastic polymer; the magnesium doping amount in the Mg-ZnO nanorods is 0.1-2 at%; the upper electrode and the lower electrode each comprise, from bottom to top, an elastic substrate layer (201), a modified functional layer (202), and a conductive network (203); the conductive network is embedded in the modified functional layer (202), and one side of the conductive network is completely embedded in the modified functional layer (202), and the other side is used for electrical connection with an external lead.

2. The bionic tenderness dual-mode flexible pressure sensor of claim 1, wherein, The common electrode is formed between the touch perception layer and the pain perception layer by evaporation or magnetron sputtering.

3. The bionic tenderness dual-mode flexible pressure sensor of claim 1, wherein, The material of the modified functional layer (202) comprises cross-linked polyvinyl alcohol and lithium chloride, wherein LiCl accounts for 2-15 wt% of the mass of the cross-linked polyvinyl alcohol.

4. The bionic tenderness dual-mode flexible pressure sensor of claim 1, wherein, The elastic polymer or the elastic substrate layer (201) is selected from one or more of polydimethylsiloxane, thermoplastic polyurethane, and polyurethane.

5. The bionic tenderness dual-mode flexible pressure sensor of claim 1, wherein, The conductive network (203) is an island-bridge grid structure formed by spin coating, photolithography, or screen printing using a conductive material.

6. A bionic allodynia dual-mode flexible pressure sensor array, characterized in that, The device comprises a plurality of bionic touch-pain bimodal flexible pressure sensors according to any one of claims 1-5 arranged in rows and columns, wherein a plurality of the upper electrodes (101), the lower electrodes (102), and the touch-pain perception piezoelectric sensing components (300) are arranged in the rows and columns accordingly; a plurality of upper electrodes in the same row are connected in series, and upper electrodes in different rows are not connected to each other; a plurality of lower electrodes in the same column are connected in series, and lower electrodes in different columns are not connected, and the connection is electrical connection.

7. A method for preparing the bionic allodynia dual-mode flexible pressure sensor according to any one of claims 1-5, characterized in that The device comprises: (1) preparing an upper electrode and a lower electrode; (2) Preparation of touch-pain piezoelectric sensing component: spin-coat the pain perception layer on the lower electrode, form the pain perception layer after curing; evaporate the common electrode on the pain perception layer; spin-coat the touch perception layer on the common electrode, form the touch perception layer after curing, at this time the lower electrode-pain perception layer-common electrode-touch perception layer structure is formed from bottom to top; (3) The entire device is formed by packaging the upper electrode and the lower electrode-pain perception layer-common electrode-touch perception layer structure through the hot pressing method.

8. The method of claim 7, wherein Specifically comprising: (1) Preparation of upper electrode (101) and lower electrode (102): a. Spin-coat, photoetch or screen-print the island-bridge structure of silver nanowire, carbon nanotube or MXene organic solution or aqueous solution with a concentration of 5-50 mg / mL on the glass substrate at 800-2000 rpm, and then anneal at 80-120℃ to obtain a conductive network; b. Modification of functional layer (202) embedding: spin-coat the mixed aqueous solution of cross-linked polyvinyl alcohol and lithium chloride to fill the conductive network on the glass substrate at 2000-7000 rpm; c. Elastic substrate layer (201) covering: spin-coat the solution of polydimethylsiloxane, thermoplastic polyurethane or polyurethane on the modified functional layer (202) at 800-2000 rpm, and form a solid film after solvent evaporation; mechanically peel off to obtain a flexible upper electrode or lower electrode; (2) Preparation of touch-pain piezoelectric sensing component (300): a. Pain perception layer (302): add 5-20wt% of the Mg-ZnO nanorod to the elastic polymer polydimethylsiloxane, thermoplastic polyurethane or polyurethane, and then spin-coat, screen-print or print the mixture on the lower electrode to form a 50-200μm pain perception layer after curing, together with the lower electrode to form the lower electrode-pain perception layer; b. Common electrode (303) deposition: On the pain perception layer, evaporate Au, Ag or Al under vacuum degree ≤8×10 -4 Pa, deposition rate 0.5-2 Å / s, to prepare the common electrode, forming the lower electrode-pain perception layer-common electrode; c. Touch perception layer (301): add 25-40wt% of the Mg-ZnO nanorod to the elastic polymer polydimethylsiloxane, thermoplastic polyurethane or polyurethane, and then spin-coat, screen-print or print the mixture on the common electrode to form a 50-200μm touch perception layer after curing, forming the lower electrode-pain perception layer-common electrode-touch perception layer structure; (3) Device packaging: bond the upper electrode and the lower electrode-touch-pain piezoelectric sensing component through the hot pressing method, with a positional deviation ≤100μm and an edge coincidence degree ≥90%; the hot pressing method has a temperature of 80-120℃, a pressure of 0.1-0.5 MPa, and a pressure maintaining time of 30-60 seconds.

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