Electrochemical sensor and biological index monitoring device

Abstract

The invention provides an electrochemical sensor and a biological index monitoring device. The electrochemical sensor comprises a flexible substrate and a working electrode arranged on at least one side of the flexible substrate; the working electrode comprises a conductive layer; the conductive layer comprises a conductive polymer PEDOT: PSS and a plasticizer; wherein the conductive polymer PEDOT: PSS is distributed in the conductive layer in the form of linear fibers. The conductive polymer PEDOT: PSS is modified by adopting the plasticizer, so that the electrochemical activity and the conductivity of the conductive polymer PEDOT: PSS can be effectively improved, the conductive polymer PEDOT: PSS can meet the application requirement of an electrochemical sensor, and the limitation that the traditional CGMS is only based on a carbon material and a noble metal material is broken through. And meanwhile, the working electrode adopting the conductive polymer PEDOT: PSS has good flexibility, still has good conductivity and electrochemical activity after being circularly bent for multiple times, and has remarkable advantages compared with a carbon-based electrode.

Description

Technical Field

[0001] The present application relates to the field of electrochemical sensing technology, and in particular to an electrochemical sensor and a biological indicator monitoring device. Background Art

[0002] Diabetes is a long-term chronic disease with no cure. How to effectively control the occurrence and development of diabetes has become a research problem that has attracted worldwide attention. Blood glucose monitoring results are of great significance for evaluating the process of glucose metabolism disorders in diabetic patients, formulating hypoglycemic plans, reflecting the effects of hypoglycemic treatment and guiding the adjustment of treatment plans. It is a rigid demand of diabetic patients. Therefore, the development of high-performance blood glucose monitoring systems is of great significance for the diagnosis and treatment of diabetes.

[0003] The implantable continuous glucose monitoring system (CGMS) developed in recent years can realize real-time dynamic monitoring of human blood glucose, providing thousands of diabetic patients with a tool for all-round blood glucose regulation. As the subcutaneous implantable probe of CGMS, the glucose electrochemical sensor is the core component with the highest barrier of CGMS, which directly determines the accuracy of blood glucose monitoring. Most of the CGMS sensors currently on the market are based on carbon materials or precious metal materials, that is, the electrodes are prepared by patterning materials on the surface of flexible substrates (such as PET, PI, etc.) through screen printing or ion sputtering. The electrode preparation materials are relatively limited. In addition, the electrode preparation process often involves the deformation process of the electrode, and the electrode is easily deformed by the fluid stress of human tissue fluid or blood after being implanted subcutaneously. Although the substrate materials such as PET and PI have good flexibility, the carbon materials and metal materials themselves are not flexible. Therefore, during the electrode deformation process, the carbon materials and metal materials are difficult to adapt to the substrate deformation and crack or even fall off, which seriously affects the sensor detection performance and may cause stress response of the body. Summary of the invention

[0004] Based on this, the present application provides an electrochemical sensor and a biological indicator monitoring device to provide a new conductive material for the electrochemical sensor and improve the flexibility of the electrode.

[0005] The first aspect of the present application provides an electrochemical sensor, comprising a flexible substrate and a working electrode arranged on at least one side of the flexible substrate; the working electrode comprises a conductive layer, the conductive layer contains a conductive polymer PEDOT:PSS and a plasticizer; wherein the conductive polymer PEDOT:PSS is distributed in the conductive layer in the form of linear fibers.

[0006] In some embodiments, the mass percentage of the plasticizer to the mass percentage of the conductive polymer PEDOT:PSS is ≥ 0.1%.

[0007] In some embodiments, the plasticizer has one of the following characteristics (1)-(5):

[0008] (1) The plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid and 1-ethyl-3-methylimidazolium tetracyanoboric acid;

[0009] (2) The plasticizer includes one or more of xylitol, glycerol and sorbitol;

[0010] (3) The plasticizer includes one or more of sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid and malic acid;

[0011] (4) The plasticizer includes one or more of Triton and fluorocarbon surfactant;

[0012] (5) The plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid, 1-ethyl-3-methylimidazolium tetracyanoboric acid, xylitol, glycerol, sorbitol, sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid, malic acid, Triton and fluorocarbon surfactant.

[0013] In some embodiments, the conductive layer further comprises a conductive material; the conductive material comprises one or more of noble metal nanomaterials, carbon nanomaterials, black phosphorus, Mxene materials, metal organic framework materials and covalent organic framework materials.

[0014] In some embodiments, the material of the flexible substrate includes one or more of polyethylene terephthalate, polyimide, polycarbonate, polyvinyl chloride, and polydimethylsiloxane.

[0015] In some embodiments, the electrochemical sensor further comprises a sensing layer located on the surface of the working electrode, wherein the sensing layer comprises an electron mediator cross-linked by a first cross-linking agent and an enzyme or a pH sensitive material for oxidizing a monitoring target.

[0016] In some embodiments, the sensing layer has at least one of the following features (6)-(10):

[0017] (6) The electron mediator includes one or more of organic conjugated small molecules and polymers thereof and transition metal complexes and polymers thereof with free amino groups or free carboxyl groups;

[0018] The transition metal includes one or more of osmium, ruthenium, iron, copper, cobalt, vanadium and manganese;

[0019] The organic conjugated small molecules include one or more of catechol and its derivatives, thiophene and its derivatives, methylene blue and its derivatives, and methylene green and its derivatives;

[0020] (7) The monitoring object includes one of glucose, lactic acid, ketone bodies, glutamic acid, glycine, uric acid, ascorbic acid, acetylcholine and cholesterol;

[0021] The enzyme for oxidizing the monitored object includes one or more of glucose oxidase, glucose dehydrogenase, lactate oxidase, lactate dehydrogenase, β-hydroxybutyrate dehydrogenase, glutamate oxidase, glycine oxidase, urate oxidase, ascorbate oxidase, acetylcholinesterase and cholesterol oxidase;

[0022] (8) the first cross-linking agent includes one or more of dialdehyde compounds, dialdehyde polymers, epoxy compounds, epoxy polymers, N,N-methylenebisacrylamide, 1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride / N-hydroxysuccinimide, genipin and dicyclohexylcarbodiimide;

[0023] (9) The pH sensitive material includes one or more of polyaniline, transition metal oxide nanomaterials and organic hydrogen ion carriers; the transition metal oxide nanomaterials include one or more of ZnO, IrO2, RuO2, WO3 and Ta2O5.

[0024] In some embodiments, when a working electrode is disposed only on one side surface of the flexible substrate, the monitored object is glucose;

[0025] The enzyme used for oxidizing the monitored object includes one or more of glucose oxidase and glucose dehydrogenase.

[0026] In some embodiments, when working electrodes are respectively disposed on at least two side surfaces of the flexible substrate, and sensing layers are respectively disposed on the surface of each of the working electrodes, the plurality of sensing layers are used to simultaneously monitor different monitoring objects.

[0027] In some embodiments, the monitoring object of the electrochemical sensor includes glucose; and the enzyme for oxidizing the monitoring object contained in the sensing layer on one side of the electrochemical sensor includes one or more of glucose oxidase and glucose dehydrogenase.

[0028] In some embodiments, a diffusion limiting layer is further included on the surface of the sensing layer; the diffusion limiting layer comprises a substrate cross-linked by a second cross-linking agent, and the substrate comprises one or more of polyvinyl pyridine, polyurethane, polyvinyl imidazole, polyacrylate, polyether urethane and polystyrene.

[0029] In some embodiments, it further comprises a biocompatible layer located on the surface of the diffusion limiting layer;

[0030] The biocompatible layer comprises one or more of polyvinyl pyrrolidone, polyvinyl alcohol, chitosan, alginate, hyaluronic acid, cellulose, collagen, gelatin, polyacrylamide, polyacrylic acid, polypropylene alcohol, sodium polystyrene sulfonate, polyethylene glycol, polypropylene glycol, phosphorylcholine grafted polymer and betaine grafted polymer.

[0031] In some embodiments, the electrochemical sensor further includes a reference electrode that cooperates with the working electrode, and the configuration of the reference electrode conductive layer is the same as that of the working electrode conductive layer.

[0032] In some embodiments, the electrochemical sensor further includes a counter electrode that cooperates with the working electrode, and the configuration of the counter electrode conductive layer is the same as that of the working electrode conductive layer.

[0033] The second aspect of the present application provides a biological indicator detection device, which includes the electrochemical sensor of the first aspect of the present application.

[0034] In some embodiments, the biological indicator detection device includes a blood glucose monitoring device.

[0035] In the electrochemical sensor and biological indicator monitoring device provided above, the conductive layer of the working electrode of the electrochemical sensor comprises a conductive polymer PEDOT:PSS and a plasticizer; by modifying the conductive polymer PEDOT:PSS with a plasticizer, the electrochemical activity and conductivity of the conductive polymer PEDOT:PSS can be effectively improved, so that the conductive polymer PEDOT:PSS can meet the application requirements of the electrochemical sensor, breaking through the limitation that the traditional CGMS is only based on carbon materials and precious metal materials. At the same time, the working electrode using the conductive polymer PEDOT:PSS has good flexibility, and still has good conductivity and electrochemical activity after multiple cycles of bending, showing significant advantages over carbon-based electrodes. BRIEF DESCRIPTION OF THE DRAWINGS

[0036] In order to more clearly illustrate the technical solutions in the embodiments of the present application and to more completely understand the present application and its beneficial effects, the following is a brief introduction to the drawings required for the description of the embodiments. Obviously, the drawings described below are only some embodiments of the present application, and those skilled in the art can obtain other drawings based on these drawings without creative work.

[0037] Figure 1 A schematic diagram of the preparation process of an electrochemical sensor with a dual-electrode structure provided in one embodiment of the present application;

[0038] Figure 2 A schematic diagram of the structure of an electrochemical sensor with a dual-electrode structure provided in one embodiment of the present application;

[0039] Figure 3 is a schematic cross-sectional view of the working electrode end;

[0040] Figure 4 Schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and ketone bodies;

[0041] Figure 5 Schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and lactate;

[0042] Figure 6 Schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and pH;

[0043] Figure 7 The sheet resistance test results of Example 1 and Comparative Example 1 are shown;

[0044] Figure 8 The cyclic voltammetric response results of the working electrode before and after bending in Example 1;

[0045] Fig. 9 The cyclic voltammetric response results of the working electrode before and after bending in Comparative Example 1;

[0046] Fig.10 is the chronoamperometric response curve of the PEDOT-based glucose sensor;

[0047] Fig.11 is the glucose standard curve of PEDOT-based glucose sensor;

[0048] Fig.12 is the chronoamperometric response curve of the carbon-based glucose sensor;

[0049] Fig.13 is the glucose standard curve of the carbon-based glucose sensor;

[0050] Fig.14 Chronoamperometric response curve of the PEDOT-based glucose sensor after coating with a diffusion limiting layer;

[0051] Fig.15 Glucose standard curve after coating the diffusion limiting layer for the PEDOT-based glucose sensor;

[0052] Fig.16 Chronoamperometric response curve of carbon-based glucose sensor after coating with diffusion limiting layer;

[0053] Fig.17 Glucose standard curve after coating the carbon-based glucose sensor with a diffusion-limiting layer. DETAILED DESCRIPTION

[0054] For ease of understanding of the present application, the present application will be described more fully below with reference to the relevant embodiments. The following provides preferred embodiments of the present application. However, the present application can be implemented in many different forms and is not limited to the embodiments described herein. On the contrary, the purpose of providing these embodiments is to make the understanding of the disclosure of the present application more thorough and comprehensive.

[0055] Unless otherwise defined, all technical and scientific terms used herein have the same meaning as those generally understood by those skilled in the art to which the present application belongs. The terms used herein in the specification of the present application are only for the purpose of describing specific embodiments and are not intended to limit the present application. The term "and / or" used herein includes any and all combinations of one or more of the related listed items.

[0056] In the present application, the technical features described in an open manner include closed technical solutions composed of the listed features, and also include open technical solutions containing the listed features.

[0057] In this application, when it comes to numerical ranges, unless otherwise specified, the above numerical ranges are deemed to be continuous and include the minimum and maximum values ​​of the range, as well as each value between such minimum and maximum values. Further, when a range refers to an integer, each integer between the minimum and maximum values ​​of the range is included. In addition, when multiple ranges are provided to describe features or characteristics, the ranges can be merged. In other words, unless otherwise specified, all ranges disclosed herein should be understood to include any and all subranges included therein.

[0058] Only some numerical ranges are specifically disclosed herein. However, any lower limit can be combined with any upper limit to form an unspecified range; and any lower limit can be combined with other lower limits to form an unspecified range, and likewise any upper limit can be combined with any other upper limit to form an unspecified range. In addition, each separately disclosed point or single value can itself be combined as a lower limit or upper limit with any other point or single value or with other lower limits or upper limits to form an unspecified range.

[0059] The temperature parameters in this application, unless otherwise specified, allow for both constant temperature treatment and treatment within a certain temperature range. The constant temperature treatment allows the temperature to fluctuate within the accuracy range of instrument control.

[0060] In the description of the invention, “plurality” means at least two, such as two, three, etc., unless otherwise clearly and specifically defined.

[0061] If there is no special description, all embodiments and optional embodiments of the present application can be combined with each other to form a new technical solution. If there is no special description, all technical features and optional technical features of the present application can be combined with each other to form a new technical solution.

[0062] If there is no special explanation, all the steps of the present application can be performed sequentially or randomly, preferably sequentially.

[0063] At present, the conductive materials used in implantable continuous glucose monitoring systems (CGMS) are mostly carbon materials or precious metal materials, and the electrode preparation materials are relatively limited. In addition, although the substrate materials such as PET and PI used in CGMS have good flexibility, carbon materials and metal materials themselves are not flexible. During the electrode deformation process, carbon materials and metal materials are difficult to adapt to the substrate deformation and crack or even fall off.

[0064] In addition, diabetes is a long-term chronic disease with no cure. How to effectively control the occurrence and development of diabetes has become a research problem that has attracted worldwide attention. Diabetes is often accompanied by a variety of acute complications, such as ketoacidosis, lactic acidosis, hyperglycemic hyperosmolar state, hypoglycemia, etc. During these complications, physiological indicators such as glucose, ketone bodies (such as β-hydroxybutyrate), lactate, and pH in the patient's blood / tissue fluid may change significantly, leading to a significant increase in the mortality rate of diabetes. Diabetic patients often monitor their blood sugar levels through blood glucose meters. However, a single blood sugar indicator is difficult to reflect the patient's true condition. Combining physiological and pathological information such as blood ketones and lactate will help to more accurately determine the condition and adjust the treatment plan in a timely manner.

[0065] In the related technology, test strips based on electrochemical sensing have been widely used to analyze the levels of biomarkers such as glucose and ketone bodies in the blood, but such test strip detection methods require frequent fingertip blood sampling and cannot be used to monitor the dynamic levels of biological molecules for a long time. The continuous glucose monitoring system (CGMS) that has emerged in recent years has achieved continuous monitoring of the dynamic level of blood glucose for up to 14 days by implanting micro-electrochemical sensors into subcutaneous tissue fluid, providing thousands of diabetic patients with a tool for all-round blood sugar regulation. However, so far, there have been no reports of implantable medical sensors that can continuously and simultaneously monitor the dynamic levels of multiple diabetes markers such as human blood glucose, blood ketones, lactate, pH, etc.

[0066] Based on the above problems, the present application provides an electrochemical sensor, in which the conductive layer of the working electrode of the electrochemical sensor comprises a conductive polymer PEDOT:PSS and a plasticizer; the conductive polymer PEDOT:PSS is modified by using a plasticizer, so that the conductivity and electrochemical activity of the conductive polymer PEDOT:PSS can meet the application requirements of the electrochemical sensor, breaking through the limitation that the traditional CGMS is only based on carbon materials and precious metal materials; at the same time, the PEDOT electrode has good flexibility, and still has good conductivity and electrochemical activity after multiple cycles of bending.

[0067] The first aspect of the present application provides an electrochemical sensor, which includes a flexible substrate and a working electrode arranged on at least one side of the flexible substrate; the working electrode includes a conductive layer, the conductive layer contains a conductive polymer PEDOT:PSS and a plasticizer; wherein the conductive polymer PEDOT:PSS is distributed in the conductive layer in the form of linear fibers.

[0068] In this application, PEDOT refers to poly(3,4-ethylenedioxythiophene) and PSS refers to poly(sodium styrene sulfonate).

[0069] It should be noted that the flexible substrate may be planar or three-dimensional; when the flexible substrate is planar, the working electrode may be disposed on only one surface of the flexible substrate, or on both opposite surfaces of the flexible substrate. When the flexible substrate is three-dimensional, the working electrode may be disposed on only one surface of the flexible substrate, or on both or more surfaces of the flexible substrate.

[0070] When the working electrode is disposed on only one side of the flexible substrate, it can be used to detect only one biological indicator; when the working electrodes are disposed on both sides or more than two sides of the flexible substrate, two or more biological indicators can be detected simultaneously.

[0071] By modifying the conductive polymer PEDOT:PSS with a plasticizer, its configuration can be changed from the intrinsic amorphous mass state to an ordered one-dimensional fiber state; when the electrode is deformed, the conductive polymer PEDOT:PSS will conform to the deformation of the substrate and slide relatively, so that the electrode's electroactive area remains unchanged and material fracture problems will not occur.

[0072] By modifying the conductive polymer PEDOT:PSS with a plasticizer, the electrostatic interaction between PEDOT and PSS can be effectively weakened, resulting in the enrichment of conductive PEDOT, accelerating the electron transfer rate, and significantly improving the conductivity and electrochemical activity of PEDOT:PSS.

[0073] Understandably, the present application can effectively improve the electrochemical activity and conductivity of the conductive polymer PEDOT:PSS by modifying the conductive polymer PEDOT:PSS with a plasticizer, so that the conductive polymer PEDOT:PSS can meet the application requirements of electrochemical sensors, breaking through the limitations of traditional CGMS based only on carbon materials and precious metal materials. At the same time, the working electrode using the conductive polymer PEDOT:PSS has good flexibility, and still has good conductivity and electrochemical activity after multiple cycles of bending, showing significant advantages over carbon-based electrodes.

[0074] In addition, the electrochemical sensor of the present application has more excellent electrochemical sensing performance, with an upper limit of glucose detection of up to 50mM and a detection sensitivity of about 10 times that of carbon-based sensors.

[0075] In some embodiments, the mass of the plasticizer accounts for a percentage of ≥0.1% of the mass of the conductive polymer PEDOT:PSS; for example, it may be, but is not limited to, 0.1%-50%, 0.1%-40%, 1%-30%, 3%-25%, 5%-20%, 10%-15%, 0.1%-10% or 1%-10%, etc. When the amount of the plasticizer is within the above range, the flexibility of the conductive polymer PEDOT:PSS can be improved while taking into account its conductivity. When the amount of the plasticizer is higher than the above range, the conductivity of the conductive polymer PEDOT:PSS will be affected; when the amount of the plasticizer is lower than the above range, the flexibility of the conductive polymer PEDOT:PSS will not be effectively improved.

[0076] As a possible embodiment, the plasticizer includes an ionic liquid. Optionally, the plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid and 1-ethyl-3-methylimidazolium tetracyanoboric acid. When the above ionic liquid is used as a plasticizer for modifying the conductive polymer PEDOT:PSS, the anions and cations of the ionic liquid will be embedded in the PEDOT chain and the PSS chain, breaking the electrostatic interaction between PEDOT and PSS, so that the conductive PEDOT can be enriched into a fibrous state, thereby significantly improving the conductivity, electrochemical activity and flexibility of PEDOT.

[0077] In some embodiments, the plasticizer includes a polyol. Optionally, the plasticizer includes one or more of xylitol, glycerol, and sorbitol. When the polyol is used as a plasticizer to modify the conductive polymer PEDOT:PSS, the polyol can hydrogen bond with PSS, thereby weakening the electrostatic interaction between PEDOT and PSS, so that the conductive PEDOT is enriched into a fibrous state, thereby significantly improving the conductivity, electrochemical activity and flexibility of PEDOT.

[0078] In some embodiments, the plasticizer includes an acid. Optionally, the plasticizer includes one or more of sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid and malic acid. When the above acid is used as a plasticizer to modify the conductive polymer PEDOT:PSS, the acid can protonate PSS, converting PSS from negative charge to neutral charge, thereby weakening the electrostatic interaction between PEDOT and PSS, so that the conductive PEDOT is enriched into a fibrous state, thereby significantly improving the conductivity, electrochemical activity and flexibility of PEDOT.

[0079] As a possible implementation, the plasticizer includes a surfactant. Optionally, the plasticizer includes one or more of a triaton and a fluorocarbon surfactant. When the above surfactant is used as a plasticizer to modify the conductive polymer PEDOT:PSS, the surfactant can cause phase separation of PEDOT and PSS, thereby weakening the electrostatic interaction between PEDOT and PSS, so that the conductive PEDOT can be enriched into a fibrous state, thereby significantly improving the conductivity, electrochemical activity and flexibility of PEDOT. .

[0080] In some optional embodiments, the plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid, 1-ethyl-3-methylimidazolium tetracyanoboric acid, xylitol, glycerol, sorbitol, sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid, malic acid, Triton and fluorocarbon surfactants.

[0081] As an example, the fluorocarbon surfactant mentioned in the present application may be Zonyl.

[0082] In some embodiments, the conductive layer further comprises a conductive material.

[0083] In some optional embodiments, the conductive material includes one or more of noble metal nanomaterials, carbon nanomaterials, black phosphorus, Mxene materials, metal organic framework materials and covalent organic framework materials. By adding the above materials, the electroactive area of ​​the sensor can be effectively increased and the electron transfer efficiency can be improved, thereby improving the electrochemical detection performance of the sensor.

[0084] It should be noted that Mxene material is a two-dimensional transition metal carbon (nitride).

[0085] In some optional embodiments, the noble metal nanomaterial includes one or more of gold, palladium, platinum, silver and rhodium.

[0086] In some optional embodiments, the carbon nanomaterial includes one or more of carbon nanotubes, graphene, graphyne and fullerene.

[0087] In some embodiments, the material of the flexible substrate includes one or more of polyethylene terephthalate, polyimide, polycarbonate, polyvinyl chloride, and polydimethylsiloxane.

[0088] As a possible implementation, the electrochemical sensor further includes a sensing layer located on the surface of the working electrode, wherein the sensing layer includes an electron mediator cross-linked by a first cross-linking agent and an enzyme or a pH sensitive material for oxidizing a monitoring object.

[0089] In some optional embodiments, the electron mediator includes one or more of organic conjugated small molecules and polymers thereof and transition metal complexes and polymers thereof with free amino groups or free carboxyl groups.

[0090] It should be noted that the “organic conjugated small molecules and their polymers” mentioned above refer to organic conjugated small molecules and their polymers; and “transition metal complexes with free amino groups or free carboxyl groups and their polymers” refer to transition metal complexes with free amino groups or free carboxyl groups and their polymers.

[0091] Optionally, the transition metal includes one or more of osmium, ruthenium, iron, copper, cobalt, vanadium and manganese.

[0092] Optionally, the organic conjugated small molecule includes one or more of catechol and its derivatives, thiophene and its derivatives, methylene blue and its derivatives, and methylene green and its derivatives.

[0093] As a possible implementation, the monitoring object includes one of glucose, lactic acid, ketone bodies, glutamic acid, glycine, uric acid, ascorbic acid, acetylcholine and cholesterol.

[0094] In some optional embodiments, the enzyme used to monitor the oxidation of the object includes one or more of glucose oxidase, glucose dehydrogenase, lactate oxidase, lactate dehydrogenase, β-hydroxybutyrate dehydrogenase, glutamate oxidase, glycine oxidase, urate oxidase, ascorbate oxidase, acetylcholinesterase and cholesterol oxidase.

[0095] In some embodiments, the first cross-linking agent includes one or more of dialdehyde compounds, dialdehyde polymers, epoxy compounds, epoxy polymers, N,N-methylenebisacrylamide, 1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride / N-hydroxysuccinimide, genipin, and dicyclohexylcarbodiimide.

[0096] It should be noted that "1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride / N-hydroxysuccinimide" means a mixture of 1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride and N-hydroxysuccinimide. In some optional embodiments, the molar ratio of 1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride and N-hydroxysuccinimide in the mixture is 1:1.

[0097] In some embodiments, the mass ratio of the electron mediator to the enzyme for oxidative monitoring is (0.5-2):1; for example, it can be, but not limited to, 0.5:1, 0.6:1, 0.7:1, 0.8:1, 0.9:1, 1:1, 1.1:1, 1.2:1, 1.3:1, 1.4:1, 1.5:1, 1.6:1, 1.7:1, 1.8:1, 1.9:1, 2:1 or a range between the above two ratios. The mass ratio of the first cross-linking agent to the enzyme for oxidative monitoring is (0.02-0.1):1; for example, it can be, but not limited to, 0.02:1, 0.03:1, 0.04:1, 0.05:1, 0.06:1, 0.07:1, 0.08:1, 0.09:1, 0.1:1 or a range between any two of the above ratios.

[0098] When the electron mediator, the enzyme for oxidizing the monitored object and the first cross-linking agent are used in combination according to the above mass ratio, the sensing layer can be stably modified and the sensor can present an obvious electrochemical response signal.

[0099] In some embodiments, the pH sensitive material includes one or more of polyaniline, transition metal oxide nanomaterials, and organic hydrogen ion carriers.

[0100] In some optional embodiments, the pH sensitive material includes one or more of polyaniline, transition metal oxide nanomaterials and organic hydrogen ion carriers, and the transition metal oxide nanomaterials include one or more of ZnO, IrO2, RuO2, WO3 and Ta2O5.

[0101] As a possible implementation manner, when a working electrode is disposed only on one surface of the flexible substrate, the monitored object is glucose.

[0102] In some optional embodiments, when a working electrode is disposed on only one surface of the flexible substrate, the monitored object is glucose, and the enzyme used to oxidize the monitored object includes one or more of glucose oxidase and glucose dehydrogenase.

[0103] In some embodiments, when working electrodes are respectively disposed on at least two side surfaces of the flexible substrate, and sensing layers are respectively disposed on the surfaces of the working electrodes, the multiple sensing layers are used to simultaneously monitor different monitoring objects.

[0104] As an example, when working electrodes are respectively arranged on the two side surfaces of the flexible substrate and sensing layers are respectively arranged on the surface of each working electrode, by respectively arranging different enzymes for oxidizing the monitored objects in the sensing layers on both sides, synchronous, real-time and quantitative monitoring of glucose and ketone bodies, glucose and lactic acid, ketone bodies and lactic acid, glucose and cholesterol, ketone bodies and uric acid, etc. can be achieved, which will help to more accurately diagnose the condition of diabetic patients and adjust the insulin treatment plan.

[0105] Alternatively, by setting a pH sensitive material in the sensing layer on one side and an enzyme for oxidizing the monitored object in the sensing layer on the other side, synchronous, real-time, quantitative monitoring of glucose and pH, ketone bodies and pH, lactic acid and pH, cholesterol and pH, uric acid and pH, etc. can be achieved, which will help to more accurately diagnose the condition of diabetic patients and adjust the insulin treatment plan.

[0106] In some embodiments, when working electrodes are respectively disposed on at least two side surfaces of the flexible substrate and sensing layers are respectively disposed on the surface of each working electrode, the monitoring object of the electrochemical sensor includes glucose, so as to achieve simultaneous detection of glucose and other biological indicators.

[0107] For example, when working electrodes are respectively arranged on the two side surfaces of the flexible substrate and a sensing layer is respectively arranged on the surface of each working electrode, synchronous, real-time and quantitative monitoring of glucose and ketone bodies, glucose and lactic acid, glucose and cholesterol, glucose and pH, etc. can be achieved.

[0108] For example, when working electrodes are respectively arranged on the three side surfaces of the flexible substrate and a sensing layer is respectively arranged on the surface of each working electrode, synchronous, real-time and quantitative monitoring of glucose, ketone bodies and lactic acid can be achieved.

[0109] In some optional embodiments, when working electrodes are respectively provided on at least two side surfaces of the flexible substrate and sensing layers are respectively provided on the surfaces of each working electrode, the monitoring object of the electrochemical sensor includes glucose, and the enzyme for oxidizing the monitoring object contained in the sensing layer on one side of the electrochemical sensor includes one or more of glucose oxidase and glucose dehydrogenase.

[0110] In some embodiments, the electrochemical sensor further comprises a diffusion limiting layer located on the surface of the sensing layer; the diffusion limiting layer comprises a substrate cross-linked by a cross-linking agent B, and the substrate comprises one or more of polyvinyl pyridine, polyurethane, polyvinyl imidazole, polyacrylate, polyether urethane and polystyrene. By providing the diffusion limiting layer, the diffusion flux of the analyte on the surface of the electrochemical sensor can be reduced; for example, when the electrochemical sensor of the present application is used to monitor glucose and ketone bodies, the upper limit of glucose detection can reach 30mM, and the upper limit of ketone body detection can reach 10mM.

[0111] In some optional embodiments, the mass ratio of the substrate to the second cross-linking agent is 1:(0.05-0.2); for example, it can be but not limited to 1:0.05, 1:0.06, 1:0.07, 1:0.08, 1:0.09, 1:0.1, 1:0.11, 1:0.12, 1:0.13, 1:0.14, 1:0.15, 1:0.16, 1:0.17, 1:0.18, 1:0.19, 1:0.2 or a range between any two of the above ratios. When the mass ratio of the substrate to the second cross-linking agent is within the above range, the diffusion limiting layer can be stably coated on the electrode surface and effectively regulate the diffusion flux of glucose on the electrode surface, thereby significantly improving the sensor's detection range, stability and other performances.

[0112] It should be noted that the "first cross-linking agent" and "second cross-linking agent" mentioned in this article are only used for descriptive purposes and cannot be understood as indicating or implying relative importance or quantity, nor can they be understood as implicitly indicating the importance or quantity of the indicated technical features.

[0113] In some embodiments, the electrochemical sensor further comprises a biocompatible layer located on the surface of the diffusion limiting layer. By providing the biocompatible layer, the biocompatibility and service life of the electrochemical sensor can be improved.

[0114] In some optional embodiments, the biocompatible layer comprises one or more of polyvinyl pyrrolidone, polyvinyl alcohol, chitosan, alginate, hyaluronic acid, cellulose, collagen, gelatin, polyacrylamide, polyacrylic acid, polypropylene alcohol, sodium polystyrene sulfonate, polyethylene glycol, polypropylene glycol, phosphorylcholine grafted polymer and betaine grafted polymer.

[0115] In some embodiments, the electrochemical sensor further includes a reference electrode that cooperates with the working electrode, and the configuration of the reference electrode conductive layer is the same as that of the working electrode conductive layer. Thus, the reference electrode also has good flexibility, and when the electrode is deformed, the performance of the reference electrode can remain stable, thereby further improving the stability of the electrochemical sensor.

[0116] It should be noted that "the setting of the reference electrode conductive layer is the same as that of the working electrode conductive layer" means that the conductive layer of the reference electrode also contains a conductive polymer PEDOT:PSS and a plasticizer, and the shape of the conductive polymer PEDOT:PSS, the type of plasticizer, and the percentage of the mass of the plasticizer to the mass of the conductive polymer PEDOT:PSS are the same as those of the conductive layer of the working electrode. When the conductive layer of the working electrode contains a conductive material, the conductive layer of the reference electrode also contains the type and amount of the conductive material.

[0117] Figure 1The schematic diagram of the preparation of an electrochemical sensor with a dual-electrode structure provided in one embodiment, wherein a group of working electrodes and a reference electrode are respectively provided on two opposite sides of a flexible substrate of the electrochemical sensor. Figure 1 As shown, the preparation method of the electrochemical sensor having a dual-electrode structure may include the following steps:

[0118] S1. A flexible polymer film is used as an electrode substrate, and the two opposite sides of the flexible substrate are named as side A and side B respectively; wherein the flexible polymer film may include but is limited to polyethylene terephthalate (PET), polyimide (PI), polycarbonate (PC), polyvinyl chloride (PVC), polydimethylsiloxane (PDMS), etc.

[0119] S2. Design as Figure 1 The stencil or mask pattern shown is used to deposit a conductive layer on the A side of the substrate by screen printing, physical vapor deposition or chemical vapor deposition in sequence, as the A side working electrode; deposit an insulating layer to expose the tip of the conductive layer (lower left corner) and the electrode contact (upper right corner), effectively controlling the electroactive area of ​​the working electrode and fully insulating it from the subsequent printed electrode; deposit a conductive layer for subsequent electron transport of the reference electrode; deposit an Ag / AgCl layer as the A side reference electrode; deposit an insulating layer to expose the tip working electrode and the reference electrode (lower left corner) and the electrode contact (upper right corner), effectively controlling the electroactive area of ​​the sensor. Among them, the materials used in depositing the conductive layer are conductive polymer PEDOT:PSS and plasticizer, and the conductive polymer PEDOT:PSS is in a one-dimensional fiber shape; the mass of the plasticizer accounts for 1%-5% of the mass of the conductive polymer PEDOT:PSS; the plasticizer includes one or more of bis(trifluoromethane)sulfonyl imide lithium, 1-butyl-3-methylimidazole tetrafluoroboric acid, 1-ethyl-3-methylimidazole tetracyanoboric acid, xylitol, glycerol, sorbitol, sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid, malic acid, Triton and DuPont Zonyl. In some optional embodiments, the materials used in depositing the conductive layer may also include conductive materials; the conductive materials include one or more of precious metal nanomaterials, carbon nanomaterials, black phosphorus, Mxene materials, metal organic framework materials and covalent organic framework materials.

[0120] S3. Turn the basement membrane over and follow the Figure 1 The scheme deposits electrode materials on the B surface of the substrate to prepare a micro dual electrode.

[0121] It should be noted that after each electrode layer is deposited, it is necessary to anneal it at 120°C for 30 minutes to make the deposited layer adhere closely to the substrate. The stencil / mask pattern can be designed according to actual needs to achieve controllable adjustment of the electrode size and number. For example, when the electrode shape is designed to be sickle-shaped and the electrode tip diameter is controlled within 300μm, it is convenient to adapt to electronic equipment (not shown) and the in vivo implantation process of the sensor. Hundreds of micro-electrodes can be arranged and deposited on the same substrate film at the same time to achieve mass production of electrodes ( Figure 1 To illustrate the electrode deposition process, only one electrode is shown).

[0122] It should be noted that Figure 1 The preparation process of the working electrode and the reference electrode is only given as an example. In practical applications, a counter electrode can be provided as required, and the counter electrode conductive layer can also adopt the same configuration as the working electrode conductive layer.

[0123] Figure 2 FIG. 1 is a schematic diagram of a structure of an electrochemical sensor having a dual-electrode structure according to an embodiment of the present invention, which can be based on but is not limited to Figure 1 The preparation process shown in the figure is prepared. Figure 2 As shown, the electrochemical sensor includes a working electrode 10, a reference electrode 20, an electrode insulating layer 30, a reference electrode contact 40 and a working electrode contact 50. The working electrode 10 and the reference electrode 20 are subcutaneously implanted parts, and electrochemical reactions occur with biomolecules in body fluids; the reference electrode contact 40 and the working electrode contact 50 are connected to an electronic device (not shown) on the surface of the skin to apply voltage to the working electrode 10 and the reference electrode 20 and output electrical response signals.

[0124] It should be noted that, in some embodiments, the implanted portion may also include a counter electrode (not shown).

[0125] Understandably, if Figure 1 As shown, by depositing two sets of working electrodes and reference electrodes on the surfaces of both sides of the flexible substrate respectively, the controllable integration and patterning of the two sets of micro-electrochemical sensors are achieved. The sensor size can be controllably adjusted according to needs, and it is expected to be used for subcutaneous implantation in living bodies.

[0126] It is not difficult to understand. Figure 1 The figure only shows how to prepare the dual electrodes on both sides of the flexible substrate by way of example; it should not be construed as limiting the dual electrodes to be prepared only on both sides of the flexible substrate. When the flexible substrate is a three-dimensional structure with multiple sides, electrodes can be respectively arranged on three or more sides of the flexible substrate as required.

[0127] In order to monitor and analyze different biological indicators, after the dual electrodes are prepared, specific sensing layers can be further modified on the working electrodes on the A side and the B side. As an example, taking the simultaneous monitoring of glucose and ketone bodies as an example, the following operations can be performed on the working electrodes on the A side and the B side respectively:

[0128] Glucose was detected by the A-side electrode, and the glucose sensor was prepared by the second-generation enzyme-type sensing principle. Glucose oxidase (or glucose dehydrogenase), electron mediators (such as transition metal complexes with free amino groups or free carboxyl groups and their polymers, organic conjugated small molecules and their polymers, etc.) and crosslinkers (such as dialdehyde compounds, dialdehyde polymers, epoxy compounds, epoxy polymers, N,N-methylenebisacrylamide, genipin, etc.) were combined by chemical crosslinking to prepare a sensing layer solution, and the sensing layer solution was modified at the tip of the A-side working electrode 10, and naturally dried at room temperature to obtain a sensing layer.

[0129] The B-side electrode is used as a sensor for β-hydroxybutyrate (the ketone body with the highest content in the human body), and β-hydroxybutyrate dehydrogenase and coenzyme NAD + The electron mediator and the cross-linking agent are combined by chemical cross-linking to prepare a sensing layer solution, the sensing layer solution is modified on the tip of the working electrode on the B side, and the solution is naturally dried at room temperature.

[0130] Since biological enzymes have Michaelis-Menten kinetics, the linear range of detection of the above sensors is limited, and it is difficult to meet the detection needs of glucose and ketone bodies in the human body. In some examples, in order to expand the detection range of the analyte, a diffusion limiting layer is coated on the surface of the double electrode after the above-mentioned preparation of the sensing layer to reduce the diffusion flux of the analyte on the sensor surface, and it is expected that the upper limit of glucose detection will reach 30mM and the upper limit of ketone body detection will reach 10mM. Among them, the substrate for preparing the diffusion limiting layer includes one or more of polyvinylpyridine, polystyrene, polyurethane, polyethoxyethyl acrylate, polyhydroxyethyl methacrylate and polyhydroxyethyl acrylate.

[0131] In some examples, a biocompatible layer may be further coated outside the diffusion limiting layer to improve the biocompatibility and service life of the sensor. The biocompatible layer may include one or more of polyvinyl pyrrolidone, polyvinyl alcohol, chitosan, alginate, hyaluronic acid, cellulose, collagen, gelatin, polyacrylamide, polyacrylic acid, polypropylene alcohol, sodium polystyrene sulfonate, polyethylene glycol, polypropylene glycol, phosphorylcholine grafted polymer and betaine grafted polymer.

[0132] The cross-sectional diagram of the working electrode end obtained after the above process is as follows: Figure 3 shown by Figure 3 It can be seen that a sensing layer 60 is disposed on the surface of the working electrode 10 , a diffusion limiting layer 70 is disposed on the surface of the sensing layer 60 , and a biocompatibility 80 is disposed on the surface of the diffusion limiting layer 70 .

[0133] Figure 4 Schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and ketone bodies; Figure 4 It can be seen that when the electrochemical sensor monitors glucose and ketone bodies simultaneously, glucose is oxidized to gluconolactone by the glucose oxidase modified on the A surface, and transfers electrons to the electron mediator, which is further electrochemically oxidized on the electron surface to generate a current signal; at the same time, β-hydroxybutyrate is oxidized to acetoacetate by the β-hydroxybutyrate dehydrogenase modified on the B surface, and transfers electrons to the coenzyme NAD in turn. + The electron mediator is further electrochemically oxidized on the electron surface to generate a current signal. By simultaneously detecting the current signals of the two electrodes, the simultaneous monitoring of glucose and ketone bodies can be achieved.

[0134] It should be noted that by changing the type of enzyme contained in the sensing layer, simultaneous real-time monitoring of other biological indicators can also be achieved; for example, simultaneous real-time monitoring of any two biological indicators including but not limited to glucose, lactate, ketone bodies, glutamate, glycine, uric acid, ascorbic acid, acetylcholine and cholesterol can be achieved.

[0135] As an example, in order to achieve simultaneous real-time monitoring of glucose and lactate, a composite sensing layer of glucose oxidase (or glucose dehydrogenase), an electron mediator and a cross-linking agent is modified on the A side to achieve electrochemical detection of glucose; and a composite sensing layer of lactate oxidase (or lactate dehydrogenase), an electron enzyme mediator and a cross-linking agent is modified on the B side to achieve electrochemical detection of lactate.

[0136] Figure 5 The schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and lactate; Figure 5 As shown, when the electrochemical sensor monitors glucose and lactate simultaneously, glucose is oxidized to gluconolactone by the glucose oxidase modified on the A surface, and electrons are transferred to the electron mediator, which is further electrochemically oxidized on the electron surface to generate a current signal; at the same time, lactate is oxidized to pyruvate by the lactate oxidase modified on the B surface, and electrons are transferred to the electron mediator, which is further electrochemically oxidized on the electron surface to generate a current signal. By simultaneously detecting the current signals of the two electrodes, the simultaneous monitoring of glucose and lactate can be achieved.

[0137] As an example, in order to achieve simultaneous real-time monitoring of glucose and pH, a composite sensing layer of glucose oxidase (or glucose dehydrogenase), an electron mediator and a cross-linking agent is modified on the A surface to achieve electrochemical detection of glucose; a pH-sensitive coating material is modified on the B surface to achieve real-time detection of the pH value by the open circuit potential method; wherein the pH-sensitive coating material may include but is not limited to one or more of polyaniline, transition metal oxide nanomaterials, and organic hydrogen ion carriers; transition metal oxide nanomaterials may include but are not limited to one or more of ZnO, IrO2, RuO2, WO3, Ta2O5, etc.

[0138] Figure 6 Schematic diagram of the principle of electrochemical sensor for simultaneous monitoring of glucose and pH; Figure 6 As shown in the figure, when the electrochemical sensor monitors glucose and pH simultaneously, glucose is oxidized to gluconolactone by the glucose oxidase modified on the A surface, and the electrons are transferred to the electron mediator, which is further electrochemically oxidized on the electron surface to generate a current signal; at the same time, the pH sensitive coating modified on the B surface can adsorb or desorb hydrogen ions, causing the ion flux on the sensor surface to change, thereby causing a change in the electrode potential. By simultaneously detecting the current signal or open circuit voltage signal of the two electrodes, the simultaneous monitoring of glucose and pH can be achieved.

[0139] It can be understood that by modifying the surfaces of two or more groups of sensors with different types of specific functionalized sensing layers, the simultaneous, real-time, and quantitative monitoring of two or more diabetes biomarkers can be achieved, which will help to more accurately diagnose the condition of diabetic patients and adjust insulin treatment plans.

[0140] In some embodiments, the electrochemical sensor further includes a counter electrode that cooperates with the working electrode, and the configuration of the counter electrode conductive layer is the same as that of the working electrode conductive layer. Thus, the counter electrode also has good flexibility, and when the electrode is deformed, the performance of the counter electrode can remain stable, thereby further improving the stability of the electrochemical sensor.

[0141] It should be noted that "the configuration of the counter electrode conductive layer is the same as that of the working electrode conductive layer" means that the conductive layer of the counter electrode also contains a conductive polymer PEDOT:PSS and a plasticizer, and the shape of the conductive polymer PEDOT:PSS, the type of plasticizer, and the percentage of the mass of the plasticizer to the mass of the conductive polymer PEDOT:PSS are the same as those of the conductive layer of the working electrode. When the working electrode conductive layer contains a conductive material, the counter electrode conductive layer also contains the type and amount of the conductive material.

[0142] The second aspect of the present application provides a biological indicator detection device, which includes the electrochemical sensor of the first aspect of the present application.

[0143] The biological indicator detection device of the present application includes the electrochemical sensor provided by the present application, and thus has at least the same advantages as the electrochemical sensor.

[0144] In some embodiments, the biological marker monitoring device includes a blood glucose monitoring device.

[0145] The technical scheme of the present application is described in detail below in conjunction with specific embodiments. It should be understood that these embodiments are only used to illustrate the present application and are not used to limit the scope of the present application. The experimental methods for which specific conditions are not specified in the following embodiments are preferably referred to the guidance given in the present application, and can also be carried out according to the experimental manual or conventional conditions in the field, or according to the conditions recommended by the manufacturer, or with reference to experimental methods known in the art.

[0146] In the following specific embodiments, the measured parameters of raw material components may have slight deviations within the range of weighing accuracy unless otherwise specified. For temperature and time parameters, acceptable deviations caused by instrument test accuracy or operation accuracy are allowed.

[0147] Example 1. Preparation of working electrode (PEDOT electrode)

[0148] 1. Pipette a 1.3 wt% aqueous solution of the conductive polymer PEDOT:PSS into a clean sample bottle, add 2 wt% of lithium bis(trifluoromethane)sulfonyl imide (plasticizer) based on the dry weight of the conductive polymer PEDOT:PSS, and stir at room temperature for 1 h to prepare PEDOT conductive ink.

[0149] 2. PEDOT conductive ink was dropped onto the PET film (flexible substrate) and spin-coated at 1500 rpm for 30 s to make the conductive ink evenly and continuously dispersed.

[0150] 3. The PET film coated with the PEDOT conductive ink was heated at 100° C. for 30 min to solidify the PEDOT into a film, thereby obtaining a working electrode (ie, a PEDOT electrode).

[0151] Comparative Example 1. Preparation of working electrode (carbon electrode)

[0152] The difference between Comparative Example 1 and Example 1 is that the same dry weight of carbon is used to replace the PEDOT aqueous solution, and no plasticizer is added, and the other conditions are the same.

[0153] The sheet resistance and electrochemical performance of the working electrodes prepared in Example 1 and Comparative Example 1 were tested respectively. The area of ​​the working electrodes was controlled to be 0.5×1 cm 2 , perform sheet resistance test and cyclic voltammetry test as follows:

[0154] Sheet resistance value test: Place the square resistance meter probe on the PEDOT electrode and the carbon electrode respectively to perform sheet resistance test.

[0155] Cyclic voltammetry test: The PEDOT electrode and the carbon electrode were placed in a 10 mM potassium ferrocyanide solution for cyclic voltammetry test.

[0156] The sheet resistance values ​​of Example 1 and Comparative Example 1 were tested as follows: Figure 7 shown.

[0157] From the results of Example 1 and Comparative Example 1, it can be seen that the sheet resistance of the carbon electrode increases significantly after being bent, while the PEDOT electrode has no significant change after being bent, indicating that the working electrode prepared using the conductive polymer PEDOT:PSS and the plasticizer has good flexibility.

[0158] The cyclic voltammetric response results of the working electrode before and after bending in Example 1 are as follows: Figure 8 As shown; the cyclic voltammetric response results of the working electrode before and after bending in Comparative Example 1 are as follows Fig. 9 Shown

[0159] Depend on Figure 8 It can be seen that when not bent, the PEDOT electrode in Example 1 presents an obvious pair of redox peaks in potassium ferrocyanide solution, wherein the oxidation peak potential is +0.3 V, the reduction peak potential is +0.22 V, and the peak potential difference is only 0.08 V, indicating that the PEDOT electrode interface has a fast electron transfer rate and has excellent electrochemical properties.

[0160] After the PEDOT electrode in Example 1 was bent 180° for 20 times, the current value and peak potential difference of the PEDOT electrode remained almost unchanged, indicating that the PEDOT electrode has good flexibility and electrochemical stability.

[0161] Depend on Fig. 9 It can be seen that when not bent, the carbon electrode in Comparative Example 1 only shows current polarization in potassium ferrocyanide solution, without obvious redox peaks, and the current size is only 1 / 10 of that of the PEDOT electrode, which indicates that the electrochemical performance of the carbon electrode is poor. After the carbon electrode in Comparative Example 1 was bent 180° for 20 times, the oxidation current and reduction current of the carbon electrode both dropped significantly, indicating that the carbon electrode has poor flexibility and electrochemical stability.

[0162] Analysis shows that the reason for the above phenomenon may be that the addition of plasticizer changes the configuration of the conductive polymer PEDOT:PSS from the intrinsic amorphous mass to an ordered one-dimensional fiber. When the electrode is deformed, the PEDOT fiber will slide relative to the deformation of the substrate, so that the electrode's electroactive area remains unchanged and the material will not break. The carbon electrode is made of carbon microparticles. When the electrode is deformed, the carbon particles will separate and crack from each other in accordance with the deformation of the substrate, which blocks the electron transfer path and seriously affects the electrical and electrochemical properties of the electrode.

[0163] Example 2. Glucose sensor preparation

[0164] 1. Using PBS as solvent, prepare a mixed solution of 10 mg / mL electron mediator, 10 mg / mL glucose enzyme, and 0.5 mg / mL cross-linking agent, and stir at room temperature for 4 hours to fully cross-link the electron mediator and glucose enzyme; wherein the electron mediator is Os(bpy)2Im-bpy(CH2)6-NH2, the glucose enzyme is glucose oxidase, and the cross-linking agent is polyethylene glycol diglycidyl ether (PEGDGE).

[0165] 2. 3 μL of the cross-linked sensing layer solution was taken and dropped onto the surfaces of the working electrodes in Example 1 and Comparative Example 1, respectively, and dried and solidified at room temperature to form a film, thereby obtaining a glucose sensor.

[0166] The two glucose sensors prepared above were characterized by placing the sensor in a PBS solution and using the chronoamperometry to measure the response current values ​​of the sensor to 0.1mM, 0.2mM, 0.3mM, 0.5mM, 1mM, 2mM, 3mM, 5mM, and 10mM glucose solutions.

[0167] The chronoamperometric response curve of the glucose sensor using the PEDOT electrode in Example 1 (hereinafter referred to as the PEDOT-based glucose sensor) is as follows: Fig.10 The glucose standard curve is shown in Fig.11 The chronoamperometric response curve of the glucose sensor using the carbon electrode in Comparative Example 1 (hereinafter referred to as the carbon-based glucose sensor) is shown in FIG. Fig.12 The glucose standard curve is shown in Fig.13 shown.

[0168] Depend on Fig.10 and Fig.11 It can be seen that with the increase of glucose concentration, the response current value of the PEDOT-based glucose sensor gradually increases, showing a good linear response (R 2 =0.9959), the standard curve equation is y(μA)=2.02x(mM)-0.15.

[0169] Depend on Fig.12 and Fig.13 It can be seen that the linear equation of the carbon-based glucose sensor is y(μA)=0.29x(mM)+0.0057, R 2 =0.9986.

[0170] From the above results, it can be seen that the detection sensitivity of PEDOT-based glucose sensor to glucose is 7 times that of carbon-based glucose sensor, indicating that the sensor using PEDOT electrode has excellent glucose sensing performance.

[0171] contrast Fig.10 and Fig.12 From the chronoamperometric curves, it can be found that the current response of the PEDOT-based glucose sensor is very stable, while the current response of the carbon-based glucose sensor has some fluctuations. This result once again proves the good electrochemical stability of the PEDOT electrode.

[0172] Example 3. Preparation of a glucose sensor coated with a diffusion limiting layer

[0173] Using ethanol as solvent, a mixed solution of 100 mg / mL poly(4-vinylpyridine) (PVP) and 5 mg / mL polyethylene glycol diglycidyl ether (PEGDGE) was prepared and stirred at room temperature for 4 hours to fully crosslink the PVP. The diffusion limiting layer solution was coated on the surface of the PEDOT-based glucose sensor and the carbon-based glucose sensor using the dip-pull technique, and cured into a film under vacuum at 40°C.

[0174] The two glucose sensors coated with diffusion limiting layers prepared above were characterized. The specific operation method was as follows: the sensor was placed in a PBS solution, and the chronoamperometry was used to measure the response current values ​​of the sensor to 1mM, 2mM, 3mM, 5mM, 7mM, 10mM, 15mM, 20mM, 25mM, 30mM, 35mM, 40mM, 45mM, and 50mM glucose solutions.

[0175] The chronoamperometric response curve of the PEDOT-based glucose sensor coated with the diffusion limiting layer in Example 1 is as follows: Fig.14 The glucose standard curve is shown in Fig.15 The chronoamperometric response curve of the carbon-based glucose sensor coated with the diffusion limiting layer in Comparative Example 1 is shown in Fig.16 The glucose standard curve is shown in Fig.17 shown.

[0176] Depend on Fig.14It can be seen that after the PVP diffusion restriction layer is coated on the surface of the PEDOT-based glucose sensor, the upper limit of the sensor's detection of glucose is increased to 50mM, and the response current is reduced to the nA level. This is because the PVP diffusion restriction layer has a porous structure, which can effectively limit the diffusion flux and diffusion rate of glucose on the electrode surface. The standard curve of the sensor at this time is calculated to be y(μA)=0.013x(mM)+0.026, R 2 =0.9958.

[0177] Depend on Fig.16 After coating the PVP diffusion restriction layer on the surface of the carbon-based glucose sensor, the upper limit of glucose detection can be increased to 50mM, but the detection sensitivity is only 1 / 10 of that of the PEDOT-based glucose sensor, and there is still the phenomenon of unstable fluctuation of the response current, which fully reflects the advantages of the PEDOT-based sensor of this application over the traditional carbon-based CGM sensor.

[0178] The technical features of the above-described embodiments may be arbitrarily combined. To make the description concise, not all possible combinations of the technical features in the above-described embodiments are described. However, as long as there is no contradiction in the combination of these technical features, they should be considered to be within the scope of this specification.

[0179] The above-mentioned embodiments only express several implementation methods of the present application, and the descriptions thereof are relatively specific and detailed, but they cannot be understood as limiting the scope of the invention patent. It should be pointed out that, for a person of ordinary skill in the art, several variations and improvements can be made without departing from the concept of the present application, and these all belong to the protection scope of the present application. Therefore, the protection scope of the patent of the present application shall be subject to the attached claims.

Claims

1. An electrochemical sensor, characterized in that It comprises a flexible substrate and a working electrode arranged on at least one side of the flexible substrate; the working electrode comprises a conductive layer, the conductive layer comprises a conductive polymer PEDOT:PSS and a plasticizer; wherein the conductive polymer PEDOT:PSS is distributed in the conductive layer in the form of linear fibers.

2. The electrochemical sensor according to claim 1, characterized in that The mass percentage of the plasticizer to the mass percentage of the conductive polymer PEDOT:PSS is ≥0.1%.

3. The electrochemical sensor according to claim 2, characterized in that The plasticizer has one of the following characteristics (1)-(5): (1) The plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid and 1-ethyl-3-methylimidazolium tetracyanoboric acid; (2) The plasticizer includes one or more of xylitol, glycerol and sorbitol; (3) The plasticizer includes one or more of sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid and malic acid; (4) The plasticizer includes one or more of Triton and fluorocarbon surfactant; (5) The plasticizer includes one or more of lithium bis(trifluoromethane)sulfonyl imide, 1-butyl-3-methylimidazolium tetrafluoroboric acid, 1-ethyl-3-methylimidazolium tetracyanoboric acid, xylitol, glycerol, sorbitol, sulfuric acid, nitric acid, phosphoric acid, oxalic acid, methanesulfonic acid, malic acid, Triton and fluorocarbon surfactant.

4. The electrochemical sensor according to claim 1, characterized in that The conductive layer further comprises a conductive material; The conductive material includes one or more of noble metal nanomaterials, carbon nanomaterials, black phosphorus, Mxene materials, metal organic framework materials and covalent organic framework materials.

5. The electrochemical sensor according to claim 1, characterized in that The material of the flexible substrate includes one or more of polyethylene terephthalate, polyimide, polycarbonate, polyvinyl chloride and polydimethylsiloxane.

6. The electrochemical sensor according to any one of claims 1 to 5, characterized in that The invention also comprises a sensing layer located on the surface of the working electrode, wherein the sensing layer comprises an electron mediator cross-linked by a first cross-linking agent and an enzyme or a pH sensitive material for oxidizing a monitoring object.

7. The electrochemical sensor according to claim 6, characterized in that The sensing layer has at least one of the following features (6)-(10): (6) The electron mediator includes one or more of organic conjugated small molecules and polymers thereof and transition metal complexes and polymers thereof with free amino groups or free carboxyl groups; The transition metal includes one or more of osmium, ruthenium, iron, copper, cobalt, vanadium and manganese; The organic conjugated small molecules include one or more of catechol and its derivatives, thiophene and its derivatives, methylene blue and its derivatives, and methylene green and its derivatives; (7) The monitoring object includes one of glucose, lactic acid, ketone bodies, glutamic acid, glycine, uric acid, ascorbic acid, acetylcholine and cholesterol; The enzyme for oxidizing the monitored object includes one or more of glucose oxidase, glucose dehydrogenase, lactate oxidase, lactate dehydrogenase, β-hydroxybutyrate dehydrogenase, glutamate oxidase, glycine oxidase, urate oxidase, ascorbate oxidase, acetylcholinesterase and cholesterol oxidase; (8) the first cross-linking agent includes one or more of dialdehyde compounds, dialdehyde polymers, epoxy compounds, epoxy polymers, N,N-methylenebisacrylamide, 1-ethyl-(3-dimethylaminopropyl)carbodiimide hydrochloride / N-hydroxysuccinimide, genipin and dicyclohexylcarbodiimide; (9) The pH sensitive material comprises one or more of polyaniline, transition metal oxide nanomaterials and organic hydrogen ion carriers; The transition metal oxide nanomaterial includes one or more of ZnO, IrO2, RuO2, WO3 and Ta2O5.

8. The electrochemical sensor according to claim 7, characterized in that When a working electrode is provided only on one side surface of the flexible substrate, the monitoring object is glucose; The enzyme used for oxidizing the monitored object includes one or more of glucose oxidase and glucose dehydrogenase.

9. The electrochemical sensor according to claim 7, characterized in that When working electrodes are respectively arranged on at least two side surfaces of the flexible substrate, and sensing layers are respectively arranged on the surface of each of the working electrodes, the plurality of sensing layers are used to simultaneously monitor different monitoring objects.

10. The electrochemical sensor according to claim 9, characterized in that The monitoring object of the electrochemical sensor includes glucose; The enzyme for oxidizing the monitored object contained in the sensing layer on one side of the electrochemical sensor includes one or more of glucose oxidase and glucose dehydrogenase.

11. The electrochemical sensor according to claim 7, characterized in that It also includes a diffusion limiting layer located on the surface of the sensing layer; the diffusion limiting layer includes a substrate cross-linked by a second cross-linking agent, and the substrate includes one or more of polyvinyl pyridine, polyurethane, polyvinyl imidazole, polyacrylate, polyether urethane and polystyrene.

12. The electrochemical sensor according to claim 11, characterized in that Also included is a biocompatible layer located on the surface of the diffusion limiting layer; The biocompatible layer comprises one or more of polyvinyl pyrrolidone, polyvinyl alcohol, chitosan, alginate, hyaluronic acid, cellulose, collagen, gelatin, polyacrylamide, polyacrylic acid, polypropylene alcohol, sodium polystyrene sulfonate, polyethylene glycol, polypropylene glycol, phosphorylcholine grafted polymer and betaine grafted polymer.

13. The electrochemical sensor according to any one of claims 1 to 12, characterized in that The electrochemical sensor further comprises a reference electrode cooperating with the working electrode, and the configuration of the reference electrode conductive layer is the same as that of the working electrode conductive layer.

14. The electrochemical sensor according to any one of claims 1 to 12, characterized in that The electrochemical sensor further comprises a counter electrode cooperating with the working electrode, and the configuration of the counter electrode conductive layer is the same as that of the working electrode conductive layer.

15. A biological indicator monitoring device, characterized in that: Comprising the electrochemical sensor according to any one of claims 1 to 14.

16. The biological indicator monitoring device according to claim 15, characterized in that: The biological indicator detection device includes a blood sugar monitoring device.

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