Flexible bioelectric body surface dry electrode based on magnetic tunnel junction and preparation method of flexible bioelectric body surface dry electrode

Abstract

The invention discloses a flexible bioelectric surface dry electrode based on a magnetic tunnel junction and a preparation method of the flexible bioelectric surface dry electrode. The flexible bioelectric body surface dry electrode comprises a flexible substrate, and a bottom electrode lead, an insulating layer, a discrete conductive island array, an MTJ spinning electron functional layer, a microneedle biological contact layer, an interface leveling insulating layer, a confluence top electrode and a top electrode lead which are arranged in sequence. A conductive island array and top-bottom dual-electrode collaborative design is adopted: a conductive island focuses a weak bioelectric signal on a body surface, so that the weak bioelectric signal vertically passes through an MTJ functional layer, magnetic moment deflection is driven by an Oersted field excited by a bioelectric current, and in-situ amplification of the signal is realized by using an MTJ tunneling magnetoresistance effect. In combination with the low-impedance contact characteristic of the microneedle, the electrode effectively reduces motion artifacts, breaks through the bottlenecks of signal attenuation and dependence on an external amplification circuit of a traditional dry electrode, has high sensitivity, high signal-to-noise ratio and wearing comfort, and is suitable for the fields of electroencephalogram and electrocardio acquisition and wearable health monitoring.

Description

Technical Field

[0001] This invention belongs to the fields of biomedical engineering and flexible electronics technology, and in particular, it relates to a flexible bioelectric surface dry electrode based on a magnetic tunnel junction and its preparation method. Background Technology

[0002] In the field of biomedical sensing technology, high-fidelity, high signal-to-noise ratio acquisition of microvolt-level bioelectrical signals such as electroencephalograms (EEG) and electrocardiograms (ECG) is the cornerstone for advancing precision medicine, neuroscience research, and the development of next-generation brain-computer interfaces. Currently, while dry electrode technology, aimed at eliminating the constraints of conductive gel and improving wearability, has made some progress, it still faces fundamental bottlenecks that are difficult to overcome in the pursuit of higher performance and a better wearable experience.

[0003] First, the signal conduction mechanism of traditional "passive" dry electrodes has physical limitations, resulting in weak source signals. Whether it's a planar metal electrode relying on pressure contact or a microneedle structure designed to pierce the stratum corneum, its core function is limited to the "passive conduction" of bioelectrical signals. The high contact impedance at the electrode-skin interface (typically in the range of hundreds of kiloohms to several megaohms) causes severe attenuation of the source signal at the acquisition point, with amplitude loss exceeding 80%. This forces subsequent circuitry to process an extremely weak initial signal that is easily overwhelmed by noise, fundamentally limiting the improvement of the overall signal-to-noise ratio.

[0004] Secondly, the heavy reliance on external high-gain amplifier circuits brings a cascading set of challenges related to power consumption, size, and noise. To extract the severely attenuated microvolt-level signal, existing solutions must rely on external multi-stage high-gain, low-noise instrumentation amplifiers. This not only significantly increases the overall power consumption of wearable devices, limiting their battery life, but also results in a bulky system that is difficult to integrate. More importantly, the long wire connections and complex front-end amplifier circuitry themselves introduce additional environmental electromagnetic interference (such as 50 / 60Hz power frequency noise) and background noise, further degrading signal quality.

[0005] Furthermore, the application paradigm of the emerging magnetic tunnel junction (MTJ) biosensing technology is not yet compatible with the needs of bioelectrical signal acquisition. Introducing highly sensitive MTJ devices into the field of biosensing is a cutting-edge exploration. However, existing technologies mostly focus on using MTJs to detect exogenous magnetic signals. For example, biochemical analysis is achieved by detecting changes in the magnetic field generated by magnetic particles of labeled biomolecules (such as DNA and proteins), or it is directly used to measure biomagnetic fields such as magnetocardiography and magnetoencephalography. These approaches are essentially detections of "magnetism," rather than the acquisition and amplification of "electricity," and their signal sources and sensing mechanisms are fundamentally different from the requirements for measuring body surface potential.

[0006] Furthermore, recent research attempting to use MTJs for physiological electrical signal detection still hasn't solved the core physical problem of potential difference coupling with the device. Recent studies have proposed combining MTJ arrays with conductive hydrogels to simultaneously acquire electrical and magnetic signals. However, such designs typically use the MTJ as a simple magnetic sensing element connected in series in the detection circuit. The key challenge lies in the fact that bioelectrical signals collected from the body surface are essentially diffusely distributed surface potential differences with extremely low energy density. The core of the MTJ—the less than 1-nanometer-thick MgO barrier layer—relies on the directional deflection of the free layer magnetic moment for its tunneling magnetoresistance (TMR) effect. Existing structures cannot focus the diffuse surface potential difference into an effective excitation that can drive the deflection of the free layer magnetic moment. This results in most of the energy failing to effectively trigger the MTJ's Tunneling Magnetoresistance (TMR) effect, and the device sensitivity falling far short of its theoretical limit.

[0007] Furthermore, in existing dual-lead MTJ sensing technology, bias voltage stability and contact impedance are core challenges. For example, although some solutions employ an operational transconductance amplifier (OTA) negative feedback bias scheme, they do not fundamentally solve the source signal attenuation problem caused by the high contact impedance (500 kΩ) at the electrode-skin interface. Moreover, when dealing with dynamically changing contact impedance, their bias scheme may still generate bias voltage fluctuations due to the existence of potential drift paths. Summary of the Invention

[0008] The purpose of this invention is to provide a flexible bioelectric surface dry electrode based on a magnetic tunnel junction and its preparation method, so as to solve the above-mentioned technical problems existing in the prior art.

[0009] Technical solution: A flexible bioelectric surface electrode based on a magnetic tunnel junction, comprising:

[0010] Flexible substrate;

[0011] Bottom electrode leads deposited on the surface of a flexible substrate;

[0012] An insulating layer covering the surface of the flexible substrate and the bottom electrode lead area, with the reserved pad area exposed;

[0013] A discrete array of conductive islands distributed on the surface of an insulating layer;

[0014] Microneedle bio-contact layers are arranged one-to-one with the discrete conductive island array;

[0015] After the microneedle biocontact layer is temporarily protected by photoresist, an MTJ spintronic functional layer array is fabricated on the surface of a discrete conductive island array in a one-to-one correspondence.

[0016] An interface-smoothing insulating layer that covers the sidewalls of the spintronic functional layer and the non-functional region of the insulating layer of the MTJ and exposes only the surface of the free layer of the MTJ.

[0017] The bus top electrode is electrically connected to the MTJ free layer;

[0018] Top electrode lead deposited on the surface of the bus top electrode;

[0019] A composite encapsulation layer covering the microneedle sidewalls, bus top electrode, top electrode lead, and MTJ spintronic functional layer sidewalls, exposing only the top / bottom electrode pads and the microneedle tip.

[0020] Preferably, the flexible substrate material is a high-modulus polyimide film with a thickness of 200~450μm and a Young's modulus of 8~18GPa, which has excellent flexibility and can withstand repeated bending at angles of 180° and above.

[0021] Preferably, the bottom electrode lead comprises, from the surface of the flexible substrate upwards, a TiN adhesive layer, a Cu conductive layer, a TiN barrier layer, and an Au lead-out layer; the thickness of the TiN adhesive layer is 15~25 nm, the thickness of the Cu conductive layer is 450~550 nm, the thickness of the TiN barrier layer is 8~12 nm, and the thickness of the Au lead-out layer is 25~35 nm.

[0022] Preferably, the insulating layer is a Si3N4 layer with a thickness of 200~500 nm.

[0023] Preferably, the discrete conductive island array is a conductive island Au array, wherein the conductive island array is distributed in three concentric rings within a circular sensing area with a diameter of 6 mm: the first ring has a diameter of 1.5 mm and contains 4 conductive islands; the second ring has a diameter of 3.0 mm and contains 6 conductive islands; and the third ring has a diameter of 4.5 mm and contains 10 conductive islands. Each conductive island has a diameter of 200~300μm and a thickness of 1~2μm, and the center-to-center distance between adjacent conductive islands is ≥500μm to avoid signal crosstalk.

[0024] Preferably, the microneedle biocontact layer has a pointed conical protrusion structure, with each pointed conical microneedle vertically deposited at the geometric center of each conductive island in the conductive island array, and the aspect ratio of the microneedles is ≥5:1. The microneedle biocontact layer includes, from the surface of the conductive island upwards, a Ti adhesion layer, a TiN barrier layer, an Au seed layer, a Ni conductive core layer, and a Pd biocompatible layer. The thickness of the Ti adhesion layer is 2~5 nm, the thickness of the TiN barrier layer is 20~30 nm, the thickness of the Au seed layer is 80~120 nm, the thickness of the Ni conductive core layer is 1500~1900 nm, and the thickness of the Pd biocompatible layer is 80~120 nm.

[0025] Preferably, the spintronic functional layer of the MTJ comprises, from the surface of the conductive island upwards, a Ti buffer layer, an IrMn pinning layer, an FeCoB reference layer, an MgO barrier layer, an FeCoB free layer, and a TiN protective layer; the thickness of the Ti buffer layer is 1~3 nm, the thickness of the IrMn pinning layer is 8~12 nm, the thickness of the FeCoB reference layer is 1.8~2.64 nm, the thickness of the MgO barrier layer is 0.6~1.2 nm, the thickness of the FeCoB free layer is 1.6~2.2 nm, and the thickness of the TiN protective layer is 3~4 nm.

[0026] Preferably, the barrier height ΔE of the MgO tunneling barrier layer is 2.0~2.6 eV, and the interface roughness between it and the FeCoB free layer is ≤0.5 nm, ensuring a tunneling magnetoresistivity ≥180%.

[0027] Preferably, the interface planarization insulating layer is a Si3N4 layer with a thickness of 1.5~2.5μm, prepared by a low-temperature deposition process at a room temperature of 25~30℃ to avoid damaging the existing functional layer structure.

[0028] Preferably, the top electrode of the busbar is a TiN / Au composite layer structure, which includes a TiN adhesive layer and an Au conductive core layer from bottom to top; wherein, the thickness of the TiN adhesive layer is 15~25 nm, which is used to improve the bonding force with the MTJ free layer; the thickness of the Au conductive core layer is 25~35 nm, which is used to ensure that the potential of all MTJ sensing units is consistent and to realize signal busbar.

[0029] Preferably, the layered structure of the top electrode lead is the same as that of the bottom electrode lead, and forms a symmetrical structure with the bottom electrode lead to reduce signal transmission differences; the top electrode lead includes, from bottom to top, a TiN adhesive layer, a Cu conductive layer, a TiN barrier layer and an Au lead-out layer; wherein, the thickness of the TiN adhesive layer is 15~25 nm, the thickness of the Cu conductive layer is 450~550 nm, the thickness of the TiN barrier layer is 8~12 nm, and the thickness of the Au lead-out layer is 25~35 nm.

[0030] Preferably, the composite encapsulation layer comprises, from the inside out, an Al2O3 passivation layer and a CVD parylene protective layer; wherein, the Al2O3 passivation layer has a thickness of 15~25 nm to achieve atomically dense protection; and the CVD parylene protective layer has a thickness of 80~120 nm to improve flexibility and biocompatibility.

[0031] Preferably, the arrangement of the conductive island array, microneedle array, and MTJ spintronic functional layer in this invention is not limited to the above-mentioned concentric circle distributed design; various discrete array arrangements such as matrix, honeycomb, and regular hexagonal arrays are also applicable. The core technical requirement is that the conductive island, microneedle, and MTJ spintronic functional layer must achieve one-to-one vertical alignment growth to ensure that the biocurrent on the body surface, after being conducted through the microneedles and conductive island, can accurately excite the Oersted field of the corresponding MTJ spintronic functional layer and drive the deflection of the FeCoB free layer magnetic moment, thereby realizing signal acquisition and amplification through independent channels.

[0032] This solution also provides a method for fabricating a flexible bioelectric surface dry electrode based on a magnetic tunnel junction. The thin film deposition process uses an ultra-high vacuum magnetron sputtering system and includes the following steps:

[0033] S1: Flexible substrate pretreatment;

[0034] S2: Using a mask template with a corresponding integrated bottom electrode lead-port pattern, patterned bottom electrode leads are deposited on a flexible substrate; then an insulating layer is deposited over the entire surface.

[0035] S3: A discrete array of conductive islands is deposited on the surface of the insulating layer using a shared mask template of the corresponding conductive island-MTJ spin functional layer;

[0036] S4: Using a mask template with the same pattern as in step S3, a full-stack MTJ thin film is continuously deposited on the surface of the conductive island array and the insulating layer to form an MTJ sensing column array.

[0037] S5: Using a mask template corresponding to the pattern of the free layer region of the MTJ sensor column, insulating material is deposited to cover the non-free layer region of the MTJ sensor column array, thereby achieving interface smoothing.

[0038] S6: Deposit the top bus electrode on the planarized surface;

[0039] S7: A patterned top electrode lead is deposited on the bus top electrode using a mask template with a corresponding top electrode lead-port integrated pattern.

[0040] S8: Using a mask template corresponding to the pad-microneedle tip protection pattern, a composite encapsulation layer consisting of an Al2O3 passivation layer and a CVD parylene protective layer is deposited, exposing the pads and microneedle tips.

[0041] Preferably, in step S1, after cleaning the polyimide flexible substrate, oxygen plasma activation treatment is performed to make the surface water droplet contact angle ≤10°.

[0042] Preferably, the mask template corresponding to the integrated bottom electrode lead-port pattern and the mask template corresponding to the integrated top electrode lead-port pattern both include lead pattern and port pattern, the lead opening thickness direction is adapted to the total deposition thickness of the bottom electrode lead or the top electrode lead, and the pattern positioning accuracy is ≤ ±2μm.

[0043] Preferably, the mask template corresponding to the conductive island-MTJ spin functional layer uses a concentric circle pattern. The pattern consists of 20 circular openings arranged in three concentric circles in the central region of the substrate. The diameter of the openings is adapted to the size of the conductive island, and the positioning deviation is ≤±1μm.

[0044] Preferably, after depositing the conductive island array in step S3, the method further includes a step of preparing a microneedle biocontact layer, specifically: covering the conductive island array with a mask template corresponding to the microneedle tip cone pattern, wherein the mask template has a tip cone opening that corresponds one-to-one with the conductive island; sequentially depositing a Ti adhesion layer, a TiN barrier layer, and an Au seed layer using magnetron sputtering; then, based on the Au seed layer, depositing a Ni conductive core layer using a pulse electroplating process, and then depositing a Pd biocompatible layer using a DC electroplating process.

[0045] Preferably, in step S4, before depositing the MTJ film, a temporary protective mask is formed using a peelable photoresist to cover the microneedle area; the deposition of the MTJ film is carried out continuously at room temperature and a chamber background vacuum of ≥5×10-8 Pa, wherein the chamber vacuum needs to be increased to ≥5×10-10 Pa before the deposition of the MgO barrier layer, and the argon purity is ≥99.99% during the deposition process.

[0046] Preferably, after the MTJ thin film deposition is completed in step S4, a first low-temperature microwave annealing is performed. Specifically, the sample is placed in a 2.45 GHz microwave annealing chamber and annealed for 15 minutes at a power of 100 W under a vacuum of ≥5×10-8 Pa to crystallize the FeCoB layer and optimize the MgO / FeCoB interface quality.

[0047] Preferably, the mask template corresponding to the pattern of the free layer region of the MTJ sensing column has a pattern of 20 circular openings with the opening diameter adapted to the size of the MTJ free layer, used to expose only the free layer region of the MTJ, and the mask positioning accuracy is ≤±1μm.

[0048] Preferably, step S6, which involves depositing the top electrode of the busbar, is a direct deposition without a mask or template. During the deposition process, the substrate temperature is ≤60℃ to avoid the impact of high temperature on the MTJ performance.

[0049] Preferably, after the top electrode lead deposition is completed in step S7, a second low-temperature microwave annealing is performed. Specifically, the sample is placed in a 2.45 GHz microwave annealing chamber and annealed for 8 minutes at a power of 50 W under a vacuum of ≥5×10-8 Pa to eliminate lead deposition stress and improve interlayer bonding.

[0050] Preferably, the mask template corresponding to the pad-microneedle tip protection pattern includes a solid block for covering the pad area and a circular solid block for covering the microneedle tip within a range of ≤5μm, to ensure the functional integrity of the pad and the microneedle tip.

[0051] This solution also provides a method for acquiring bioelectrical signals, including the following steps:

[0052] The microneedle bio-contact layer of the dry electrode is attached to the skin of the target area on the body surface, and the microneedles penetrate the stratum corneum of the skin to achieve stable contact;

[0053] The pads of the bottom electrode lead are connected to a low-noise, low-temperature drift bandgap reference voltage source, and a precision current-limiting resistor of 100 kΩ to 1 MΩ is connected in series between the reference voltage source and the bottom electrode.

[0054] Connect the pads of the top electrode lead and the bottom electrode lead together to the signal reading circuit;

[0055] Bioelectric signals on the body surface are conducted through microneedles and conductive islands, forming a weak biocurrent flowing through the MTJ free layer. The Oersted field perpendicular to the membrane surface excited by this current directly drives the deflection of the magnetic moment of the FeCoB free layer. Based on the tunneling magnetoresistance effect, the microvolt-level bioelectric signal is converted into a resistance change signal and output to the signal reading circuit through the bus top electrode and the top electrode lead.

[0056] Beneficial effects: This solution utilizes a biomimetic microneedle array to penetrate the stratum corneum of the skin, reducing contact impedance to less than one percent of that of traditional electrodes. Furthermore, it leverages the Oersted field generated when biocurrent flows through the free layer of the magnetic tunnel junction to drive magnetic moment deflection, achieving in-situ, passive signal amplification through the magnetoresistive effect, with a gain exceeding 1000 times. This solution employs a fully cryogenic process to integrate a high-performance magnetic tunnel junction on a flexible polyimide substrate, solving key technical challenges in flexible wearable integration. It achieves high-fidelity, high signal-to-noise ratio acquisition of microvolt-level EEG and ECG signals without relying on complex external amplification circuits, resulting in extremely low power consumption. Attached Figure Description

[0057] Figure 1 This is a side view of the layered structure of the MTJ flexible bioelectric surface electrode of the present invention.

[0058] Figure 2 This is a perspective view of the concentric circular array of 20 units of the flexible bioelectric surface dry electrode based on MTJ according to the present invention.

[0059] Figure 3 This is a diagram of the bioelectric signal-magnetic field-magnetic moment deflection coupling model of the present invention.

[0060] Figure 4 This is a schematic diagram of the electrode flexibility bending performance test of the present invention.

[0061] Figure 5 This is a schematic diagram of the electrode preparation process of the present invention. Detailed Implementation

[0062] To better illustrate the technical solution of the present invention, the following will be combined with... Figures 1 to 5 The present invention will be described in detail below with reference to specific embodiments. It should be noted that in the description of the present invention, the terms "upper" and "lower," etc., indicate the directional relationship based on the layered structural relationship shown in the accompanying drawings, and are used only for simplification and not to limit the actual orientation. The term "connection" includes direct electrical connection or indirect connection through an intermediate layer. The technical solutions are described in detail below with reference to embodiments, but the embodiments do not constitute a limitation on the scope of protection. In the absence of conflict, the embodiments and features in the embodiments of the present invention can be combined with each other. The technical solutions in the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are only some embodiments of the present invention, not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of the present invention.

[0063] Example 1

[0064] This embodiment provides an overall framework for a bioelectric surface electrode based on a magnetic tunnel junction (MTJ) structure. Specifically, the bioelectric surface electrode of this embodiment includes (e.g.) Figure 1 (as shown)

[0065] The flexible substrate is a high-modulus polyimide material.

[0066] Bottom electrode leads, disposed on the surface of the flexible substrate, are used to provide a stable potential reference, construct the bottom path of the signal detection closed loop, and serve as external connection pads. The bottom electrode pads can be connected to a low-noise, low-temperature drift precision DC voltage source, providing a stable DC bias voltage with a driving capability of several milliamps and precise adjustable range of 0.5~1 V. The bottom electrode pads need to be connected to a 100 kΩ~1 MΩ precision current-limiting resistor to provide a stable static operating point for the MTJ device, so that the free layer magnetic moment is in a high-sensitivity operating point that is easy to deflect, ensuring that the Oersted field generated by weak biocurrents can efficiently drive the deflection of the free layer magnetic moment, and significantly improving the response sensitivity of the MTJ device to bioelectric signals.

[0067] An insulating layer is disposed on the surface of the flexible substrate and in the lead area of ​​the signal output terminal, with exposed pad areas reserved for subsequent external circuitry. This insulating layer forms a reliable electrical isolation barrier between the bottom electrode and the subsequently deployed conductive islands, completely blocking the leakage path of the DC bias voltage of the bottom electrode. This insulating layer does not function as a capacitor dielectric; it only needs to meet the requirements of pinhole-free, low-stress micro / nano fabrication processes, providing a stable electrical isolation environment for the direct signal path between the conductive island and the MTJ free layer. The potential difference at different points on the body surface forms a driving electric field, prompting electrons to form a direct closed-loop path through the microneedle-conductive island-MTJ free layer-capacitor top electrode. Based on this, the low-frequency potential difference on the body surface is directly converted into a weak biocurrent flowing through the MTJ free layer. This current generates a local induced magnetic field perpendicular to the membrane surface through the Oersted effect, efficiently driving the deflection of the free layer's magnetic moment, and then achieving precise conversion of the bioelectrical signal into a resistance change signal through the tunneling magnetoresistance (TMR) effect.

[0068] A discrete array of conductive islands, arranged in a concentric circle pattern on the surface of the insulating layer, maintains strict electrical isolation from the patterned bottom electrode leads, and preserves a complete insulating layer structure between the conductive islands and the bottom electrode. This concentric array structure reduces the area for collecting surface potential differences, minimizes current diffusion loss, and achieves high-density convergence of biocurrents. Since there is no direct electrical contact between the two, it efficiently converges diffuse biocurrent signals from different sites on the skin surface, such as those from electroencephalography (EEG) (frequency range 0.5–100 Hz). This converts the surface potential difference into a weak electronic current flowing through the conductive islands and precisely delivers it to the corresponding MTJ functional layer above the array, providing a stable biocurrent input for in-situ signal amplification based on the tunneling magnetoresistance effect of the MTJ structure.

[0069] The microneedle bio-contact layer is disposed on the surface of the conductive island array. The conical tips of the microneedles can penetrate the stratum corneum of the skin, significantly reducing the contact impedance at the skin-electrode interface. Ionic currents within the skin are transmitted to the conductive islands below through the conductive core layer of the microneedles. Through the double-layer reaction at the electrode-skin interface, the efficient conversion of ionic current to electronic current is completed, providing a stable carrier for subsequent biocurrent transmission. The insulating layer between the conductive islands and the bottom electrode only serves the function of reliable electrical isolation, completely blocking DC bias leakage, allowing electronic current to flow directly through the MTJ free layer and generate an induced magnetic field, driving the magnetic moment deflection of the free layer. Through the TMR effect, the in-situ conversion of biocurrent signals into resistance change signals is realized.

[0070] The MTJ spintronic functional layer is fabricated on its surface in a one-to-one correspondence with the discrete conductive island array. It is a magnetic tunnel junction (MTJ) structure. The free layer of this structure has excellent conductivity and can serve as a path for biocurrents. When the converged biocurrents flow through the free layer, a local induced magnetic field perpendicular to the membrane surface is generated based on the Oersted effect. The MTJ free layer is designed as a perpendicular magnetic anisotropy (PMA) structure, which causes the magnetic field to drive the magnetic moment of the free layer to deflect, changing the angle between it and the magnetic moment of the reference layer. Then, based on the tunneling magnetoresistance (TMR) effect, the weak biocurrent signal is accurately converted into a significant resistance change signal.

[0071] An interface-smoothing insulating layer is prepared after the MTJ array is fabricated and before the bus top electrode is deposited. This insulating layer covers the non-MTJ functional area of ​​the flexible substrate and the sidewalls of the MTJ functional layer, exposing only the top area of ​​the MTJ free layer. This eliminates the surface steps and morphological undulations caused by multilayer thin film deposition, providing a smooth substrate for the uniform deposition of the subsequent bus top electrode, while avoiding unintended electrical contact between the non-free layer area and the top electrode.

[0072] The top bus electrode is a continuous metal layer deposited on the planarized insulating layer and the exposed MTJ free layer. Twenty MTJ units are connected in parallel to the top bus electrode. This can collect the resistance change signals generated by the biocurrent of each unit, which is equivalent to increasing the output signal current. It also helps to average device noise and improve the overall signal-to-noise ratio. Furthermore, it is completely isolated from the conductive island and microneedle structure below, ensuring that all biocurrent signals flowing through each MTJ free layer are collected and exported by the top electrode, thus ensuring the accuracy and integrity of signal collection.

[0073] The top electrode lead, positioned on the flat area of ​​the bus top electrode, is used to export the bioelectrical signal amplified by the tunneling magnetoresistance effect. This top electrode lead is fabricated using an integrated top / bottom electrode lead-port mask, eliminating the need for direct contact with the highly protruding microneedles. It only needs to connect to the flat area of ​​the bus top electrode to achieve signal aggregation and export from 20 MTJ units, completely avoiding interference from microneedle height differences on lead deposition. The top and bottom electrode leads together constitute the signal output terminal, with the top electrode lead serving as the forward path and the bottom electrode lead as the potential reference path. This signal output terminal includes composite top and bottom metal leads and pads. The patterned top and bottom electrode leads are mutually insulated, forming an independent biocurrent detection closed loop for each sensing unit, ensuring the stability and anti-interference capabilities of signal acquisition.

[0074] The electrode also includes an encapsulation layer that covers the microneedle sidewalls, the surface of the bus top electrode, the surface of the top electrode lead, and the exposed MTJ functional layer sidewalls. A pad-microneedle tip protection mask is used to fully expose the pad layer at the ends of the top / bottom electrodes and the tip area of ​​the microneedle (diameter ≤ 5 μm), ensuring that the microneedle tip can penetrate the stratum corneum of the skin to establish a low-impedance contact path, while forming a dense protection for the core functional layer of the device, improving the biocompatibility and long-term wear stability of the electrode.

[0075] Preferably, the flexible substrate is a high-modulus polyimide film with a thickness controlled between 200 and 450 μm and a Young's modulus between 8 and 18 GPa. When the thickness is ≤200 μm, the flexibility is excellent but the mechanical support is insufficient; when the thickness is ≥450 μm, the support is strong but the bending performance decreases. The 200-450 μm range can balance the requirements of flexible wearables and the substrate stiffness requirements of micro-nano processing technology, and has a bending performance of ≥180°, which can adapt to the skin fit requirements and micro-nano processing technology requirements in wearable scenarios.

[0076] Preferably, the bottom electrode lead structure comprises, from bottom to top (towards skin contact), a TiN adhesion layer, a Cu conductive layer, a TiN barrier layer, and an Au lead-out layer. The TiN adhesion layer has a thickness of 15–25 nm, the Cu conductive layer has a thickness of 450–550 nm, the TiN barrier layer has a thickness of 8–12 nm, and the Au lead-out layer has a thickness of 25–35 nm.

[0077] Preferably, the insulating layer is made of high resistivity Si3N4 (resistivity > 10). 16 The thickness (Ω·cm) is designed to be 200~500 nm. This thickness achieves reliable electrical isolation between the bottom electrode leads and the conductive island array, completely blocking the leakage path of the DC bias voltage of the bottom electrode, while avoiding the increase in modulus burden on the flexible substrate due to excessive insulation layer thickness, thus adapting to the bending requirements of wearable scenarios. Its high resistivity characteristics can completely isolate the conductive island array from the bottom electrode leads below both physically and electrically, providing a stable electrical environment for the direct signal path of the "conductive island-MTJ free layer".

[0078] Preferably, the discrete conductive island array is a cylindrical array composed of high-purity Au. The array is arranged as follows: within a circular sensing region with a diameter of 6 mm above the insulating layer, three concentric rings are defined. From the inside out, the first ring has a diameter of 1.5 mm and uniformly distributes 4 conductive islands; the second ring has a diameter of 3.0 mm and uniformly distributes 6 conductive islands; the third ring has a diameter of 4.5 mm and uniformly distributes 10 conductive islands; a total of 20 conductive islands constitute the array. Each Au conductive island has a diameter of 200~300 μm and a thickness of 1~2 μm. The center-to-center distance between adjacent conductive islands is ≥500 μm to ensure the conductivity of the Au conductive layer, while forming a good step coverage with the Si3N4 insulating layer, reducing interface stress.

[0079] Preferably, the microneedles are arranged in a three-layer concentric circle distributed pattern, with 20 microneedles evenly distributed in an "inner layer-middle layer-outer layer" configuration, completely overlapping with the concentric circle array of the conductive island below (e.g., Figure 2 (As shown in the diagram). The inner concentric circle has a diameter of 1.5 mm and contains 4 microneedles evenly arranged; the middle concentric circle has a diameter of 3.0 mm and contains 6 microneedles evenly arranged; the outer concentric circle has a diameter of 4.5 mm and contains 10 microneedles evenly arranged. The spacing between the three concentric circles is 1.5 mm. Individual microneedle dimensions: The microneedle has a pointed, conical protrusion structure with an aspect ratio ≥ 5:1 to ensure its penetration ability into the stratum corneum. The tip diameter is smaller than the base width, forming a sharp contact end that can penetrate the stratum corneum, reducing contact resistance and improving current convergence efficiency.

[0080] Preferably, the microneedles employ a multilayer composite structure, comprising, from bottom to top, the following layers from the Au conductive island: a Ti adhesion layer (2-5 nm thick), a TiN barrier layer (20-30 nm thick), an Au seed layer (80-120 nm thick), a Ni conductive core layer (1500-1900 nm thick), and a Pd biocompatible layer (80-120 nm thick). The materials and thicknesses of each layer are adapted to ultra-high vacuum magnetron sputtering technology to ensure conductivity and biocompatibility.

[0081] Preferably, the MTJ spintronic functional layer is a magnetic tunnel junction (MTJ) structure, which is vertically distributed in a 1:1 ratio with the conductive island array. Each MTJ structure is independently located in the region directly above its corresponding conductive island, and its diameter is 200 μm. Before fabrication, a special spintronic functional layer mask precisely matched to the conductive island array is used for patterning to ensure that the deposition position and range of the MTJ structure are precisely controllable. The MTJ structure, from bottom to top (pointing towards skin contact direction) from the Au conductive island, sequentially includes: a Ti buffer layer (thickness 1~3 nm), an IrMn pinning layer (thickness 8~12 nm), an FeCoB reference layer (thickness 1.8~2.64 nm), an MgO barrier layer (thickness 0.6~1.2 nm), an FeCoB free layer (thickness 1.6~2.2 nm), and a TiN protective layer (thickness 3~4 nm).

[0082] Preferably, the interface smoothing insulating layer is made of high-density Si3N4 with a thickness of 1.5~2.5 μm. This thickness is much greater than that of the underlying electrical isolation insulating layer, primarily because its function is to smooth the surface morphology and protect against unintended electrical contacts: the 1.5~2.5 μm thickness is sufficient to completely fill the step difference between the MTJ functional layer and the flexible substrate, eliminating surface undulations caused by multilayer thin film deposition, providing a smooth and continuous support surface for the uniform deposition of the subsequent bus top electrode, and preventing the top electrode from breaking or experiencing localized short circuits due to surface irregularities; at the same time, this thickness ensures complete coverage of the sidewalls of the MTJ functional layer, completely isolating the non-free layer region from electrical contact with the bus top electrode, ensuring the uniqueness of the signal transmission path.

[0083] Preferably, the top electrode of the bus adopts a TiN / Au composite layer structure, which includes, from bottom to top, the following layers from the interface planarization insulating layer and the MTJ free layer: a TiN adhesive layer (thickness of 15~25 nm) and an Au conductive core layer (thickness of 25~35 nm). This composite layer can realize the electrical parallel connection of 20 MTJ free layers, ensuring the efficient collection and transmission of resistance change signals generated by each MTJ unit, and improving the overall detection sensitivity and response speed.

[0084] Preferably, the top electrode lead is positioned above the bus top electrode and is patterned using the same integrated top / bottom electrode lead-port mask as the bottom electrode lead, ensuring precise alignment of their wiring paths, spacing, and port positions. The other end of the top electrode lead extends to the substrate edge pad, enabling the aggregation and export of signals from 20 MTJ units. The lead maintains a safe distance from the high-protrusion micropins, preventing contact and short circuits. The top electrode lead adopts the same structure as the bottom electrode lead, comprising, from bottom to top: a TiN adhesion layer, a Cu conductive layer, a TiN barrier layer, and an Au lead-out layer. The TiN adhesion layer has a thickness of 15–25 nm, the Cu conductive layer has a thickness of 450–550 nm, the TiN barrier layer has a thickness of 8–12 nm, and the Au lead-out layer has a thickness of 25–35 nm.

[0085] Preferably, the encapsulation layer is a composite insulating structure, comprising, from the surface of the functional layer of the device outwards, an Al2O3 passivation layer (15-25 nm thick) deposited by atomic layer deposition (ALD) and a CVD parylene protective layer (80-120 nm thick); wherein the Al2O3 passivation layer is deposited by atomic layer deposition and the parylene protective layer is deposited by chemical vapor deposition, and the two work together to achieve non-damaging dense protection, flexible bending tolerance and improved biocompatibility of the device.

[0086] Preferably, before fabricating each functional layer of the device, a corresponding dedicated mask should be used for precise masking to ensure the forming accuracy and process compatibility of each layer structure.

[0087] This invention also provides a method for fabricating a bioelectric surface electrode based on a magnetic tunnel junction (MTJ) structure. Except for the electroplating process used for the metal core layer and biocompatible layer of the microneedles, all other thin film deposition processes employ an ultra-high vacuum magnetron sputtering system. After deposition of some functional layers, a low-temperature microwave annealing process can be used, with an annealing microwave power of 100 W and an annealing time of 15 min. Specifically, the method includes the following steps:

[0088] S1: Flexible substrate pretreatment, the polyimide flexible substrate is cleaned and activated by oxygen plasma to improve the adhesion of the substrate surface;

[0089] S2: Deposition of bottom electrode leads and Si3N4 insulating layer. A top / bottom electrode lead-port integrated mask is used to cover the pretreated flexible substrate, and patterned bottom electrode leads are formed by ultra-high vacuum magnetron sputtering. Then, while maintaining a vacuum environment, Si3N4 insulating layer is deposited on the entire surface of the substrate by radio frequency reactive magnetron sputtering.

[0090] S3: Discrete Au conductive island array fabrication: A concentric circle mask shared by conductive islands and MTJ spin functional layers is used to cover the surface of the Si3N4 insulating layer. Au thin film is deposited in the area defined by the mask by ultra-high vacuum magnetron sputtering. After deposition, the mask is removed to directly form an interlocking Au conductive island array on the surface of the Si3N4 insulating layer.

[0091] S4: The conductive island array is covered with a mask template corresponding to the microneedle tip cone pattern, and Ti adhesion layer, TiN barrier layer and Au seed layer are sputtered and deposited in sequence. Then, the microneedle morphology is defined by a tip cone photoresist mask. Then, a Ni conductive core layer is pulsed and a Pd biocompatible layer is DC electroplated based on the Au seed layer to obtain the microneedle biocontact layer.

[0092] S5: Fabrication of MTJ spintronic functional layer array. The same mask as the concentric circle mask shared by the conductive island-MTJ spin functional layer in step S3 is used to cover the Au conductive island array and the surface of the Si3N4 insulating layer. The full stack of MTJ thin film is continuously deposited by magnetron sputtering at room temperature and ultra-high vacuum. After deposition, the mask is removed to form an independent MTJ sensing pillar array corresponding to each Au conductive island.

[0093] S6: The interface planarization insulating layer is prepared by using an MTJ cell protection-interface planarization insulating layer mask to cover the MTJ sensing pillar array and the substrate surface. The insulating layer film is deposited by ultra-high vacuum magnetron sputtering, exposing only the MTJ sensing pillar area, while the rest of the area is filled and covered by the insulating layer. After the deposition is completed, the mask is removed to achieve the interface planarization treatment of the device surface.

[0094] S7: The top bus electrode is fabricated without the aid of a mask. TiN adhesion layer and Au conductive core layer are deposited sequentially on the insulating layer and the exposed MTJ sensing column surface after the interface is flattened, forming a TiN / Au composite top bus electrode, which realizes the electrical parallel connection of multiple MTJ sensing columns.

[0095] S8: Top electrode lead fabrication: The same top / bottom electrode lead-port integrated mask as the bottom electrode lead is used to cover the surface of the bus top electrode. TiN adhesion layer, Cu conductive layer, TiN diffusion barrier layer and Au lead-out layer are sequentially deposited by ultra-high vacuum magnetron sputtering to form the top electrode lead. After deposition, the mask is removed.

[0096] S9: Encapsulation layer deposition

[0097] A pad-microneedle tip protective mask is used to cover the device surface. First, an Al2O3 passivation layer is deposited using atomic layer deposition (ALD). Then, without removing the mask, a parylene protective layer is deposited using chemical vapor deposition (CVD). After deposition, the mask is removed.

[0098] Compared with the prior art, the present invention has the following advantages:

[0099] This invention uses a flexible polyimide substrate, which ensures excellent bendability of the device while also possessing a certain structural rigidity. It can closely conform to the complex curved surface of the human body and support the stable shape of the microneedle array and MTJ spintronic functional layer, avoiding electrode failure caused by excessive substrate deformation. At the same time, the Al2O3 / parylene composite encapsulation layer can effectively protect the device from external environmental corrosion and extend the device's service life.

[0100] Furthermore, this invention employs a microneedle array as the bio-contact structure. The conical tips of the microneedles can directly pierce the high-impedance barrier of the stratum corneum of the skin, forming a stable contact with the active tissue of the epidermis, significantly reducing the contact impedance at the skin-electrode interface. Compared with traditional planar electrodes, the microneedle structure of this invention can more efficiently convert the ionic current in the skin into the electronic current flowing through the MTJ free layer, reducing signal attenuation loss at the interface and acquiring bioelectrical signals with a higher signal-to-noise ratio. This provides a high-quality original current basis for subsequent signal conversion based on the Oersted effect and tunneling magnetoresistance effect.

[0101] Furthermore, the essence of surface bioelectrical signals (EEG, ECG, etc.) is the potential distribution difference on the body surface generated by cellular electrical activity. After the microneedles contact the skin, the physical quantity directly collected is the potential difference between the skin and the electrode. The focusing effect of the discrete conductive island array is to converge the diffuse potential difference in different regions of the body surface through the concentric circle array structure, forming a stable potential gradient. This potential gradient drives the ionic current in the skin to flow to the microneedles, and is converted into an electronic current Ibio flowing through the MTJ free layer through the double-layer reaction. The external circuit applies a preset stable constant voltage excitation to the bottom electrode and connects it in series with a 100 kΩ~1 MΩ precision current-limiting resistor to provide a stable bias path for the biocurrent, so that the MTJ free layer works in the high-sensitivity range of magnetic moment deflection.

[0102] When electron current flows through the free layer of MTJ (such as...) Figure 3As shown in the figure, a local induced magnetic field H perpendicular to the membrane surface is generated based on the Oersted effect. The magnetic field strength is proportional to the biocurrent H ∝ Ibio. This induced magnetic field drives the free layer magnetic moment to deflect, changing the angle θ between it and the reference layer magnetic moment, thereby inducing a dynamic change in the tunneling magnetoresistance of the MTJ. The rate of resistance change satisfies the quantitative relationship ΔR / R = TMR0·(1-cosθ). Under the small-signal approximation, H ∝ I, Ibio∝ U, and the magnetic moment deflection angle satisfies the simplified linear relationship θ ≈ k·U (k is the equivalent coupling coefficient). At the same time, using the approximate formula 1-cosθ ≈ θ² / 2, the final MTJ resistance change rate can be simplified to: ΔR / R ≈ (TMR0·(k·U)²) / 2; where TMR0 is the intrinsic magnetoresistance ratio of the junction, and k is the coupling coefficient of potential-current-magnetic field. For an MTJ with TMR0=200% and k=2.5 rad / μV, an 8 μV input potential difference signal can be converted into a biocurrent of considerable strength, thereby generating an induced magnetic field sufficient to drive the magnetic moment deflection, ultimately producing a relative resistance change of approximately 400%. A simple readout circuit (such as a constant current excitation circuit) can then achieve an equivalent voltage gain on the order of hundreds, thus enabling in-situ, efficient amplification of weak bioelectrical signals. In practical applications, due to the roughness of the MTJ barrier layer interface, the bending stress of the flexible substrate, and external environmental noise, the actual value of the resistance change rate will decrease by 10%–20%, and the equivalent voltage gain will also decrease accordingly. However, this decrease is within a controllable range and will not change the core signal amplification characteristics of the device, still enabling effective detection and amplification of weak bioelectrical signals.

[0103] Furthermore, the MTJ functional layers are fabricated on a flexible substrate using a fully low-temperature process. This invention precisely controls the thickness of each functional layer of the MTJ structure, combining ultra-high vacuum magnetron sputtering deposition with a customized microwave annealing process: the deposited MTJ structure is transferred to a 2.45 GHz microwave annealing chamber, and the chamber is evacuated to a vacuum level ≥ 5 × 10⁻⁶. -8 Pa, without additional heating devices in the annealing chamber, followed by a microwave frequency of 2.45 GHz continuous wave mode, an output power of 100 W, and an annealing time of 15 min, can precisely ensure that the barrier height ΔE of the MgO tunneling barrier layer is between 2.0 and 2.6 eV. At the same time, it forms an atomically flat interface (interface roughness ≤ 0.5 nm) between the barrier layer and the free layer, minimizing interface scattering and ensuring the efficiency and stability of magnetic moment deflection-resistance conversion under Oersted field drive. More importantly, it avoids the thermal damage to the flexible substrate caused by traditional high-temperature annealing, successfully solving the integration problem of MTJ structure and flexible substrate in wearable scenarios such as EEG, and is compatible with the stringent requirements of flexible wearable devices for substrate heat resistance. It provides reliable device structure support for the efficient coupling of biocurrent-Oersted field-magnetic moment response on the body surface.

[0104] Furthermore, this invention designs a synergistic structure of conductive island array and top and bottom dual electrodes: it can efficiently focus the weak diffuse potential difference on the body surface and convert it into a high-density biological tunneling current that vertically passes through the nanoscale MgO barrier layer. Relying on the Oersted field excited by the current, it precisely drives the deflection of the free layer magnetic moment, which is then converted into a detectable change in tunneling resistance, greatly improving the signal transmission efficiency and response speed, and laying a good foundation for subsequent signal processing.

[0105] Furthermore, the bioelectric surface dry electrode of this embodiment adopts a top and bottom dual electrode lead structure design, which can efficiently converge the weak potential difference diffused on the body surface and force it into a high-density bio-tunneling current that vertically passes through the nanoscale MgO barrier layer. The interface interaction between the FeCoB free layer and the MgO barrier layer can induce strong vertical magnetic anisotropy (PMA), making the free layer magnetic moment perpendicular to the membrane surface, which greatly improves the response sensitivity to the vertical induced magnetic field. The Oersted field excited by this current can directly drive the deflection of the MTJ free layer magnetic moment. Relying on the magnetoresistance effect of the MTJ device itself, only an external conventional constant voltage source is needed to complete the in-situ conversion and amplification of the potential difference to the Oersted field-magnetic moment deflection-resistance signal, without relying on specific external circuit components, realizing a simplified configuration of the core function. Meanwhile, this structure is not limited to high-precision bioelectrical acquisition scenarios. In a preferred embodiment, performance is further optimized through a synergistic design of a stable potential reference supply and differential signal reading: specifically, a low-noise, low-temperature drift (≤8 ppm / ℃) bandgap reference voltage source can be connected externally to the bottom electrode pad, and a 100 kΩ~1 MΩ precision current-limiting resistor can be connected in series to provide a stable excitation bias for the MTJ device, ensuring the linear coupling relationship between the Oersted field strength and the biocurrent; the bandgap reference voltage source can be a REF5050 model, with a low-frequency noise spectral density of less than 1 μV / √Hz. The top and bottom electrodes are connected together to a differential instrumentation amplifier (such as AD8221) with a high common-mode rejection ratio (CMRR > 100 dB @ 50 Hz) to perform secondary filtering and amplification on the output resistance change signal. This optional scheme ensures that the inherent noise of the bias voltage is much smaller than that of the bioelectric signal, and avoids the amplifier's background noise from drowning out the signal amplified by the MTJ Ørsted field drive. Ultimately, it achieves the optimization target of signal-to-noise ratio (SNR) ≥ 42 dB at the system level, effectively solving the signal stability problem of the MTJ structure in high-precision scenarios.

[0106] Therefore, this invention systematically solves the problems of large signal attenuation of traditional dry electrode sources, noise and power consumption caused by reliance on high-gain external amplifier circuits, as well as the core challenges of MTJ integration on flexible substrates and bioelectrical acquisition adaptability, achieving a breakthrough in high-fidelity, high signal-to-noise ratio acquisition of microvolt-level bioelectrical signals while ensuring wearable comfort.

[0107] Example 2

[0108] The following will be discussed in conjunction with the accompanying drawings (such as...). Figure 5 The embodiments of the present invention will be described in detail below (as shown). It should be noted that the embodiments are intended to enable those skilled in the art to better understand the present invention, and not to limit its scope of protection. This embodiment provides a flexible bioelectric surface dry electrode (such as...) based on a magnetic tunnel junction (MTJ) structure. Figure 1 As shown, its layered structure from bottom to top (in the direction of skin contact) includes:

[0109] Flexible substrate: Utilizing a high-modulus polyimide (PI) film with a thickness of 300 μm and a Young's modulus of 10 GPa, it exhibits excellent flexibility and can withstand repeated bending at angles of 180° and above (e.g., ...). Figure 4 (As shown).

[0110] Signal output terminal: Located on the surface of a flexible substrate, it includes two parallel composite metal leads and square pads (2.0 mm × 2.0 mm) at their ends. The leads employ a multilayer structure to ensure low resistance and good interface characteristics, consisting of, from bottom to top: a 20 nm TiN adhesion layer, a 500 nm Cu conductive layer, a 10 nm TiN diffusion barrier layer, and a 30 nm Au lead-out layer.

[0111] Insulating layer: disposed on the surface of the flexible substrate and the lead area of ​​the signal output terminal, with the pad area reserved for subsequent external circuits. The insulating layer is made of 300nm silicon nitride (Si3N4).

[0112] Discrete conductive island array: Deployed within a circular sensing area with a diameter of 6 mm, using a three-layer concentric ring arrangement. From the inside out, the first ring has a diameter of 1.5 mm and evenly distributes 4 conductive islands; the second ring has a diameter of 3.0 mm and evenly distributes 6 conductive islands; the third ring has a diameter of 4.5 mm and evenly distributes 10 conductive islands; the three layers of rings together form an array structure of 20 conductive islands. This array is made of Au material and fabricated using a pre-fabricated concentric circular array mask. The width of each conductive island is 200 μm and the height is 2 μm.

[0113] Microneedle biocontact layer: A three-layer concentric circular distributed arrangement is used, fabricated with an independent photomask. Twenty microneedles are evenly distributed in an "inner layer-middle layer-outer layer" configuration, perfectly aligned with the concentric array of conductive islands below. From the inside out, the microneedles consist of a conductive core layer (pulse-plated Ni, approximately 1700 nm) and a biocompatible surface layer (DC-plated Pd, approximately 100 nm). A metal seed layer is also provided between the conductive core layer and the conductive islands. This metal seed layer, from bottom to top, consists of a 3 nm Ti adhesion layer, a 25 nm TiN barrier layer, and a 100 nm Au seed layer.

[0114] Spintronic functional layer: Located directly above each conductive island array, this layer is fabricated using the same photomask (concentric circle array photomask) as the conductive islands, forming a precise spatial alignment with the conductive island array below. This layer is a magnetic tunnel junction (MTJ) structure, which is the core for in-situ amplification of bioelectrical signals. Its specific stacking order from bottom to top (pointing towards skin contact) is as follows: 1.5 nm Ti buffer layer, 10 nm thick IrMn pinning layer (used to provide a fixed exchange bias field), 2.0 nm thick FeCoB reference layer (whose magnetic moment is fixed by the pinning layer), 0.8 nm thick MgO tunneling barrier layer (providing a tunneling barrier for electrons), 1.6 nm thick FeCoB free layer, and a 4 nm thick TiN protective layer above the free layer (used to prevent oxidation of the core functional layer).

[0115] Interface planarization insulating layer: This layer is prepared after the MTJ array is fabricated and before the top bus electrode is deposited, covering the surface of the MTJ array and the free layer. This insulating layer uses Si3N4 material, which is the same as the bottom insulating layer, and its thickness is controlled to 2 μm. It is formed by photolithography-sputtering process using a pre-prepared interface planarization insulating layer mask.

[0116] The top bus electrode is directly deposited on the surface of the interface-planarized insulating layer and the exposed MTJ sensing pillar, employing a TiN / Au composite layer structure. Starting from the interface-planarized insulating layer and the MTJ structural surface, along the direction pointing towards skin contact, the composite layer consists of a 20 nm TiN adhesion layer and a 30 nm Au conductive bus layer, without the need for a mask.

[0117] Top electrode lead fabrication: A mask identical to that used for the bottom electrode leads is applied to the bus layer surface to ensure structural consistency and alignment accuracy between the top and bottom electrode leads. The top electrode lead employs the same multilayer composite structure as the bottom electrode lead, consisting of a 20 nm TiN adhesion layer, a 500 nm Cu conductive layer, a 10 nm TiN barrier layer, and a 30 nm Au lead-out layer from bottom to top. This achieves low-resistance, high-stability signal transmission performance that matches the bottom electrode leads.

[0118] Encapsulation layer: A pad-microneedle tip protective mask is used to cover the device surface, creating exposed areas only in the pad and microneedle tip regions. All other functional layer areas are completely covered by this encapsulation layer. This encapsulation layer is a composite insulating structure, prepared using an atomic layer deposition (ALD) process to deposit an Al2O3 + CVD parylene dual-layer encapsulation process. From the surface of the functional layer of the device outwards, it sequentially includes a 20 nm ALD Al2O3 passivation layer and a 100 nm CVD parylene protective layer.

[0119] The fabrication method of the above-mentioned bioelectric signal acquisition electrode based on spin electron layer and its magnetoresistive amplification method is as follows:

[0120] (1) Flexible substrate pretreatment

[0121] A high-modulus polyimide (PI) film with a thickness of 300 μm and a Young's modulus of 10 GPa is used, exhibiting excellent flexibility and withstanding repeated bending at angles of 180° and above. The film is cut into circular substrates with a diameter of 10 mm using a precision cutter or UV laser cutter (avoiding thermal damage to the substrate). The substrates are then ultrasonically cleaned sequentially in acetone, anhydrous ethanol, and deionized water for 10 minutes each (power 100 W, frequency 40 kHz). After each cleaning, the edges of the substrate are handled with clean tweezers, avoiding contact with the effective area, and immediately purged from the sides with a high-purity nitrogen gun (purity ≥ 99.99%) to remove surface droplets. The cleaned substrates are then placed in a clean glass petri dish and dried in an 80°C oven for 30 minutes to ensure complete removal of moisture and residual solvents. Key Control Points: To enhance subsequent film adhesion, the dried substrates are immediately placed in the sample chamber of an oxygen plasma treatment machine for immediate oxygen plasma surface activation treatment (oxygen flow rate 40 sccm, chamber pressure 30 Pa, RF power 120 W, time 5 minutes). Within 5 minutes after treatment, 1-2 substrates are randomly selected and tested using a contact angle meter. A 2 μL droplet of deionized water is placed on the substrate surface, and its static contact angle is measured. Acceptance Criterion: Water droplet contact angle ≤ 10°. Handling of Non-Acceptance: If the contact angle > 10°, it indicates insufficient surface activation. The substrate can be reinserted into the plasma treatment machine, the treatment time extended by 1-2 minutes, or the RF power finely adjusted by ±10 W, and the test repeated until the standard is met.

[0122] In a preferred embodiment, due to the high surface energy and strong activity after plasma treatment, the substrates readily adsorb dust and organic matter from the air. Substrates that pass inspection must proceed to the next step (coating in signal output preparation) within 30 minutes. The transfer process requires the use of cleanroom fixtures and must be conducted entirely within a Class 100 laminar flow hood to prevent dust particles from adsorbing onto the substrate surface. The entire pretreatment process should be carried out in a cleanroom or laminar flow hood to minimize particulate contamination. When handling organic solvents such as acetone, it must be done in a fume hood, and appropriate personal protective equipment must be worn.

[0123] (2) Bottom electrode lead fabrication process

[0124] Fabrication of the integrated top / bottom electrode lead-port mask: A high-precision photolithography process was used to fabricate the integrated top / bottom electrode lead-port mask. A circular quartz chromium plate with a diameter of 10 mm and a thickness of 4 μm was selected as the substrate. The mask pattern includes an effective area matching the flexible substrate, and symmetrically distributed 100 μm wide lead patterns and 2 mm × 2 mm square port patterns. The distance between the leads and functional areas is ≥500 μm. The linewidth accuracy of the mask is controlled to ±5 μm, and the pattern alignment deviation is ≤3 μm. The lead aperture thickness direction is adapted to a total deposition thickness of 560 nm, meeting the deposition requirements of a 20 nm TiN adhesion layer, a 500 nm Cu conductive layer, a 10 nm TiN barrier layer, and a 30 nm Au lead-out layer. The bottom electrode leads are disposed on the surface of a flexible PI substrate treated with oxygen plasma activation and are formed simultaneously with the port structure through a single deposition process. Subsequently, the prepared integrated top / bottom electrode lead-port mask was precisely fitted and fixed onto the surface of the flexible substrate. The flexible substrate sample was then transferred to the chamber of the ultra-high vacuum magnetron sputtering system, and the vacuum was evacuated to a base vacuum level ≥5×10⁻⁶. -8 Pa, maintain room temperature conditions, and add a sample stage at a constant speed of 15 rpm in all thin film deposition steps to ensure that the thickness uniformity deviation of each layer is ≤±3%.

[0125] In this embodiment, a 20 nm TiN adhesion layer was deposited using a high-purity Ti target with a purity ≥99.99% via reactive sputtering in an Ar / N2 mixed atmosphere (volume ratio 80:20). The argon gas (99.99% purity) flow rate was 20 sccm, the nitrogen gas (99.99% purity) flow rate was 5 sccm, the sputtering power was set to 100 W, the deposition pressure was maintained at 2.25 mTorr, and the deposition rate was controlled at 0.15 nm / s using a quartz crystal microbalance (QCM). The deposition thickness was 20 nm. This layer serves as a bonding layer to the subsequent Cu layer, providing adhesion and Cu atom diffusion barrier functions. Room temperature deposition avoids thermal damage to the substrate.

[0126] Further, a 500 nm Cu conductive layer was deposited using a high-purity Cu target with a purity ≥99.99%. Parameters: argon flow rate 30 sccm, operating pressure 1.875 mTorr, DC sputtering power 120 W, deposition rate 0.5 nm / s. Control: 500 nm thickness was deposited under QCM control. Thickness uniformity was monitored in real time to ensure a deviation ≤ ±3%. The Cu layer serves as the core conductive carrier for the lead, and its low resistivity ensures attenuated transmission of bioelectrical signals.

[0127] Next, a 10 nm TiN barrier layer is deposited using a high-purity Ti target with a purity ≥99.99%. The deposition is performed via reactive sputtering in an Ar / N2 mixed atmosphere (volume ratio 80:20) (Ar flow rate 20 sccm, N2 flow rate 5 sccm, both with a purity of 99.99%), with a sputtering power of 100 W, a deposition pressure of 2.25 mTorr, and a deposition rate of 0.12 nm / s. This layer isolates the atomic interdiffusion between the Cu layer and the surface Au layer, preventing the formation of high-resistivity Cu-Au intermetallic compounds and ensuring the long-term conductivity stability of the lead wire.

[0128] Next, a 30 nm thick Au lead-out layer was deposited using DC magnetron sputtering. The target material was a high-purity Au target with a purity ≥99.99%, the sputtering power was 60 W, the working gas was pure Ar (purity 99.99%), the gas pressure was maintained at 1.5 mTorr, and the deposition rate was controlled at 0.08 nm / s. After deposition, the mask was removed to obtain a bottom electrode lead structure integrated with the port.

[0129] Furthermore, after the bottom electrode lead fabrication is completed, the edge morphology and size of the pattern are observed using SEM to confirm that there are no burrs and the line width deviation is ≤ ±5 μm; the interlayer adhesion is verified by tape peeling test to ensure that the metal layer does not fall off; the conduction resistance is tested using a probe station to check for pinholes or open circuit defects; and the alignment accuracy of the mask is checked with a metallographic microscope to ensure that the alignment deviation between the lead and the subsequent functional layer is ≤ 3 μm.

[0130] (3) Insulation layer preparation process

[0131] The mask is prepared using a high-precision photolithography process to create a special mask for the insulating layer. The substrate is a circular quartz chrome plate with a diameter of 10 mm and a thickness of 4 μm. The mask pattern is designed to completely block only the 2 mm × 2 mm port area, while the rest of the area is completely cut out. The line width accuracy is controlled to ±5 μm, and the alignment deviation is ≤3 μm, which can accurately match the structural dimensions of the flexible substrate and the bottom electrode lead.

[0132] Furthermore, the prepared insulating layer mask is precisely fitted and fixed onto the surface of the bottom electrode lead, transferred to the chamber of the ultra-high vacuum magnetron sputtering system, and evacuated to a base vacuum level ≥5×10⁻⁶. -8At room temperature, a high-purity Si target with a purity ≥99.99% was used. Radio frequency reactive sputtering was performed in an Ar / N2 mixed atmosphere (volume ratio 70:30, with Ar flow rate 21 sccm and N2 flow rate 9 sccm, both gases 99.99% pure). The sputtering power was set to 80 W, the deposition pressure to 1.8 mTorr, and the deposition rate to 0.08 nm / s. A 300 nm Si3N4 insulating layer was precisely deposited in the exposed area. After deposition, the chamber was kept under vacuum for 5 minutes to allow initial stress release. The pressure was then slowly released to atmospheric pressure, and the mask was removed, resulting in an insulating layer structure with exposed port areas and full coverage of other functional areas. This insulating layer effectively isolates the bottom electrode leads from subsequent functional layers, preventing unintended electrical conduction. Its dense structure also blocks external moisture and impurities, ensuring device stability.

[0133] (4) Conductive island fabrication process

[0134] The mask was fabricated using a high-precision photolithography process, creating a shared concentric circular mask for the conductive island and MTJ spin functional layer. A 1 cm × 1 cm circular quartz chrome plate with a thickness of 5 μm was used as the substrate. The mask pattern consisted of 20 circular openings arranged in three concentric circles within a 6 mm diameter area at the center of the substrate: 4 in the inner layer, 6 in the middle layer, and 10 in the outer layer. Each opening had a diameter of 200 μm, matching the MTJ structure diameter of 200 μm or less than or equal to the diameter of the MTJ free layer. The opening diameter precisely matched the MTJ functional layer diameter with a deviation of ≤ ±1 μm. The linewidth accuracy was controlled to ±5 μm, with an alignment deviation of ≤3 μm, ensuring precise matching of the flexible substrate and the subsequent deposition and positioning requirements of the MTJ functional layer.

[0135] Furthermore, the prepared shared concentric circular mask is precisely fitted and fixed onto the substrate surface where the insulating layer has been deposited, and then transferred to the chamber of the ultra-high vacuum magnetron sputtering system, where a vacuum is evacuated to a base vacuum level ≥5×10⁻⁶. -8 Pa, maintaining room temperature; using a high-purity Au target with a purity ≥99.99%, high-purity Ar gas with a purity of 99.99% was introduced, with a gas flow rate controlled at 30 sccm, the deposition gas pressure adjusted to 2.0 mTorr, the sputtering power set to 80 W, and the deposition rate controlled at 0.1 nm / s, precisely depositing a 2 μm Au conductive island layer in the exposed opening area; after deposition, maintaining the chamber vacuum for 10 minutes, then depressurizing and removing the mask, thus obtaining a concentric circle distributed Au conductive island structure corresponding one-to-one with the mask opening (e.g., ...). Figure 2 (As shown).

[0136] (5) Preparation of microneedle biocontact layer

[0137] The mask was fabricated using a high-precision photolithography process to create a dedicated mask for microneedles. A 1 cm × 1 cm circular quartz chromium plate with a thickness of 65 μm was used as the substrate. The 65 μm thickness of the microneedle mask was designed to meet the photolithographic precision requirements of the pointed cone-shaped openings. The mask edge positioning marks are fully compatible with the concentric circle mask shared by the conductive island-MTJ spin functional layer. A total of 20 pointed cone-shaped openings were incorporated, each corresponding to an opening on the conductive island. The bottom diameter of each cone-shaped opening is 10 μm, and the tip diameter is ≤5 μm. The bottom diameter of the opening is ≤200 μm of the conductive island diameter. The linewidth accuracy is controlled to ±5 μm, the alignment deviation is ≤3 μm, and the perpendicularity of the opening sidewall is ≥89°, adapting to the requirements of multilayer deposition of microneedles. By aligning the mask positioning marks with the substrate reference point, the centers of the microneedles and conductive islands are precisely coincident, with an alignment deviation ≤±3 μm, ensuring efficient transmission of biocurrent to the MTJ layer. The flexible polyimide substrate with the mask aligned was then transferred to the ultra-high vacuum magnetron sputtering system chamber, and evacuated to a base vacuum level ≥5×10⁻⁶. -8 Under ambient temperature conditions, Ti / TiN / Au seed layers were sputtered sequentially. The Au seed layer was then subjected to plasma activation treatment (Ar gas, 100 W, 5 min). Following this, a Ni core layer was pulse-plated and a Pd biocompatible layer was DC-plated. Specific parameters are as follows:

[0138] 3 nm Ti metal bonding layer: A high-purity Ti target with a purity of ≥99.99% was used, and a high-purity Ar gas with a purity of 99.99% was introduced at a flow rate of 15 sccm, a deposition gas pressure of 1.5 mTorr, a sputtering power of 60 W, a deposition rate of 0.05 nm / s, and a vacuum was maintained for 5 min after sputtering.

[0139] 25 nm TiN barrier layer: A high-purity Ti target with a purity ≥99.99% was used, and an Ar / N2 mixed atmosphere was introduced (volume ratio 75:25, Ar flow rate 20 sccm, N2 flow rate 6.7 sccm, and gas purity 99.99%), deposition pressure 2.0 mTorr, sputtering power 100 W, and deposition rate 0.1 nm / s;

[0140] 100 nm Au conductive seed layer: A high-purity Au target with a purity ≥99.999% was used, and a high-purity Ar gas with a purity of 99.99% was introduced at a flow rate of 25 sccm, a deposition gas pressure of 2.5 mTorr, a sputtering power of 80 W, and a deposition rate of 0.12 nm / s;

[0141] 1700 nm Ni core layer: A pulse electroplating process was carried out based on the Au seed layer, using a nickel sulfamate electrolyte system (nickel sulfamate 300 g / L + boric acid 40 g / L). The electrolyte pH was controlled at 4.0-4.2 and the temperature at 50±2°C. The pulse parameters were set as follows: forward current density 0.5 A / dm², reverse current density 0.1 A / dm², forward time 10 ms, reverse time 1 ms. The forward / reverse current density ratio (5:1) and duty cycle were selected to maximize hydrogen release and reduce scorching while ensuring the deposition rate, thereby obtaining a low-stress, dense nickel microneedle core. To avoid stress in the coating, the deposition was performed in 5 stages, each with a thickness of approximately 340 nm. After each deposition, the core was rinsed in flowing deionized water (resistivity ≥18 MΩ·cm) at 40-50°C for 1 minute and then dried with nitrogen. After electroplating, electrochemical polishing is used to remove surface burrs, improve the smoothness of the microneedle tip, and avoid skin irritation. The final Ni conductive layer surface should be uniformly silvery-white, free from defects such as cracks and charring.

[0142] 100 nm Pd biocompatible top layer: Based on the formed Ni conductive core layer, a DC electroplating process was carried out. The palladium sulfamate electrolyte system (palladium sulfamate 20 g / L + ammonium chloride 50 g / L + sodium citrate 30 g / L) was selected. The current density was controlled at 0.5 A / dm², the deposition rate at 0.02 nm / s, the electrolyte temperature at 40-50°C, the pH value at 7.5-8.5, and the deposition time at about 1-2 minutes to ensure that the Pd layer surface is uniformly silvery white and that the Ni conductive core layer has good electrical conductivity.

[0143] Furthermore, after electroplating, the sample surface was ultrasonically cleaned with deionized water to remove residual electroplating solution for 3 minutes, with the ultrasonic power controlled at 50 W to avoid breakage of the microneedle tips due to high power. After cleaning, the microneedle morphology was inspected using an optical microscope to ensure that the tips were free from deformation and defects. After drying with nitrogen, the microneedle biocontact layer was prepared, thus obtaining the pointed conical microneedle structure corresponding one-to-one with the conductive island (e.g., Figure 2 (As shown).

[0144] In some embodiments, the Ti adhesive layer uses low power and low deposition rate to enhance interfacial adhesion with the substrate and subsequent TiN layers, preventing interlayer delamination after microneedle formation. The TiN barrier layer uses an Ar / N2 mixed atmosphere and medium power parameters to ensure film density and effectively block the interdiffusion of Au and Ni atoms. The Au seed layer is matched with medium to high sputtering power to improve film conductivity and provide a uniform conductive substrate for subsequent Ni and Pd electroplating. The Ni core layer is deposited in batches using pulse electroplating, combined with reverse current and intermediate water washing processes, which can significantly reduce the internal stress of the coating and prevent cracks or deformation of the cone structure. The Pd top layer uses mild DC electroplating parameters to ensure uniform coverage of the microneedle tip, improving biocompatibility while avoiding damage to the structural integrity of the Ni core layer.

[0145] (6) Spin-electron functional layer (magnetic tunnel junction MTJ) fabrication process

[0146] Before reusing the shared concentric circle mask for the conductive island-MTJ spin functional layer, a temporary protective mask needs to be prepared on the microneedle surface: a flexible photoresist film is applied using a flexible photoresist lamination process, and after pre-baking (90℃, 30 min), only the protective layer in the microneedle area is retained. Then, the sample with the protective mask is precisely aligned and bonded to the shared concentric circle mask, ensuring that the mask openings only cover the conductive island area, the microneedles are completely blocked by the photoresist, and the alignment deviation is ≤3 μm.

[0147] Furthermore, after mask protection, all films were continuously deposited at room temperature (25-30°C) in an environment that did not disrupt the chamber vacuum, maintaining a chamber background vacuum of ≥5×10⁻⁶. -8 Pa ensures that the interfaces of each film layer are pure, free from oxidation and impurities. The spin-electron functional layer on the surface of the conductive island is a magnetic tunnel junction (MTJ) structure, consisting of a 1.5 nm Ti buffer layer and a 10 nm Ir layer from bottom to top. 22 Mn 78 Pinning layer, 2.0 nm Fe 45 Co 35 B 20 Reference layer, 0.8 nm MgO barrier layer, 1.6 nm Fe 40 Co 40 B 20 The deposition process parameters and functions of each layer are as follows: Free layer, 4 nm TiN protective layer:

[0148] 1.5 nm Ti buffer layer: A high-purity Ti target with a purity ≥99.99% was selected, and DC magnetron sputtering was employed. The sputtering power was set to 80 W, a pure Ar atmosphere, a chamber pressure of 1 mTorr, and a deposition rate of 0.05 nm / s. During the deposition process, the film thickness was precisely controlled to 1.5 nm in real time using a quartz crystal oscillator. The low-speed deposition mode ensured that the Ti layer uniformly covered the Au conductive islands and the PI substrate surface, playing a dual role in adhesion and atomic diffusion barrier. The ultra-thin thickness design ensured interfacial adhesion while minimizing the impact of film stress on the overall device flexibility.

[0149] 10 nm Ir 22 Mn 78 Pinning layer: Ir with an atomic ratio of Ir:Mn = 22:78 is selected. 22 Mn 78 The alloy target was sputtered using DC magnetron sputtering at a power of 50-60 W, a pure Ar atmosphere, a chamber pressure of 1.5 mTorr, and a deposition rate of 0.08 nm / s. The co-cavity sputtering process does not disrupt the vacuum environment and avoids interface contamination. The underlying Ti buffer layer effectively prevents Ir and Mn atoms from diffusing to the Au conductive islands, ensuring the stability of the Ir target. 22 Mn 78 Stability of the exchange bias field in the pinned layer. Appropriate sputtering power ensures stable compositional transfer and ordering of the antiferromagnetic IrMn alloy, providing a structural basis for generating a strong exchange bias field, while also controlling the deposition heat input and protecting the flexible substrate.

[0150] 2.0 nm Fe 45 Co 35 B 20 Reference layer: Fe with an atomic ratio of Fe:Co:B = 45:35:20 is selected. 45 Co 35 B 20 An alloy target was used, and DC magnetron sputtering was employed with a sputtering power of 60 W, a pure Ar atmosphere, a chamber pressure of 1 mTorr, and a deposition rate of ≤0.05 nm / s. This low-rate deposition mode yielded amorphous Fe. 45 Co 35 B 20 The thin film, with a thickness precisely controlled to 2.0 nm, allows for optimization of crystallization in subsequent annealing processes, ensuring its compatibility with Ir. 22 Mn 78 The pinned layers form a stable magnetic coupling relationship. The composition range is 40%–50% Fe atoms, 30%–40% Co atoms, and 15%–25% B atoms. This range is optimal for the formation of FeCoB amorphous alloys. Furthermore, the reference layer has a higher Fe content than the free layer, which enhances the interaction with Ir. 22 Mn 78The switching bias effect of the pinning layers stabilizes the magnetic moment of the reference layer.

[0151] 0.8 nm MgO barrier layer: The chamber background vacuum level is ≥5×10⁻⁶ during continuous deposition of the MTJ functional layer. -8 Pa, where the chamber vacuum level needs to be further increased to ≥5×10 before the MgO barrier layer deposition. -10 To avoid residual gas contamination of the barrier layer interface, a high-purity MgO target with a purity ≥99.99% was selected. Radio frequency magnetron sputtering was employed, with a sputtering power of 40 W, an Ar / O2 mixed atmosphere with a volume ratio of 98:2, a chamber pressure of 1 mTorr, and a deposition rate of 0.02 nm / s. This step is the core critical process for the fabrication of the magnetic tunnel junction. The 2% O2 content optimizes the stoichiometry of the MgO film, reducing oxygen vacancy defects. The extremely low deposition rate and precise thickness control of 0.8 nm are key technical means to obtain tunnel junctions with high barrier height and low pinhole density. Furthermore, strict control of the O2 ratio prevents oxidation of the underlying FeCoB reference layer. The 0.8 nm thickness design balances the tunneling probability of bioelectric signals with the insulation performance of the barrier layer, providing reliable device structure support for efficient signal conversion of Oersted field-driven magnetic moment deflection.

[0152] 1.6 nm Fe 40 Co 40 B 20 Free layer: Fe with an atomic ratio of Fe:Co:B = 40:40:20 is selected. 40 Co 40 B 20 An alloy target was used, employing DC magnetron sputtering with a sputtering power of 60 W, a pure Ar atmosphere, a chamber pressure of 1 mTorr, and a deposition rate ≤0.05 nm / s. The free layer thickness was precisely controlled to 1.6 nm. This thickness design optimizes the magnetic moment response speed, adapting to the Oersted field driving requirements excited by weak bioelectrical signals on the body surface. The composition range of the free layer was adjustable: 35%–45% Fe atoms, 35%–45% Co atoms, and 15%–25% B atoms. The lower Fe content reduces magnetic anisotropy, significantly improving the sensitivity of the free layer's magnetic moment response to the Oersted field. This ensures that the current induced by capacitive coupling from the body surface potential difference can efficiently drive the magnetic moment deflection, guaranteeing the conversion efficiency of the Oersted field-magnetic moment deflection-resistance signal. The free layer used the same sputtering parameters as the reference layer, ensuring the crystal orientation and performance matching of the two FeCoB thin films, laying the foundation for stable tunneling magnetoresistance effects in the MTJ device.

[0153] 4 nm TiN protective layer: A high-purity TiN target with a purity of ≥99.99% was selected, and radio frequency magnetron sputtering was adopted. The sputtering power was set to 100 W, Ar / N2 mixed atmosphere with a volume ratio of 90:10, chamber pressure of 1.2 mTorr, and deposition rate of 0.08 nm / s. The high sputtering power ensured the formation of a dense TiN encapsulation protective layer. During the deposition process, the sample temperature was monitored in real time by an infrared thermometer to ensure that the surface temperature of the PI substrate was always below 100℃. This effectively protected the magnetic tunnel junction while adhering to the design requirements of low-temperature process throughout the entire process.

[0154] Further, after completing the continuous deposition of the MTJ functional layer, the temporary protective mask of photoresist and the residual MTJ film on the surface were removed. A gentle lift-off process was used, followed by ultrasonic cleaning with acetone (50 W power, 3 min time), then rinsing with anhydrous ethanol and drying with high-purity nitrogen to avoid mechanical scratching damage to the device. Subsequently, FeCoB crystallization low-temperature microwave annealing was performed, specifically as follows: the sample with the complete structure deposition of bottom electrode leads (TiN / Cu / TiN / Au), 60 nm Si3N4 insulating layer, 2 μm Au conductive island, microneedle bio-contact layer and spintronic functional layer (MTJ) was laid flat on a quartz substrate and transferred to a 2.45 GHz microwave annealing chamber, and the chamber vacuum was evacuated to ≥5 × 10⁻⁶. -8 To maintain a consistent high-vacuum process, high-purity Ar gas (5 sccm flow rate) was introduced as a protective atmosphere to prevent film oxidation. No additional heating device was installed in the annealing chamber. The microwave frequency was set to 2.45 GHz continuous wave mode, the output power to 100 W, and the annealing time to 15 min. The distance between the microwave antenna and the sample surface was adjusted to 5 cm to ensure uniform microwave energy irradiation of the lead and functional layer regions of the sample. A miniature infrared thermometer (resolution ±0.1℃) was attached to the sample surface in the 2.45 GHz microwave annealing chamber to monitor the PI substrate surface temperature in real time. The PI substrate's heat distortion temperature was >200℃. After the annealing process, the chamber was kept under vacuum until the sample naturally cooled to room temperature (cooling rate ≤5℃ / min) before removal to prevent stress cracking at the film interface caused by sudden temperature changes.

[0155] Preferably, this microwave annealing process can selectively heat the FeCoB layer, raising its temperature to approximately 140°C to achieve optimized crystallization. This temperature is the optimal range for FeCoB crystallization, maximizing the tunneling magnetoresistance (TMR) performance of the MTJ. Simultaneously, it eliminates the internal stress of the lead metal layer, promotes slight atomic diffusion at the interface between the Au lead and the Ti buffer layer, strengthens ohmic contact performance, and provides stable interface support for the efficient conduction of biocurrent-Oersted field.

[0156] In a preferred embodiment, the 5 nm Ti buffer layer is a common process parameter for traditional rigid substrate MTJ devices. This thickness, in rigid substrate systems such as silicon, can fully ensure the interfacial adhesion between the MTJ functional layer and the substrate, while effectively blocking the long-range diffusion of metal atoms, thus meeting the structural stability requirements of rigid devices. Based on the heat resistance and bending performance requirements of the flexible PI substrate of this invention, the thickness of the Ti buffer layer is reduced from 5 nm to 1.5 nm. The core reason for thinning lies in balancing the three requirements of interface adhesion, flexibility, and thermal stress control: On the one hand, the ultra-thin thickness of 1.5 nm can minimize the constraint of the rigid metal layer on the flexibility of the PI substrate while ensuring the interfacial adhesion of the Ti layer to the Au conductive island and MTJ pinning layer, and the basic function of blocking Au / IrMn atomic interdiffusion. This avoids cracks or interlayer delamination caused by stress concentration in the thick Ti layer during device bending, and ensures that the functional stability of the Oersted field-driven magnetic moment deflection is not affected by bending. On the other hand, combined with the low-temperature sputtering process of ≤30℃ throughout, the ultra-thin Ti layer can reduce the heat accumulation during the deposition process, prevent local temperature rise of the substrate caused by thick film deposition (exceeding the heat resistance threshold of 100℃ of the PI substrate), and avoid thermal deformation of the substrate.

[0157] Furthermore, after annealing, the sample is placed on an ultra-high vacuum probe stage (vacuum degree ≥ 5 × 10⁻⁶). -7 On a plate (Pa), the electrical performance of a single MTJ unit was tested using the four-probe method. A DC bias voltage of 0.1~1.0 V was applied, and the tunneling magnetoresistance ratio (TMR) was calculated. The criterion was TMR ≥ 180%. The surface roughness of the MgO barrier layer was tested using atomic force microscopy (AFM). The criterion was interface roughness ≤ 0.5 nm. The MTJ barrier height ΔE was tested. The criterion was ΔE = 2.0~2.6 eV. This range is the optimal range for the FeCoB / MgO system to achieve high spin polarization tunneling, which directly determines the TMR amplification factor (up to 200% or more), providing core performance assurance for the efficient conversion of Oersted field-magnetic moment deflection-resistance signal.

[0158] Furthermore, only after all the above performance indicators meet the standards can the subsequent interface planarization and top electrode deposition processes proceed to ensure that the core performance of the MTJ meets the design requirements. Moreover, after 1000 bending cycles, the ΔE fluctuation is ≤±0.08 eV, demonstrating excellent mechanical bending stability, which can meet the practical application requirements of wearable bioelectronic dry electrodes, ultimately achieving synergistic optimization of the device's mechanical bending stability and Oersted field-controlled electrical performance.

[0159] (7) Interface smoothing insulation layer preparation process

[0160] The mask was fabricated using a high-precision photolithography process to prepare the interface planarization insulating layer mask. The substrate was a 1cm×1cm circular quartz chrome plate with a thickness of 5 μm. The positioning marks on the edge of the mask were fully compatible with the bottom electrode lead mask, the concentric circle mask shared by the conductive island-MTJ, and the microneedle-specific mask. A total of 20 circular openings were set, which precisely corresponded to the centers of the 20 openings of the conductive island-MTJ shared mask. The diameter of a single opening was ≤ the diameter of the MTJ free layer, and the dimensional deviation was ≤ ±1 μm. This ensured that only the MTJ free layer area was exposed, while the conductive island, microneedles, and non-functional areas of the flexible substrate were completely shielded, which met the selective deposition requirements of the interface planarization insulating layer Si3N4.

[0161] Furthermore, the prepared interface-smoothed insulating layer mask is precisely fitted and fixed onto the annealed MTJ functional layer and substrate surface, with an alignment deviation ≤3 μm. It is then transferred to the ultra-high vacuum magnetron sputtering system chamber and evacuated to a base vacuum level ≥5×10⁻⁶. -8 The chamber temperature was maintained at 25-30℃. A high-purity Si target with a purity ≥99.99% was used, and an Ar / N2 mixed atmosphere with a volume ratio of 70:30 was introduced (Ar flow rate 21 sccm, N2 flow rate 9 sccm, both gas purity 99.99%). The RF reactive sputtering power was set to 180 W, the deposition pressure to 2.5 mTorr, and the deposition rate to 0.3 nm / s. A Si3N4 interface planarization insulating layer with a thickness of 2 μm was precisely deposited on the uneven steps outside the exposed area of ​​the mask. After deposition, the chamber was kept under vacuum for 15 min. After the film stress was released, the pressure was slowly released to atmospheric pressure, and the mask was removed. This achieved a planarized surface structure with the MTJ free layer completely exposed and the remaining area filled and covered by the insulating layer. During the deposition process, the substrate temperature was monitored with an infrared thermometer to ≤50℃ to prevent thermal deformation of the PI substrate caused by prolonged high-power sputtering.

[0162] The thickness differences formed by multi-step deposition (such as 2 μm Au conductive islands and 1700 nm Ni microneedle cores) cause significant uneven steps on the device surface, which can easily lead to uneven top electrode film thickness, open circuits, or short circuits, affecting the consistency of electrical performance. Preferably, this process fills the steps in the non-functional area by selectively depositing a Si3N4 insulating layer, making the surface roughness ≤0.5 nm, providing a uniform substrate for the top electrode, while accurately exposing the MTJ free layer to ensure effective electrical contact between it and the top electrode and avoid tunneling magnetoresistance effect failure. Furthermore, the use of the same Si3N4 material and consistent sputtering parameters as the bottom insulating layer can improve interlayer compatibility, avoid stress cracking, ensure insulation performance and device flexibility, and meet the application requirements of wearable bioelectric electrodes.

[0163] (8) Fabrication process of bus top electrode

[0164] This process requires no mask assistance; deposition is directly performed on the surface of the interface-planarized insulating Si3N4 layer and the exposed MTJ free layer. A TiN / Au composite bus layer is sequentially prepared using ultra-high vacuum magnetron sputtering, enabling the electrical parallel connection of 20 MTJ sensing pillars and ensuring efficient signal aggregation and transmission. The composite layer, from bottom to top, consists of a 20 nm TiN adhesion layer and a 30 nm Au conductive core layer. Specific process parameters are as follows:

[0165] 20 nm TiN adhesion layer: A high-purity Ti target with a purity ≥99.99% was selected, transferred to the chamber of an ultra-high vacuum magnetron sputtering system, and evacuated to a base vacuum level ≥5×10⁻⁶. -8 Pa, maintaining the chamber temperature at room temperature of 25~30℃; introducing an Ar / N2 mixed atmosphere with a volume ratio of 80:20 (Ar flow rate 20 sccm, N2 flow rate 5 sccm, gas purity 99.99%), setting the sputtering power to 100 W, deposition pressure to 2.25 mTorr, and deposition rate to 0.12 nm / s; monitoring the film thickness in real time using a quartz crystal oscillator, accurately depositing to 20 nm. This layer can enhance the interfacial adhesion between the composite bus top electrode and the substrate, while preventing Au atoms from diffusing into the MTJ functional layer.

[0166] 30 nm Au conductive core layer: To maintain the vacuum environment of the chamber, replace the target with a high-purity Au target with a purity ≥99.999%; introduce high-purity Ar gas with a purity of 99.99% at a flow rate of 30 sccm, adjust the deposition gas pressure to 2.0 mTorr, set the sputtering power to 80W, and the deposition rate to 0.1 nm / s; accurately deposit a 30 nm thick Au layer. This layer serves as the core conductive layer and enables the low-resistance parallel connection of 20 MTJ free layers, ensuring the efficient collection and transmission of bioelectrical signals.

[0167] Furthermore, after deposition, the chamber is kept under vacuum for 10 minutes to allow the stress in the membrane layer to be fully released. Then, the pressure is slowly released to atmospheric pressure to complete the preparation of the busbar layer.

[0168] (9) Top electrode lead fabrication process

[0169] The fabrication process for the top electrode lead reuses an integrated mask template for both the top and bottom electrode leads and the port. Precise alignment with the bottom electrode lead is achieved through pre-set alignment marks on the substrate, with an alignment deviation of ≤3 μm. This ensures precise matching of the wiring path, spacing, and port position of the top and bottom electrode leads. Subsequently, the lead is transferred to the ultra-high vacuum magnetron sputtering system chamber and evacuated to a base vacuum level ≥5×10⁻⁶. -8 Pa, maintain the chamber temperature at room temperature of 25~30℃.

[0170] Furthermore, the top electrode lead adopts the same multilayer composite structure and process parameters as the bottom electrode lead, consisting of a 20 nm TiN adhesion layer, a 500 nm Cu conductive layer, a 10 nm TiN barrier layer, and a 30 nm Au lead-out layer from bottom to top, ensuring that the signal transmission performance of the two is matched. The specific process is as follows:

[0171] 20 nm TiN Adhesion Layer: A high-purity Ti target with a purity ≥99.99% was selected, and an Ar / N2 mixed atmosphere with a volume ratio of 80:20 was introduced (Ar flow rate 20 sccm, N2 flow rate 5 sccm, both gas purity 99.99%); the sputtering power was set to 100 W, the deposition pressure to 2.25 mTorr, and the deposition rate to 0.12 nm / s; the film thickness was monitored in real time using a quartz crystal oscillator, and the layer was precisely deposited to 20 nm. This layer enhances the interfacial adhesion between the top electrode lead and the bus layer, while simultaneously preventing Cu atoms from diffusing into the MTJ functional layer.

[0172] 500 nm Cu conductive layer: Maintain the vacuum environment of the chamber without damage, replace with a high-purity Cu target with a purity ≥99.999%; introduce high-purity Ar gas with a purity of 99.99% at a flow rate of 35 sccm, adjust the deposition gas pressure to 2.5 mTorr, set the sputtering power to 150W, and the deposition rate to 0.5 nm / s; monitor the substrate temperature ≤50℃ during the deposition process using an infrared thermometer to avoid deformation of the PI substrate due to high temperature, and ensure the low-resistivity conductivity of the 500 nm thick Cu layer.

[0173] 10 nm TiN barrier layer: Without disrupting the vacuum environment, a high-purity Ti target is used again, and the gas ratio, sputtering power, deposition gas pressure parameters and rate parameters of the TiN adhesive layer are used. A 10 nm TiN layer is precisely deposited, which effectively isolates the atomic interdiffusion between the Cu conductive layer and the surface Au lead-out layer, prevents the formation of high-resistivity Cu-Au intermetallic compounds, and ensures the long-term conductivity stability of the lead wire.

[0174] 30 nm Au lead-out layer: Maintaining the vacuum environment, replace the target with a high-purity Au target with a purity ≥99.999%; introduce high-purity Ar gas with a purity of 99.99% at a flow rate of 30 sccm, a deposition pressure of 2.0 mTorr, a sputtering power of 80 W, and a deposition rate of 0.1 nm / s; deposit a 30 nm thick Au layer to improve the oxidation resistance of the lead surface and the compatibility of the pads, and realize the summary and export of signals from 20 MTJ units.

[0175] Furthermore, after deposition, the chamber is kept under vacuum for 15 minutes to allow the stress in each film layer to be fully released. Then, the pressure is slowly released to atmospheric pressure, and the mask is removed. The core reasons for keeping the process parameters consistent with the bottom electrode leads are: first, to ensure that the resistance and conductivity stability of the top and bottom electrode leads are consistent, avoiding impedance mismatch problems during signal transmission and ensuring that the biopotential difference on the body surface is efficiently converted into the Oersted field excitation current driving the MTJ; second, to simplify the process flow, reduce mask replacement and parameter adjustment steps, and improve mass production efficiency; and third, to ensure consistent compatibility between the leads and the flexible PI substrate, avoiding bending performance differentiation due to parameter differences, and ensuring the long-term stability of the Oersted field-magnetic moment deflection coupling mechanism of the device in wearable scenarios.

[0176] In a preferred embodiment, the previous MTJ crystallization low-temperature microwave annealing only completed the crystallization of the FeCoB layer and the optimization of the MTJ interface. The subsequent deposited planarized TiN insulating layer, TiN / Au bus top electrode, top / bottom electrode lead multilayer structure has problems such as interface stress accumulation, insufficient metal layer atomic diffusion, and high ohmic contact resistance. Therefore, a stress relief annealing process is required to eliminate the above defects.

[0177] Furthermore, the device with the top electrode lead deposited was laid flat on a quartz substrate, transferred to a 2.45 GHz microwave annealing chamber, and evacuated to a chamber vacuum level ≥ 5 × 10⁻⁶. -8 To maintain consistency in the high vacuum process, high-purity Ar gas (flow rate 5 sccm) was introduced as a protective atmosphere to prevent oxygen residue from oxidizing the MTJ layer. The microwave power was set to 50 W (less than 50% of the MTJ core annealing power), the annealing time to 8 min, and the antenna-sample distance was maintained at 5 cm. During annealing, the sample surface temperature was monitored by infrared thermography to be ≤80℃. The temperature rise range of the TiN / Au bus layer and the top electrode lead metal layer was 60-70℃. This temperature can effectively eliminate interlayer interface stress, strengthen ohmic contact, and avoid high temperature damage to the structure of the crystallized MTJ functional layer. The MTJ resistance change was monitored during annealing. Annealing was stopped when the TMR value reached its peak. After annealing, the sample was naturally cooled to room temperature under vacuum (cooling rate ≤5℃ / min) before being removed. The annealing process is placed after the top electrode lead and before the encapsulation layer. This can prevent the encapsulation layer material (especially parylene) from failing due to its low temperature tolerance (it is easy to decompose at >100℃). At the same time, it ensures the optimization effect of annealing on the newly added metal layer. After the supplementary annealing is completed, the on-resistance of the top electrode lead is retested using a probe station. The resistance change rate is required to be ≤±2%. At the same time, the TMR value of the MTJ unit is randomly checked to ensure that the TMR is ≥180%. Only after meeting the standard can it enter the encapsulation process.

[0178] (10) Packaging process

[0179] A pad-microneedle tip protection mask is used, with a 1 cm × 1 cm circular quartz chrome plate substrate, 5 μm thick, and edge positioning marks compatible with all masks in the entire process. The mask is precisely attached to the device surface, ensuring a perfect match between two types of solid masking areas and the target area: one type consists of two 2 mm × 2 mm square solid blocks, precisely covering the pad areas of the top / bottom electrodes; the other type consists of 20 circular solid blocks with a diameter ≤ 5 μm, precisely covering the microneedle tip area ≤ 5 μm. Except for the pad and microneedle tip masking areas, the remaining functional layer areas are all cutout designs. The deviation is controlled to ≤ ± 2 μm during alignment, ensuring the boundary accuracy between the packaged area and the exposed area. No subsequent etching process is required, and the target area is exposed without damage in one step through the mask design.

[0180] Further, the device with the mask attached is transferred to the atomic layer deposition (ALD) system chamber, and a low-temperature Al2O3 first passivation layer is deposited using trimethylaluminum (TMA) and deionized water as precursors: the chamber is evacuated to a base vacuum level ≥5×10⁻⁶. -6 The deposition temperature was set at 80℃, with precursor pulse durations of 0.1 s (TMA) and 0.05 s (H2O), a nitrogen purging time of 10 s, and 1000 deposition cycles. A quartz crystal microbalance (QCM) was used to monitor the film thickness in real time, precisely depositing a 20 nm dense Al2O3 passivation layer. This self-limiting growth film possesses 100% conformal coverage, forming a uniform and defect-free protective layer on the steep sidewalls of microneedles and at MTJ nanoscale steps. It effectively blocks water vapor and ion penetration, and the low-temperature deposition process completely avoids plasma bombardment, preventing damage to the MTJ MgO barrier layer.

[0181] Furthermore, the device surface mask is kept intact and directly transferred to the chemical vapor deposition (CVD) parylene system. The transfer process must be carried out in a clean, dry atmosphere to avoid mask displacement or device surface contamination, ensuring the feasibility of the "etch-free" process. Using parylene dimer as raw material, the system is divided into three independent temperature-controlled zones: a pyrolysis zone at 680°C, a vaporization zone at 150°C, and a deposition zone maintained at room temperature (25~30°C), with the chamber vacuum controlled at 100 mTorr. The parylene active monomers generated by pyrolysis are transported to the deposition zone via insulated tubing, where they spontaneously adsorb and polymerize on the device surface. The deposition rate is controlled at 0.1 nm / s, precisely depositing a 100 nm thick parylene protective layer. The deposition zone remains at room temperature throughout, completely isolating heat from the sample and preventing any thermal damage to the PI flexible substrate and MTJ functional layer. This layer possesses excellent flexibility and biocompatibility, tightly encapsulating the Al2O3 passivation layer, buffering bending stress, preventing inorganic layer cracking, and improving device comfort in contact with skin.

[0182] Further, after deposition, the device is removed and the mask is gently removed. At this point, the device surface has achieved the target state of an Al2O3 / CVD parylene composite encapsulation layer deposited by atomic layer deposition (ALD) process, with completely exposed pads, completely exposed microneedle tips, and full coverage of the remaining functional areas. The exposed area of ​​the pads and microneedle tips is inspected using an optical microscope to ensure no encapsulation material residue; the thickness uniformity of the encapsulation layer is measured using a profilometer, requiring a thickness deviation ≤ ±5 nm; and the surface roughness of the encapsulation layer is tested using an atomic force microscope (AFM) to ensure a roughness ≤ 0.5 nm.

[0183] This process employs a synergistic approach combining high-precision mask design, low-temperature atomic layer deposition (ALD) for non-destructive deposition, and room-temperature CVD flexible encapsulation. This completely eliminates the need for traditional ion beam etching, fundamentally preventing physical damage to the MTJ functional layer from etching. The ALD process deposits Al2O3 to provide dense hard protection, while CVD-deposited parylene provides excellent soft buffering. The combination of these two elements forms a composite encapsulation structure of "hard protection + soft buffering," perfectly matching the wearable application requirements of flexible body surface dry electrodes and ensuring the long-term stability of the device under repeated bending and humid heat environments.

[0184] (11) Performance verification and reliability testing methods

[0185] The comprehensive performance tests of the microneedle magnetoresistive brain-computer interface dry electrode of this invention were all conducted under the standard environment of room temperature 25±2℃ and atmospheric pressure 101.3 kPa. The specific test items and judgment criteria are as follows:

[0186] Electrodes were fixed to a bending tester and repeatedly bent around a cylinder with a radius of curvature of 5 mm at ±90° and a frequency of 1 Hz. After every 1000 bends, the lead resistance was measured using a four-probe method, and the MTJ magnetoresistive response curve was detected using a standard signal source and a reading circuit. The test method and judgment criteria refer to existing technology [Tanaka et al., Small, 2024]: This study confirms that MTJ devices prepared by magnetron sputtering on PI flexible substrates, after optimization by annealing in Ar plasma at 140℃ (20 min), exhibit excellent mechanical-electrical performance stability under the same radius of curvature and bending angle.

[0187] Existing technologies have only verified the flexible bending stability of PI-based MTJ structures, failing to address issues such as insufficient conductivity and high signal transmission loss when flexible MTJ devices are used as dry electrodes on the body surface. This invention, based on the aforementioned flexible MTJ structure, introduces a conductive island array to improve surface contact conductivity and employs a symmetrical dual-lead arrangement design to reduce signal interference. Simultaneously, a 2.45 GHz microwave annealing process is used, leveraging the selective heating characteristics of microwaves to precisely control the FeCoB / MgO interface temperature to 140°C and maintain it for the corresponding duration. This inherits the thermal compatibility advantages of low-temperature annealing on the PI substrate from previous literature while avoiding potential damage to the MTJ tunneling barrier layer from plasma bombardment. Ultimately, the invention achieves the criteria of lead resistance change rate <5% and MTJ tunneling magnetoresistance ratio (TMR) attenuation <10% after 1000 bends, ensuring that the stability of the biocurrent-Oersted field-magnetic moment deflection coupling mechanism is unaffected by bending.

[0188] Accelerated oxidation and environmental stability testing: The finished electrode was aged for 30 days in a constant temperature and humidity chamber at 37℃ and 90% relative humidity. After the experiment, the contact impedance was measured using an electrochemical workstation (10 Hz - 100 kHz), and the morphology of the encapsulation layer was observed using SEM. Judgment criteria: Contact impedance change rate ≤15%, and no bulging, cracking, or moisture intrusion marks on the encapsulation layer. This ensures that the electrode can stably excite an Oersted field to achieve signal conversion in wearable scenarios.

[0189] Bioelectric signal acquisition test: Electrodes were integrated into a headband / chest patch device to acquire EEG and ECG signals, and the data were compared with those from clinical-grade medical electrodes. Signal quality, key feature detection rate, and similarity under resting and mild motion conditions were analyzed. Judgment criteria: Clear signals were acquired without conductive gel; key features were highly consistent with medical electrodes; motion artifacts were effectively suppressed, verifying the feasibility and superiority of the Oersted field-driven MTJ signal conversion scheme in practical bioelectric acquisition scenarios.

[0190] Comparative Example 1

[0191] A performance comparison of different MTJ material systems was conducted. The spin-electron functional layer was replaced with a traditional NiFe (2.0 nm) / Al₂O₃ (0.8 nm) MTJ system, while the remaining structure and process parameters remained identical to those in Example 1 of this invention. At room temperature, a 10 μV simulated EEG signal at a frequency of 10 Hz was applied to the electrode surface using a standard skin model. The readout circuit employed a combination of a 100 kΩ current-limiting resistor and a constant current source (10 μA). The signal amplification gain of the electrode was detected using the readout circuit. The signal amplification gain of the control group electrode was only 18 times, significantly lower than the 128 times of the FeCoB / MgO system of this invention. The core reason for the performance difference between the two lies in the different synergistic mechanisms of the barrier layer and the electrode layer: The MgO barrier layer used in this invention is a single crystal with a rock salt structure, which forms a highly matched crystal orientation with the FeCoB electrode layer. It can selectively tunnel spin-up electrons through the symmetry filtering effect, and the spin polarization can reach more than 80%, thereby giving the MTJ device an intrinsic tunneling magnetoresistance (TMR) ratio of more than 200%. In contrast, in the traditional NiFe / Al2O3 system, the amorphous Al2O3 barrier layer lacks crystal symmetry and has no obvious screening effect on spin-polarized electrons. The spin polarization is usually less than 30%, and the corresponding TMR ratio is generally less than 50%, which makes it difficult to achieve efficient amplification of weak bioelectric signals, and even more difficult to support the high-sensitivity coupling control of biocurrent-Oersted field-magnetic moment deflection.

[0192] Comparative Example 2

[0193] A performance comparison was made between the planar contact layer and the microneedle contact layer. The microneedle bio-contact layer was replaced with a planar Pd film of the same thickness (100 nm), while the remaining structure and process parameters were completely consistent with Example 1 of this invention. Electrode output signals were collected under a dynamic motion scenario simulating skin stretching amplitude ±5% and friction frequency 1 Hz, and the amplitude of motion artifact noise and signal-to-noise ratio (SNR) were analyzed. The motion artifact noise of the control group electrode output signal reached ±50 μV, with a SNR of only 22 dB; the motion artifact noise of the electrode of this invention was only ±4.2 μV, with a SNR as high as 45 dB. The microneedle structure can stably penetrate the stratum corneum of the skin, reducing the relative displacement of the skin-electrode interface, and plays an irreplaceable role in suppressing motion artifacts and improving dynamic signal quality.

[0194] Comparative Example 3

[0195] A performance comparison of the dual-lead architecture without the MTJ spintronic functional layer was conducted. The microneedle bio-contact layer and dual-lead structure were retained, but the MTJ spintronic functional layer was removed and replaced with a traditional metal conductive path. The remaining structure and process parameters were completely consistent with Example 1 of this invention. Following the GB / T 39538-2020 standard, "Performance Testing Methods for Flexible Electronic Devices," a 24-hour continuous test was performed on a simulated skin model (resistivity 1 kΩ·cm) to monitor contact impedance, bias voltage fluctuation, and output signal-to-noise ratio. The test environment was room temperature 25±2℃ and atmospheric pressure 101.3 kPa. The control group electrode contact impedance fluctuation range was 120-180 kΩ, far exceeding the 38-48 kΩ of this invention; the 24-hour bias voltage fluctuation reached ±15 mV, and the signal-to-noise ratio was only 28 dB, significantly inferior to the electrode of this invention with a bias fluctuation of ≤±1 mV and a signal-to-noise ratio of ≥42 dB. The core reason for the performance gap between the two is that the control group lacks the coupling regulation mechanism of biocurrent-Oersted field-magnetic moment deflection. It cannot achieve in-situ amplification and noise suppression of weak bioelectric signals through MTJ devices, and can only rely on traditional metal pathways for passive signal transmission, resulting in large signal loss and weak anti-interference ability, which makes it difficult to meet the needs of high-precision bioelectric acquisition.

[0196] In summary, this invention employs a collaborative architecture of microneedle bio-contact layer, discrete conductive island array, and MTJ capacitively coupled dual-lead. By focusing and converging the diffuse biopotential difference on the body surface through the conductive islands, combined with the dual effect of in-situ magnetoresistive amplification of weak potential signals achieved by the MTJ structure based on the biocurrent-Oersted field-magnetic moment deflection coupling mechanism, it effectively solves the technical pain points of signal drift and insufficient gain of traditional dual-lead electrodes, achieving a performance breakthrough of low contact impedance, high bias stability, and high signal-to-noise ratio.

[0197] The electrode of this invention has three core advantages: high signal-to-noise ratio, low static power consumption, and strong power frequency suppression.

[0198] The core support for the high signal-to-noise ratio comes from the dual-layer structure design of the microneedle bio-contact layer and the (MTJ) spintronic functional layer. Microneedles can penetrate 10-20 μm thick stratum corneum, breaking through high impedance barriers to establish stable low impedance electrical coupling pathways and reducing signal interface attenuation. Their aperiodic topology can disrupt motion artifact coherence. Combined with the increased contact area and pressure from mechanical anchoring (using conservative values ​​P2=2P1, S2=10S1, the interference voltage attenuation can reach A=20lg(ΔZ1 / ΔZ2), where ΔZ1 and ΔZ2 represent the skin-electrode contact impedance fluctuation amplitudes of traditional planar electrodes and the microneedle electrode of this invention, respectively). Through mechanical anchoring, the microneedles can increase the contact pressure to P2=2P2 and expand the effective contact area to S2=10S1. Based on the inverse relationship between impedance fluctuation and contact pressure and area, the impedance fluctuation ratio ΔZ2 / ΔZ1=(P1·S1) / (P2·S2) is derived. Further converted to an interference voltage attenuation value in decibels A=20lg(ΔZ1 / ΔZ2), the calculated attenuation is approximately 26 dB, achieving 20 Interference signals exceeding dB are attenuated. Simultaneously, the FeCoB / MgO system MTJ with spin electron functional layers achieves an atomically smooth interface with a roughness ≤0.5 nm between the MgO barrier layer and the FeCoB free layer through process control. This ensures efficient coupling of biocurrent, Oersted field, and magnetic moment deflection, enabling in-situ amplification of weak microvolt-level signals via tunneling magnetoresistance.

[0199] The key to low quiescent power consumption lies in abandoning the traditional electrode-dependent external high-gain instrumentation amplifier and instead utilizing the magnetoresistive effect of the spintronic functional layer to amplify the signal. This process requires only nanowatt to microwatt level extremely low bias power supply, mainly relying on the Oersted field generated when biocurrent flows through the free layer to drive the magnetic moment deflection, and then converting the change in magnetic moment into the rate of change of resistance and output voltage through the tunneling magnetoresistive effect. The power consumption of the entire signal acquisition and amplification link can be reduced to the tens of microwatt level, which is one to two orders of magnitude lower than the traditional external amplification scheme.

[0200] The strong power frequency suppression capability stems from the synergistic effect of the magnetoresistive response characteristics of the spin electron functional layer and the interference suppression capability of the microneedle structure. The MTJ is only sensitive to the Oersted field component perpendicular to the film surface, while power frequency electromagnetic interference is mostly a magnetic field component parallel to the body surface, and the coupling efficiency between the two is extremely low. At the same time, the discrete array structure of the microneedles can weaken the common-mode coupling strength of power frequency interference at the electrode interface, achieving strong power frequency suppression under the dual effect.

[0201] The preferred embodiments of the present invention have been described in detail above. However, the present invention is not limited to the specific details in the above embodiments. Within the scope of the technical concept of the present invention, various equivalent transformations can be made to the technical solutions of the present invention, and these equivalent transformations all fall within the protection scope of the present invention.

Claims

1. A flexible bioelectric surface electrode based on a magnetic tunnel junction, characterized in that, From bottom to top, they include: Flexible substrate; Bottom electrode leads deposited on the surface of a flexible substrate; An insulating layer covering the surface of the flexible substrate and the bottom electrode lead area, with the reserved pad area exposed; A discrete array of conductive islands distributed on the surface of an insulating layer; Microneedle bio-contact layers are arranged one-to-one with the discrete conductive island array; After the microneedle biocontact layer is temporarily protected by photoresist, an MTJ spintronic functional layer array is fabricated on the surface of a discrete conductive island array in a one-to-one correspondence. An interface-smoothing insulating layer that covers the sidewalls of the spintronic functional layer and the non-functional region of the insulating layer of the MTJ and exposes only the surface of the free layer of the MTJ. The bus top electrode is electrically connected to the MTJ free layer; Top electrode lead deposited on the surface of the bus top electrode; A composite encapsulation layer covering the microneedle sidewalls, bus top electrode, top electrode lead, and MTJ spintronic functional layer sidewalls, exposing only the top / bottom electrode pads and the microneedle tip.

2. The flexible bioelectric surface electrode according to claim 1, characterized in that, The flexible substrate material is a high-modulus polyimide film with a thickness of 200~450μm and a Young's modulus of 8~18GPa. It has excellent flexibility and can withstand repeated bending at angles of 180° and above.

3. The flexible bioelectric surface electrode according to claim 1, characterized in that, The bottom electrode lead consists of a TiN adhesive layer, a Cu conductive layer, a TiN barrier layer, and an Au lead-out layer, arranged sequentially from the surface of the flexible substrate upwards. The thickness of the TiN adhesive layer is 15~25 nm, the thickness of the Cu conductive layer is 450~550 nm, the thickness of the TiN barrier layer is 8~12 nm, and the thickness of the Au lead-out layer is 25~35 nm.

4. The flexible bioelectric surface electrode according to claim 1, characterized in that, The insulating layer is a Si3N4 layer with a thickness of 200~500nm.

5. The flexible bioelectric surface electrode according to claim 1, characterized in that, The discrete conductive island array is a conductive island Au array, which is distributed in three concentric rings within a circular sensing area with a diameter of 6 mm: the first ring has a diameter of 1.5 mm and contains 4 conductive islands; the second ring has a diameter of 3.0 mm and contains 6 conductive islands. The third ring has a diameter of 4.5 mm and is divided into 10 conductive islands. Each conductive island has a diameter of 200~300μm and a thickness of 1~2μm. The center-to-center distance between adjacent conductive islands is ≥500μm to avoid signal crosstalk.

6. The flexible bioelectric surface electrode according to claim 1, characterized in that, The microneedle biocontact layer has a pointed conical protrusion structure, with each pointed conical microneedle vertically deposited at the geometric center of each conductive island in the conductive island array, and the aspect ratio of the microneedles is ≥5:

1. The microneedle biocontact layer consists of a Ti adhesion layer, a TiN barrier layer, an Au seed layer, a Ni conductive core layer, and a Pd biocompatible layer, arranged sequentially from the surface of the conductive island upwards. The thickness of the Ti adhesion layer is 2~5 nm, the thickness of the TiN barrier layer is 20~30 nm, the thickness of the Au seed layer is 80~120 nm, the thickness of the Ni conductive core layer is 1500~1900 nm, and the thickness of the Pd biocompatible layer is 80~120 nm.

7. The flexible bioelectric surface electrode according to claim 1, characterized in that, The spintronic functional layer of the MTJ consists of, from the surface of the conductive island upwards, a Ti buffer layer, an IrMn pinning layer, an FeCoB reference layer, an MgO barrier layer, an FeCoB free layer, and a TiN protective layer. The thickness of the Ti buffer layer is 1–3 nm, the thickness of the IrMn pinning layer is 8–12 nm, the thickness of the FeCoB reference layer is 1.8–2.64 nm, the thickness of the MgO barrier layer is 0.6–1.2 nm, the thickness of the FeCoB free layer is 1.6–2.2 nm, and the thickness of the TiN protective layer is 3–4 nm.

8. The flexible bioelectric surface electrode according to claim 7, characterized in that, The barrier height ΔE of the MgO tunneling barrier layer is 2.0~2.6 eV, and the interface roughness between it and the FeCoB free layer is ≤0.5 nm, ensuring a tunneling magnetoresistivity ≥180%.

9. The flexible bioelectric surface electrode according to claim 1, characterized in that, The interface smoothing insulating layer is a Si3N4 layer with a thickness of 1.5~2.5μm.

10. The flexible bioelectric surface electrode according to claim 1, characterized in that, The top electrode of the busbar is a TiN / Au composite layer structure, which includes a TiN adhesive layer and an Au conductive core layer from bottom to top. The TiN adhesive layer has a thickness of 15~25nm and is used to improve the bonding force with the MTJ free layer. The Au conductive core layer has a thickness of 25~35nm and is used to ensure that the potential of all MTJ sensing units is consistent, so as to realize signal busbar.

11. The flexible bioelectric surface electrode according to claim 1, characterized in that, The top electrode lead has the same layered structure as the bottom electrode lead and forms a symmetrical structure with the bottom electrode lead to reduce signal transmission differences. The top electrode lead consists of a TiN adhesive layer, a Cu conductive layer, a TiN barrier layer and an Au lead-out layer from bottom to top. The thickness of the TiN adhesive layer is 15~25 nm, the thickness of the Cu conductive layer is 450~550 nm, the thickness of the TiN barrier layer is 8~12 nm, and the thickness of the Au lead-out layer is 25~35 nm.

12. The flexible bioelectric surface electrode according to claim 1, characterized in that, The composite encapsulation layer consists of an Al2O3 passivation layer and a CVD parylene protective layer from the inside out. The Al2O3 passivation layer has a thickness of 15-25 nm and is used to achieve atomically dense protection. The CVD parylene protective layer has a thickness of 80-120 nm and is used to improve flexibility and biocompatibility.

13. The flexible bioelectric surface electrode according to claim 1, characterized in that, The MTJ spintronic functional layer is configured such that when biocurrents flow through the MTJ spintronic functional layer, an Oersted field perpendicular to the membrane surface is excited. The Oersted field directly drives the deflection of the magnetic moment of the FeCoB free layer. Furthermore, the MTJ spintronic functional layer converts microvolt-level bioelectric signals into resistance change signals based on the tunneling magnetoresistance effect, thereby achieving in-situ amplification of the signals.

14. The flexible bioelectric surface electrode according to claim 1, characterized in that, The array arrangement of the discrete conductive island array, the corresponding MTJ spintronic functional layer array, and the microneedle biocontact layer is not limited to a concentric circle distribution. It can also adopt any discrete array structure such as matrix, honeycomb, or regular hexagonal arrays, and the one-to-one correspondence between the three in the vertical direction remains unchanged.

15. A method for preparing a flexible bioelectric surface electrode based on a magnetic tunnel junction according to any one of claims 1 to 14, characterized in that, The thin film deposition process employs an ultra-high vacuum magnetron sputtering system, including the following steps: S1: Flexible substrate pretreatment; S2: Using a mask template with a corresponding integrated bottom electrode lead-port pattern, patterned bottom electrode leads are deposited on a flexible substrate; then an insulating layer is deposited over the entire surface. S3: A discrete array of conductive islands is deposited on the surface of the insulating layer using a shared mask template of the corresponding conductive island-MTJ spin functional layer; S4: Using a mask template with the same pattern as in step S3, a full-stack MTJ thin film is continuously deposited on the surface of the conductive island array and the insulating layer to form an MTJ sensing column array. S5: Using a mask template corresponding to the pattern of the free layer region of the MTJ sensor column, insulating material is deposited to cover the non-free layer region of the MTJ sensor column array, thereby achieving interface smoothing. S6: Deposit the top bus electrode on the planarized surface; S7: A patterned top electrode lead is deposited on the bus top electrode using a mask template with a corresponding top electrode lead-port integrated pattern. S8: Using a mask template corresponding to the pad-microneedle tip protection pattern, a composite encapsulation layer consisting of an Al2O3 passivation layer and a CVD parylene protective layer is deposited, exposing the pads and microneedle tips.

16. The preparation method according to claim 15, characterized in that, In step S4, before depositing the MTJ film, a temporary protective mask is formed using a peelable photoresist to cover the microneedle area. MTJ thin film deposition at room temperature and chamber background vacuum ≥5×10 -8 The process is carried out continuously under conditions of Pa, wherein the vacuum level of the chamber needs to be increased to ≥5×10 before the deposition of the MgO barrier layer. -10 Pa, the argon purity during the deposition process is ≥99.99%.

17. The preparation method according to claim 15, characterized in that, After MTJ thin film deposition is completed in step S4, a first low-temperature microwave annealing is performed. Specifically, the sample is placed in a 2.45 GHz microwave annealing chamber with a vacuum degree ≥5×10⁻⁶. -8 Under Pa conditions, annealing at 100 W power for 15 minutes crystallized the FeCoB layer and optimized the MgO / FeCoB interface quality.

18. The preparation method according to claim 15, characterized in that, After the top electrode lead deposition is completed in step S7, a second low-temperature microwave annealing is performed. Specifically, the sample is placed in a 2.45 GHz microwave annealing chamber with a vacuum degree ≥5×10⁻⁶. - 8 Under the condition of Pa, annealing at 50 W power for 8 minutes eliminates lead deposition stress and improves interlayer bonding.

19. A method for acquiring bioelectrical signals, characterized in that, Using the flexible bioelectric surface dry electrode according to any one of claims 1 to 12 includes the following steps: The microneedle bio-contact layer of the dry electrode is attached to the skin of the target area on the body surface, and the microneedles penetrate the stratum corneum of the skin to achieve stable contact; The pads of the bottom electrode lead are connected to a low-noise, low-temperature drift bandgap reference voltage source, and a precision current-limiting resistor of 100 kΩ to 1 MΩ is connected in series between the reference voltage source and the bottom electrode. Connect the pads of the top electrode lead and the bottom electrode lead together to the signal reading circuit; Bioelectric signals on the body surface are conducted through microneedles and conductive islands, forming a weak biocurrent flowing through the MTJ free layer. The Oersted field perpendicular to the membrane surface excited by this current directly drives the deflection of the magnetic moment of the FeCoB free layer. Based on the tunneling magnetoresistance effect, the microvolt-level bioelectric signal is converted into a resistance change signal and output to the signal reading circuit through the bus top electrode and the top electrode lead.

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