An ultra-flexible bioelectric-photoelectric signal patch sensor and its preparation method and application

Abstract

The present invention provides an ultra-flexible bioelectric-photoelectric signal patch sensor and its preparation method and application. The patch sensor includes a packaging layer, a photodetector layer, a substrate layer, an adhesion layer, and an electrode layer arranged in sequence; wherein: the packaging layer is made of a water- and oxygen-resistant material; the photodetector layer is a near-infrared organic photodiode; the substrate layer is a layer made of a water- and oxygen-resistant material; the adhesion layer is made of a hydrogel material; and the electrode layer is made of a conductive polymer material. The present invention also provides a preparation method for the above-mentioned patch sensor and its application as a device capable of simultaneously monitoring biophotoelectric signals and bioelectric signals. The patch sensor of the present invention is ultra-thin and ultra-flexible, with a bending radius that matches the undulations of the skin surface and can reach the order of 10 μm. It can be in conformal contact, and on the basis of being light and non-sensitive, it fits the skin better to obtain the best physiological signals.

Description

Technical Field

[0001] The present invention relates to an ultra-flexible bioelectric-photoelectric signal patch sensor and a preparation method and application thereof, belonging to the technical field of flexible electronics. Background Art

[0002] Flexible wearable electronics are a crucial piece of hardware for future health monitoring. Unlike the rigidity of traditional electronic devices, emerging wearable electronics possess distinct mechanical properties, including foldability, bendability, and stretchability. These properties are poised to revolutionize many fields, including health monitoring. Compared to traditional measurement methods, flexible electronics offer unique advantages, including high skin adhesion, low contact impedance, suitability for long-term wear, and minimal harm to the human body. These advantages will meet the demands of future wearable health monitoring devices.

[0003] Exercise heart rate, blood oxygen saturation, and other indicators are important indicators for monitoring cardiovascular health. Photoplethysmography (PPG) is a common method for monitoring human pulse information. It uses the periodic changes in blood volume within the skin under the action of the pulse. The incident light transmits the volume change signal to the photodetector through transmission or reflection, thereby obtaining pulse blood flow change information. When the signals generated by two different wavelengths of incident light are analyzed proportionally, considering the different molar extinction coefficients of oxygenated and deoxygenated hemoglobin at specific wavelengths, blood oxygen saturation characteristics can be further extracted. Flexible PPG sensors will become an indispensable hardware component in cardiovascular health monitoring, and intrinsically flexible organic photodiodes will be a suitable choice for flexible PPG sensors. In addition, because near-infrared light can effectively pass through skin tissue without significant attenuation, it can collect signals from a wider area and is less susceptible to interference from motion or other information, thereby effectively improving test accuracy and signal stability. For measuring pulse wave signals and blood oxygen saturation, light around 940nm provides the best signal within the near-infrared window utilized by PPG sensors. However, selecting organic semiconductor materials with good photoelectric response in this window is difficult. In particular, the inability to build a PPG sensor with performance comparable to silicon photodiodes while also possessing extreme flexibility has limited further development in this field.

[0004] Compared to the pulse signal provided by PPG based on photoelectric volume changes, the electrocardiogram (ECG) signal provides a pulse signal based on bioelectricity. The combination of the two can improve the accuracy and reliability of pulse information. In addition, PPG and ECG signals carry additional information about arterial blood flow and heart beats in blood vessels, respectively. By analyzing and fitting the two signals, more health monitoring indicators can be obtained. For example, the difference between the R peak of the ECG and the contraction peak in the PPG signal is called the pulse arrival time (PAT). This indicator has a linear relationship with blood pressure and can be used to estimate the blood pressure changes of the subject by establishing a model. However, due to the limitations of physical form and conductivity, few commercial ECG electrodes such as Ag / AgCl gel electrodes or various metal dry electrodes can be integrated with flexible PPG sensors to form flexible small-sized sensors to achieve the above functions.

[0005] In summary, the construction of an integrated PPG-ECG patch that simultaneously monitors photoelectric and electrical biosignals on a single ultra-flexible patch sensor would be crucial for the design of a new generation of wearable health monitoring devices. However, currently, there are technical bottlenecks in developing both high-performance near-infrared-sensitive ultra-flexible photodetectors and flexible electrodes with high conductivity and low interfacial impedance. Summary of the Invention

[0006] To solve the above technical problems, the purpose of the present invention is to provide an ultra-flexible bioelectric-photoelectric signal patch sensor and its preparation method and application, so as to achieve close fit with human skin and real-time and accurate monitoring of health parameters, and provide prototype demonstration and ideas for future ultra-flexible wearable health monitoring devices.

[0007] To achieve the above-mentioned object, the present invention first provides an ultra-flexible bioelectric-photoelectric signal patch sensor, which comprises an encapsulation layer, a photodetector layer, a substrate layer, an adhesion layer, and an electrode layer arranged in sequence; wherein:

[0008] The material of the encapsulation layer is water and oxygen resistant material;

[0009] The photodetector layer is a near-infrared organic photodiode;

[0010] The substrate layer is made of a water and oxygen resistant material;

[0011] The material of the adhesion layer is hydrogel material;

[0012] The electrode layer is made of conductive polymer material.

[0013] In the above patch sensor, preferably, the encapsulation layer satisfies one or both of the following conditions:

[0014] The material of the encapsulation layer is Parylene or Teflon;

[0015] The thickness of the encapsulation layer is 0.1 μm-5 μm.

[0016] In the above patch sensor, the photodetector layer preferably includes external wires and a positive electrode, a first interface layer, a photosensitive layer, a second interface layer, and a negative electrode disposed in that order. More preferably, the positive electrode has a thickness of 10-200 nm, the first interface layer has a thickness of 5-50 nm, the photosensitive layer has a thickness of 50-300 nm, the second interface layer has a thickness of 5-50 nm, and the negative electrode has a thickness of 50-200 nm.

[0017] In the above patch sensor, preferably, the photodetector layer satisfies one or a combination of two or more of the following conditions:

[0018] The thickness of the photodetector layer is less than 1 μm;

[0019] The photosensitive layer is a near-infrared sensitive organic bulk heterojunction;

[0020] The first interface layer and the second interface layer are made of oxide or polymer;

[0021] The material of the positive electrode is metal;

[0022] The negative electrode is a transparent conductive film;

[0023] The external connecting wire is a metal film.

[0024] According to a specific embodiment of the present invention, preferably, the photosensitive layer is a PTB7-Th polymer donor and a COTIC-4F non-fullerene small molecule and PC 71 More preferably, the PTB7-Th polymer donor, COTIC-4F non-fullerene small molecule, PC 71 The mass ratio of BM fullerene mixed acceptor is 1:(1-2):(0.1-1).

[0025] According to a specific embodiment of the present invention, preferably, the oxides used as the materials of the first interface layer and the second interface layer include ZnO, SnO2, MoO x , NiO, one or more combinations thereof, the polymer includes PEI (polyethyleneimine), PEIE (polyethoxyethyleneimine), PEDOT:PSS, one or more combinations thereof; more preferably, the first interface layer is MoO x Hole transport layer, the second interface layer is a PEIE-Zn electron transport layer. MoO x represents manganese oxides, such as MnO, MnO2, Mn2O3, but not limited thereto.

[0026] According to a specific embodiment of the present invention, preferably, the metal used as the positive electrode material includes silver, aluminum, or a combination of the two.

[0027] According to a specific embodiment of the present invention, preferably, the material of the negative electrode is ITO conductive glass.

[0028] According to a specific embodiment of the present invention, preferably, the external wire is a Cr / Au wire patterned and grown on a cPI substrate.

[0029] In the above patch sensor, preferably, the substrate layer satisfies one or a combination of two or more of the following conditions:

[0030] The substrate layer is a flat layer made of water and oxygen resistant material;

[0031] The material of the substrate layer is parylene, PI or PET;

[0032] The surface of the substrate layer is planarized with photoresist;

[0033] The thickness of the substrate layer is 0.5 μm-5 μm.

[0034] In the above patch sensor, preferably, the adhesion layer satisfies one or a combination of two or more of the following conditions:

[0035] The material of the adhesive layer is a hydrogel with adhesiveness and biocompatibility;

[0036] The thickness of the adhesion layer is 10 μm-50 μm.

[0037] According to a specific embodiment of the present invention, preferably, the hydrogel is a peelable ultra-thin hydrogel; more preferably, the peelable ultra-thin hydrogel comprises a hydrogel film and two hydrophobically modified release films attached to the surfaces on both sides of the hydrogel film. Among them, the peelable ultra-thin hydrogel can be the peelable ultra-thin hydrogel disclosed in CN113754897A (application number 202110980155.6, application date 2021.08.25, invention name is a peelable ultra-thin hydrogel, preparation method and application). The preparation method of the peelable ultra-thin hydrogel comprises: applying a one-step reaction type hydrogel stock solution between two substrates, and obtaining the peelable ultra-thin hydrogel by double-roll pressing; the substrate is a hydrophobically modified release film. One-step reaction type hydrogel stock solution polyacrylamide / sodium alginate hydrogel.

[0038] In the above patch sensor, preferably, the electrode layer satisfies one or a combination of two or more of the following conditions:

[0039] The material of the electrode layer is a conductive polymer material with high conductivity, low skin interface impedance and biocompatibility;

[0040] The length and width of the electrodes on the electrode layer are 1mm-50mm respectively;

[0041] The spacing between the electrodes on the electrode layer is 1 cm to 5 cm;

[0042] The thickness of the electrode layer is 1 μm-50 μm.

[0043] According to a specific embodiment of the present invention, the conductive polymer is preferably PEDOT:PSS modified with EG and LiTFSI. The high-performance conductive polymer used in the present invention is PEDOT:PSS modified with EG and LiTFSI, which can be patterned on a hydrogel (high water content) substrate while maintaining high conductivity. It also exhibits excellent skin interface impedance and conductivity, which is difficult to achieve with other conductive polymers.

[0044] The EG and LiTFSI modified PEDOT:PSS is the PEDOT:PSS / EG / LiTFSI conductive ink described in the invention patent application with application number 202210227885.3 (invention name: PEDOT:PSS / EG / LiTFSI conductive ink and ultra-flexible electrode and electrophysiological signal monitoring method), that is, the EG and LiTFSI modified PEDOT:PSS is PEDOT:PSS / EG / LiTFSI conductive ink, which is obtained by modifying the PEDOT:PSS solution with ethylene glycol and lithium bistrifluoromethanesulfonyl imide, wherein the viscosity of the PEDOT:PSS / EG / LiTFSI conductive ink is 1-1000cP and the surface tension is 30-40mN / m.

[0045] According to a specific embodiment of the present invention, preferably, the modification comprises the following steps:

[0046] Step 1: mixing ethylene glycol with the PEDOT:PSS solution to obtain a PEDOT:PSS / EG solution; wherein the amount of ethylene glycol is 2-10 wt %, preferably 5-10 wt %, and more preferably 5 wt % or 10 wt % of the mass of the PEDOT:PSS solution;

[0047] Step 2: Add LiTFSI to the PEDOT:PSS / EG solution to obtain the PEDOT:PSS / EG / LiTFSI conductive ink, wherein the amount of LiTFSI is 5-45.5wt% of the solid content of the PEDOT:PSS solution, preferably 10-45.5wt%, and more preferably 45.5wt%.

[0048] The present invention achieves a simultaneous and substantial improvement in conductivity and stretchability by simultaneously adding EG and LiTFSI to the PEDOT:PSS system and controlling the corresponding ratios. The conductivity is significantly better than that of a system modified with either EG or LiTFSI alone.

[0049] This invention mixes different concentrations of EG and LiTFSI with a PEDOT:PSS aqueous solution. By controlling the ratio of EG to LiTFSI, a PEDOT:PSS / EG / LiTFSI conductive ink with viscosity, dispersibility, stability, and surface tension suitable for inkjet printing can be obtained. Furthermore, by adjusting the stirring time, filtration process, and ultrasonication time, excellent printing effects can be achieved.

[0050] The PEDOT:PSS / EG / LiTFSI conductive ink obtained by the present invention has a viscosity of 1-1000 cP and a surface tension of 30-40 mN / m. After being placed at room temperature for 60 days, its properties do not change significantly and the nozzle will not be clogged.

[0051] According to a specific embodiment of the present invention, by controlling the ratio of EG to LiTFSI, a long-range ordered structure of the conductive domain (PEDOT) can be achieved, generating a partial crystalline network within the PSS phase, thereby ensuring good conductivity even under high strain. Preferably, the mass ratio of EG to LiTFSI is 0.04-2, preferably 0.109-1, and more preferably 0.109, 0.167, 0.219, 0.25, 0.333, 0.5, or 1.

[0052] According to a specific embodiment of the present invention, preferably, in step 1, ethylene glycol and the PEDOT:PSS solution are mixed by stirring. Preferably, the stirring speed is controlled to 900 r / min and the time is 15 min.

[0053] According to a specific embodiment of the present invention, the total thickness of the patch sensor is preferably 10 μm to 100 μm, more preferably 10 μm to 90 μm. The patch sensor of the present invention is on the order of 10 μm and can achieve a bending radius on the order of μm, which is compatible with skin wrinkles. Existing flexible devices are only flexible in a macroscopic state, with bending radii ranging from millimeters to centimeters, which do not have the application value of ultra-flexible patch sensors.

[0054] The present invention also provides a method for preparing the ultra-flexible bioelectric-photoelectric signal patch sensor, which comprises the following steps:

[0055] depositing a sacrificial layer on a rigid substrate;

[0056] Depositing a substrate layer on the sacrificial layer, and performing negative photoresist coating, exposure, and baking to planarize the substrate;

[0057] preparing a photodetector layer on the substrate layer;

[0058] performing packaging on the photodetector layer;

[0059] preparing an adhesion layer, and then patterning a conductive polymer material on the surface of the adhesion layer to form an electrode layer;

[0060] The packaged device is peeled off from the rigid substrate to form a flexible film, and the patterned adhesive layer is laminated to obtain the ultra-flexible bioelectric-photoelectric signal patch sensor.

[0061] In the above preparation method, preferably, the material of the sacrificial layer includes one or a combination of two or more fluorine-containing compounds such as perfluoroalkane, perfluorosilane and fluorine-containing acrylic polymer.

[0062] In the above preparation method, preferably, after the substrate layer is deposited on the sacrificial layer, it is necessary to perform negative photoresist coating, exposure, and baking to make the substrate flat; photoresist is usually relatively hard, but after being flattened by the method of the present invention, it can have the characteristics of high temperature resistance and water and oxygen resistance, and at the same time have high flexibility to meet the needs of patch sensors; wherein, the negative photoresist coating, exposure, and baking can be carried out in the following specific manner: first, the photoresist is diluted with a developer, and then vortex mixing and ultrasonic debubbling are performed. After coating, it is exposed to 365nm ultraviolet light for more than 3 minutes, baked at 95°C, and hardened in a glove box at 180°C. The final photoresist thickness is less than 1μm and is flexible.

[0063] In the above preparation method, preferably, the photodetector layer is prepared by the following steps: forming a negative electrode by magnetron sputtering deposition, preparing a second interface layer (for example: an electron transport interface layer) and a photosensitive layer (for example: an organic bulk heterojunction layer) by spin coating, and preparing a first interface layer (for example: a hole transport interface layer) and a positive electrode by thermal evaporation deposition.

[0064] In the above preparation method, preferably, the encapsulation layer is formed by spin coating.

[0065] In the above preparation method, preferably, the adhesive layer is formed by the following method: using a double-roll coater with adjustable spacing, taking a hydrogel stock solution that has not yet gelled and coating it, and then reacting to obtain an ultra-thin hydrogel adhesive layer.

[0066] In the above preparation method, preferably, before patterning, the electrode layer is formed by applying a precursor of the conductive polymer onto the adhesive layer by a printing method such as printing or coating, and then performing patterning.

[0067] In the above preparation method, preferably, the patterning of the above electrode layer is performed by printing, scraping through a mask, or dripping through a mask.

[0068] The present invention also provides an application of the above-mentioned ultra-flexible bioelectric-photoelectric signal patch sensor as a device capable of simultaneously monitoring biophotoelectric signals and bioelectric signals.

[0069] The beneficial effects of the present invention are:

[0070] 1. The ultra-flexible bioelectric-photoelectric signal patch sensor of this invention is ultra-thin and ultra-flexible, with a bending radius that matches the undulations of the skin surface, reaching the order of 10 μm. It also provides conformal contact, making it lightweight and non-sensitive while conforming more closely to the skin to obtain optimal physiological signals. Existing methods have not yet been able to create a patch sensor that combines both high flexibility and high performance.

[0071] 2. The photosensitivity of the intrinsically flexible high-performance organic photodetector layer of the ultra-flexible bioelectric-photoelectric signal patch sensor of the present invention is comparable to that of silicon photodiodes, reaching 0.55A / W@940nm. It can replace rigid silicon photodiodes in scenarios where flexible photodetectors are required, and the preparation method is suitable for the flexibility of various rigid detectors.

[0072] 3. The breathability and biocompatibility of ultra-thin hydrogels make long-term wear possible, replacing various commercial viscose products currently in use. They are thinner, reaching 10μm in thickness, and have better biocompatibility.

[0073] 4. High-performance conductive polymer electrodes can greatly reduce the electrode area and spacing. Due to their ultra-high conductivity and ultra-low skin interface impedance, the conductivity is higher than 3000S / cm and the skin interface impedance is less than 5kΩ@100Hz. Bioelectric signals such as electrocardiogram can be collected in an extremely small size, which is conducive to the miniaturization of related devices. BRIEF DESCRIPTION OF THE DRAWINGS

[0074] Figure 1 This is a schematic diagram of the structure of the ultra-flexible bioelectric-photoelectric signal patch sensor;

[0075] Figure 2 A representative physical picture of the ultra-flexible bioelectric-photoelectric signal patch sensor;

[0076] Figure 3 This is a flow chart of the preparation process of the ultra-flexible bioelectric-photoelectric signal patch sensor;

[0077] Figure 4A and Figure 4B This is a photoelectric performance test result diagram of the ultra-flexible bioelectric-photoelectric signal patch sensor;

[0078] Figure 5 This is a biocompatibility test diagram of the adhesion layer;

[0079] Figure 6 This is the skin contact impedance performance diagram of the electrode layer;

[0080] Figure 7A and Figure 7B Performance diagram of the ultra-flexible bioelectric-photoelectric signal patch sensor for PPG-ECG applications;

[0081] Figures 8A-8D Performance comparison results of ultra-flexible bioelectric-photoelectric signal patch sensor, commercial PPG sensor and ECG electrode.

[0082] Figure 9 Comparative microscopic images of patch sensors with different thicknesses. DETAILED DESCRIPTION

[0083] In order to have a clearer understanding of the technical features, objectives and beneficial effects of the present invention, the technical solution of the present invention is now described in detail below, but it should not be understood as limiting the scope of implementation of the present invention.

[0084] Figure 1 Schematic diagram of the structure of an ultra-flexible bioelectric-photoelectric signal patch sensor according to a specific embodiment of the present invention.

[0085] In some embodiments, reference Figure 1 As shown, the ultra-flexible bioelectric-photoelectric signal patch sensor includes:

[0086] Encapsulation layer 1;

[0087] The photodetector layer 2 is located below the packaging layer 1 and includes a photosensitive layer 21, a first interface layer 221, a second interface layer 222, a positive electrode 23, a negative electrode 24, and an external wire 25.

[0088] a substrate layer 3, located below the photodetector layer 2;

[0089] an adhesive layer 4 located below the substrate layer 3;

[0090] The electrode layer 5 is located below the adhesion layer 5 .

[0091] In some embodiments, encapsulation layer 1 is an ultra-thin, water- and oxygen-resistant film that protects underlying materials from water, oxygen, and external contaminants at a minimal thickness. Specifically, materials for encapsulation layer 1 include, but are not limited to, parylene and Teflon. The thickness of encapsulation layer 1 ranges from 0.1 μm to 5 μm.

[0092] In some specific embodiments, the photodetector layer 2 is a high-performance organic photodetector with intrinsic flexibility and a responsivity comparable to that of commercial silicon photodiodes, which can ensure the quality of biophotoelectric signals. 71 The organic bulk heterojunction composed of BM fullerene mixed acceptors (the mass ratio of the three is preferably 1: (1-2): (0.1-1)), the first interface layer 221 is MoO x The hole transport layer, the second interface layer 222 is a PEIE-Zn electron transport layer, the material of the positive electrode 23 is Ag, the material of the negative electrode 24 is ITO conductive glass, and the external wire 25 is a Cr / Au wire patterned on the cPI substrate. The thickness of the photodetector layer 2 is less than 1μm (to show the specific composition of the photodetector layer 2, Figure 1 In the schematic method, Figure 1 The dimensions of each layer shown in the figure do not represent the actual size relationship between the layers).

[0093] In some specific embodiments, the substrate layer 3 is a water- and oxygen-resistant and flat transparent film, preferably Parylene, which is then planarized using a negative photoresist. The thickness of the substrate layer 3 is between 0.5 μm and 5 μm.

[0094] In some specific embodiments, the adhesion layer 4 is an ultra-thin hydrogel material, which is made of acrylamide / sodium alginate hydrogel and has a thickness of 10 μm-50 μm.

[0095] In some specific embodiments, the electrode layer 5 is a patterned conductive polymer material, and its material is PEDOT:PSS modified with EG and LiTFSI; the thickness of the electrode layer is 1 μm-50 μm.

[0096] Figure 2 The following is a representative physical picture of the ultra-flexible bioelectric-photoelectric signal patch sensor according to the present invention. Figure 2 It can be seen that the bending radius of the ultra-flexible bioelectric-photoelectric signal patch sensor of the present invention matches the undulations of the skin surface, and can achieve good fit with the human body.

[0097] Some specific embodiments of the present invention also provide a method for preparing an ultra-flexible bioelectric-photoelectric signal patch sensor, the process of which is as follows: Figure 3 As shown, it includes the following steps:

[0098] Step S31: depositing a sacrificial layer on a rigid substrate;

[0099] In some specific embodiments, the method for forming a sacrificial layer is as follows: depositing a fluorine-containing compound such as perfluoro(silane) alkane, a fluorine-containing acrylic polymer, etc. on a clean rigid substrate, the rigid substrate can be a 1mm thick soda-lime silica glass, and the material of the sacrificial layer is a mixed solution made by mixing the Delo series coating solution of Shenzhen Zhongfu Technology Co., Ltd. with the 3M Novec series fluorine solution (the mass ratio of the two is preferably 1:5); the deposition method is spin coating, the spin coating speed is 5000rpm, and the time is 30s. The use of this sacrificial layer can optimize the surface energy and surface state of the substrate, and control the bonding force between the flexible substrate and the rigid substrate without damaging the quality of the subsequent chemical vapor deposition film. It can achieve a subsequent heat treatment temperature of up to 250°C and repeated high vacuum environment without damaging the integrity and removability of the flexible substrate. Among them, the concentration of the mixed solution is preferably 10-50%, which can avoid the generation of ripples when the concentration is too high and affect the flatness of subsequent operations.

[0100] Step S32: depositing a substrate layer on the sacrificial layer, and performing photoresist coating, exposure, and baking to planarize the substrate;

[0101] In some specific embodiments, the method for forming the substrate layer is: chemical vapor deposition of a substrate layer of polyparaxylene on a sacrificial layer, using an SCS parylene coater for deposition, and using Parylene C dimer as a raw material; after deposition, a negative photoresist is applied, exposed, and baked to planarize the substrate, wherein SU-8 negative photoresist is specifically used and diluted with a developer for application, and the dilution ratio is less than 1:1. The thickness of the photoresist film after application is less than 1 μm to obtain flexibility; then, 365 nm ultraviolet exposure for 3 minutes, baking at 95° C., and hardening at 180° C. are performed.

[0102] Step S33: preparing a photodetector layer on the substrate layer;

[0103] In some specific embodiments, the method for forming the photodetector layer is: depositing a transparent conductive film cathode by magnetron sputtering, preparing an electron transport interface layer (second interface layer) and an organic bulk heterojunction layer by spin coating, and depositing a hole transport interface layer (first interface layer) and a cathode by thermal evaporation;

[0104] Among them, magnetron sputtering uses an ITO target material AC source of 100W sputtering for 20 minutes to deposit a transparent conductive film of about 100nm. The physical position of this layer must be controlled in the middle of the packaging layer and the substrate layer to relieve the internal stress of the ultra-flexible device in the bending state;

[0105] The electron transport interface layer is prepared by spin coating a PEIE-Zn chelate precursor at a rotation speed of 3000 rpm for 60 s and a post-treatment temperature of 180°C. The preparation method of PEIE-Zn can refer to the preparation method of PEI-Zn.

[0106] Among them, the organic bulk heterojunction uses PTB7-Th polymer donor and COTIC-4F non-fullerene and PC 71 The BM fullerene mixed acceptor and chlorobenzene mixed solution (the mass ratio of the three is preferably 1: (1-2): (0.1-1), and the total concentration of the chlorobenzene mixed solution is 10-30 g / L) was spin-coated at a speed of 1000 rpm for 30 s.

[0107] The raw materials used for thermal evaporation of the hole transport interface layer and the positive electrode are MoO3 powder and Ag particles, respectively. The vacuum degree during evaporation is less than 2×10 -4 Pa.

[0108] Step S34: performing packaging on the photodetector layer;

[0109] In some specific embodiments, the method for forming the encapsulation layer is: spin-coating the fluorinated polymer Teflon (dissolved in Novec FC series fluorinated liquid, inert to the organic layer below and does not damage its performance) and then taking it out and placing it in a parylene coater, using chemical vapor deposition to deposit polyparaxylene to achieve water and oxygen isolation effects, and pasting external wires 25 before encapsulation to facilitate subsequent measurement scenarios.

[0110] Step S35: preparing an ultrathin hydrogel substrate;

[0111] In some specific embodiments, the method for forming the adhesion layer is: using a double-roll coater with adjustable spacing, taking a hydrogel stock solution that has not yet gelled and reacting it to obtain an ultra-thin hydrogel adhesion layer, wherein the component of the hydrogel stock solution that has not yet gelled is acrylamide / sodium alginate hydrogel.

[0112] Step S36: performing patterning of a conductive polymer electrode on the surface of the ultrathin hydrogel;

[0113] In some specific embodiments, the electrode layer is formed by patterning a conductive polymer electrode by masked drop coating onto the surface of the adhesion layer; the electrode component is a PEDOT:PSS conductive polymer modified with EG and LiTFSI. The present invention has previously attempted to use conventional PEDOT:PSS, but due to its hydrophilicity and water solubility, it swells on the hydrogel substrate, making patterning impossible. The PEDOT:PSS conductive polymer modified with EG and LiTFSI employed in the present invention, due to the addition of EG and LiTSI, introduces hydrogen and ionic crosslinking within the conductive polymer, resulting in enhanced conductivity and reduced water solubility, enabling patterning and stable stability on the surface of ultrathin hydrogels.

[0114] Step S37: The encapsulated flexible device can be peeled off from the rigid substrate to form a flexible film, which is then aligned and bonded to the hydrogel adhesive layer patterned with the conductive polymer electrode, thereby obtaining an ultra-flexible bioelectric-photoelectric signal patch sensor;

[0115] In some specific embodiments, due to the presence of the sacrificial layer on the rigid substrate, the encapsulated flexible device can be peeled off from the rigid substrate to become a flexible film without affecting the bonding and functionality between the functional layers in the flexible film; after alignment and bonding with the hydrogel adhesion layer patterned with conductive polymer electrodes, it becomes an ultra-flexible bioelectric-photoelectric signal patch sensor, which can realize the simultaneous monitoring of bioelectric-photoelectric signals at the same position by a single sensor.

[0116] Example 1

[0117] This embodiment provides an ultra-flexible bioelectric-photoelectric signal patch sensor having the following features: Figure 1 The structure shown in FIG. 1 is a patch sensor comprising:

[0118] Encapsulation layer: Made of polyparaxylene, with a thickness of 1.5 μm;

[0119] Photodetector layer: including positive electrode, first interface layer, photosensitive layer, second interface layer, negative electrode, external wire; the positive electrode is made of Ag with a thickness of 100nm; the first interface layer is MoO x The hole transport interface layer has a thickness of 7.5 nm; the photosensitive layer is composed of PTB7-Th polymer donor, COTIC-4F non-fullerene small molecule and PC 71The organic bulk heterojunction is composed of a BM fullerene mixed acceptor (the mass ratio of the three is 1:(1-2):(0.1-1)), with a thickness of 120nm; the second interface layer is a PEIE-Zn electron transport layer with a thickness of 30nm; the negative electrode is made of ITO conductive glass with a thickness of 100nm; the external wire is a Cr / Au wire patterned on a cPI substrate with a thickness of 100nm;

[0120] Substrate layer: Made of polyparaxylene, it is planarized with negative photoresist. The thickness of the substrate layer is 2μm.

[0121] Adhesion layer: a hydrogel film made of acrylamide / sodium alginate hydrogel (prepared according to Example 1 of CN113754897A), with a thickness of 10 μm;

[0122] Electrode layer: The material is PEDOT:PSS modified with EG and LiTFSI, prepared according to Example 4 of the invention patent application with application number 202210227885.3;

[0123] The thickness of the ultra-flexible bioelectric-photoelectric signal patch sensor is about 15 μm.

[0124] The performance test of the ultra-flexible bioelectric-photoelectric signal patch sensor prepared in this example was carried out:

[0125] The ultra-flexible bioelectric-photoelectric signal patch sensor obtained in Example 1 was subjected to a photoelectric device performance test and an electrical performance test, and a biocompatibility tester for the adhesion layer to verify its reliability as a bioelectronic device.

[0126] Figure 4A and Figure 4B The photoelectric performance test results are shown in the figure. Figure 4A ) and 940nm( Figure 4B ) wavelength monochromatic light, the response of the patch sensor of Example 1 reaches 0.4A / W@660nm and 0.55A / W@940nm under a bias of -1V, while the response of the silicon photodiode is 0.42A / W@660nm and 0.6A / W@940nm. It can be seen that the optoelectronic device performance of the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 can match that of commercial silicon photodiodes.

[0127] The biocompatibility of the hydrogel film was characterized by human embryonic lung fibroblast cell line, IMR-90 culture medium, and DAPI fluorescence staining. Figure 5As shown. The experimental process is as follows: IMR-90 human embryonic lung fibroblast cell line (Wuhan Procell Life Science Co., Ltd.) was maintained in IMR-90 medium (Wuhan Procell Life Science Co., Ltd.); cells were passaged every four days at a confluence of 60%-80% by incubating with 0.25% trypsin containing EDTA (Gibco) at 37°C for 3 minutes and re-seeding on a tissue culture (TC) treated 6-well plate at a ratio of 1:3; cells were cultured in an incubator at 37°C with 5% CO2; in order to produce conditioned IMR-90 medium, the hydrogel membrane was cut to the precise size of the culture surface and then immersed in 6 ml of IMR-90 medium for at least 48 hours; in order to exert the potential effect of the hydrogel on the cells, the hydrogel membrane was cultured in IMR-90 medium at 4°C until the cells were characterized. By Figure 5 It can be seen that the cells survive well, proving that using hydrogel as the adhesion layer has good safety and reliability.

[0128] Figure 6 is the electrode layer skin contact impedance performance diagram. Figure 6 As shown, the conductivity of the electrode layer of the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 is higher than 3000 S / cm, and the skin interface impedance is less than 5 kΩ@100 Hz, which is better than the currently commonly used Ag / AgCl gel electrode.

[0129] Figure 7A and Figure 7B The performance diagram of the ultra-flexible bioelectric-photoelectric signal patch sensor for PPG-ECG applications. Figure 7A and Figure 7B It can be seen that the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 can realize the monitoring of PPG-ECG signals. Due to its advanced photoelectric response, conductivity and skin interface impedance, the patch sensor can accurately obtain the PPG-ECG signal on the chest. Among them, the photoelectric sensor is only 0.04 cm 2 , the size of a single electrode is only 0.09cm 2 , the distance between electrodes is only 1 cm.

[0130] Comparative Example 1

[0131] The performance of the common finger clip blood oxygen sensor and Ag / AgCl gel electrode on the market was compared under the same conditions. The results are as follows: Figures 8A-8D As shown. Figures 8A-8DIt can be seen that under equivalent LED module current and signal amplification, the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 has a higher PPG signal strength than a commercial finger-clip blood oxygen sensor. At the same test location, the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 has a higher ECG signal strength than a commercial Ag / AgCl gel electrode.

[0132] In addition to its superior performance, the ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 also has unparalleled advantages over similar products in terms of test position and sensor size.

[0133] Comparative Example 2

[0134] Compared with the commonly used packaging materials on the market as packaging layers, including acrylic resin, epoxy resin, etc., as well as common flexible packaging materials PDMS, SEBS, etc.

[0135] Among them, acrylic resin requires ultraviolet light curing, and epoxy resin requires AB glue reaction, both of which will have adverse effects on the organic semiconductor layer.

[0136] To ensure the flexibility of patch sensors, spin-coating with materials such as PDMS and SEBS is recommended. However, PDMS silicone rubbers generally have high oxygen permeability, making them in thin layers unable to guarantee the long-term operation of water- and oxygen-sensitive organic semiconductor materials. Common solvents for SEBS, such as toluene and chloroform, can also damage the underlying organic semiconductor. Therefore, a mild preparation environment, the selection of an encapsulation layer solvent orthogonal to the organic semiconductor, and a balance between encapsulation layer thickness and mechanical flexibility are crucial.

[0137] In this comparative example, the performance of the lower organic photoelectric device seriously degraded after being affected by the upper packaging material. Under the same lighting conditions, the photocurrent at 0V decreased by 60%, and the photovoltage decreased by about 20%.

[0138] Flexibility comparison experiment

[0139] The ultra-flexible bioelectric-photoelectric signal patch sensor of Example 1 is ultra-thin and ultra-flexible, and its bending radius matches the undulations of the skin surface and can reach the order of 10 μm, and can conformally contact ( Figure 9 ), on the basis of being light and imperceptible, it fits the skin better to obtain the best physiological signals.

[0140] In contrast, patch sensors are on the order of 100μm. When the thickness increases, gaps will appear between the patch sensor and the skin, which will affect the adhesion of the patch and the detection performance of the sensor.

Claims

1. An ultra-flexible bioelectric-photoelectric signal patch sensor, comprising an encapsulation layer, a photodetector layer, a substrate layer, an adhesion layer, and an electrode layer arranged in sequence; wherein: The material of the encapsulation layer is water and oxygen resistant material; The photodetector layer is a near-infrared organic photodiode; The substrate layer is made of a water and oxygen resistant material; The material of the adhesion layer is hydrogel material; The material of the electrode layer is a conductive polymer material; Wherein, the material of the encapsulation layer is parylene or fluorinated polymer; The photodetector layer includes an external wire and a positive electrode, a first interface layer, a photosensitive layer, a second interface layer, and a negative electrode arranged in sequence; The photosensitive layer is a near-infrared sensitive organic phase heterojunction; the photosensitive layer is a PTB7-Th polymer donor and COTIC-4F non-fullerene small molecules and PC 71 Organic bulk heterojunction composed of BM fullerene mixed acceptors; The material of the first interface layer and the second interface layer is oxide or polymer; the oxide includes ZnO, SnO2, MoO x , NiO, or a combination of two or more thereof, the polymer comprising PEI, PEIE, PEIE-Zn, or PEDOT:PSS; The material of the electrode layer is a conductive polymer material with high conductivity, low skin interface impedance and biocompatibility; the conductive polymer is PEDOT:PSS modified by EG and LiTFSI.

2. The patch sensor according to claim 1, wherein The thickness of the encapsulation layer is 0.1 μm-5 μm.

3. The patch sensor according to claim 1, wherein: The photodetector layer satisfies one or a combination of two or more of the following conditions: The thickness of the photodetector layer is less than 1 μm; The PTB7-Th polymer donor, COTIC-4F non-fullerene small molecule, PC 71 The mass ratio of BM fullerene mixed acceptor is 1:(1-2):(0.1-1); The first interface layer is MoO x hole transport layer, the second interface layer is a PEIE-Zn electron transport layer; The positive electrode is a metal; The negative electrode is a transparent conductive film; The external connecting wire is a metal film.

4. The patch sensor according to claim 3, wherein: The positive electrode is silver, aluminum or a combination of both.

5. The patch sensor according to claim 3, wherein: The material of the negative electrode is ITO conductive glass.

6. The patch sensor according to claim 3, wherein: The external wires are Cr / Au wires patterned and grown on a cPI substrate.

7. The patch sensor according to claim 1, wherein: The thickness of the positive electrode is 10-200 nm, the thickness of the first interface layer is 5-50 nm, the thickness of the photosensitive layer is 50-300 nm, the thickness of the second interface layer is 5-50 nm, and the thickness of the negative electrode is 50-200 nm.

8. The patch sensor according to claim 1, wherein: The substrate layer satisfies one or a combination of two or more of the following conditions: The substrate layer is a flat layer made of water and oxygen resistant material; The material of the substrate layer is parylene, PI or PET; The surface of the substrate layer is planarized with photoresist; The thickness of the substrate layer is 0.5 μm-5 μm.

9. The patch sensor according to claim 1, wherein: The adhesive layer satisfies one or a combination of two or more of the following conditions: The material of the adhesive layer is a hydrogel with adhesiveness and biocompatibility; The thickness of the adhesion layer is 10 μm-50 μm.

10. The patch sensor according to claim 9, wherein: The hydrogel is a peelable ultra-thin hydrogel.

11. The patch sensor according to claim 10, wherein: The peelable ultra-thin hydrogel comprises a hydrogel film and two hydrophobically modified release films attached to both side surfaces of the hydrogel film.

12. The patch sensor according to claim 1, wherein: The electrode layer satisfies one or a combination of two or more of the following conditions: The length and width of the electrodes on the electrode layer are 1 mm to 50 mm respectively; The distance between the electrodes on the electrode layer is 1 cm to 5 cm; The thickness of the electrode layer is 1 μm-50 μm.

13. The patch sensor according to claim 1, wherein: The EG and LiTFSI-modified PEDOT:PSS is a PEDOT:PSS / EG / LiTFSI conductive ink, which is obtained by modifying a PEDOT:PSS solution with ethylene glycol and lithium bis(trifluoromethanesulfonyl)imide. The PEDOT:PSS / EG / LiTFSI conductive ink has a viscosity of 1-1000 cP and a surface tension of 30-40 mN / m.

14. The patch sensor according to any one of claims 1 to 13, wherein: The total thickness of the patch sensor is 10μm-100μm.

15. The patch sensor according to claim 14, wherein: The total thickness of the patch sensor is 10 μm-90 μm.

16. A method for preparing the ultra-flexible bioelectric-photoelectric signal patch sensor according to any one of claims 1 to 15, comprising the following steps: depositing a sacrificial layer on a rigid substrate; Depositing a substrate layer on the sacrificial layer, and performing negative photoresist coating, exposure, and baking to planarize the substrate; preparing a photodetector layer on the substrate layer; performing packaging on the photodetector layer; preparing an adhesion layer, and then patterning a conductive polymer material on the surface of the adhesion layer to form an electrode layer; The packaged device is peeled off from the rigid substrate to form a flexible film, and the patterned adhesive layer is laminated to obtain the ultra-flexible bioelectric-photoelectric signal patch sensor.

17. The preparation method according to claim 16, wherein The material of the sacrificial layer includes one or a combination of two or more of perfluoroalkane, perfluorosilane, and fluorinated acrylic polymer.

18. The preparation method according to claim 16, wherein The photodetector layer is prepared by the following steps: forming a negative electrode by magnetron sputtering deposition, preparing a first interface layer and a photosensitive layer by spin coating, and preparing a second interface layer and a positive electrode by thermal evaporation deposition.

19. The preparation method according to claim 16, wherein The patterning process includes printing, scraping through a mask, or dispensing through a mask.

20. Use of the ultra-flexible bioelectric-photoelectric signal patch sensor according to any one of claims 1 to 15 for simultaneous monitoring of biophotoelectric signals and bioelectric signals.

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