Electrode for implantable biosensor and application thereof

Abstract

The invention provides an electrode for an implantable biosensor and application of the electrode, a conducting layer of the electrode is located on an insulating substrate of the biosensor, a sensing layer and a limiting layer are sequentially arranged on the conducting layer of the electrode, the sensing layer comprises enzyme required by reaction, and the sensing layer further comprises cationic polymer. The biosensor provided by the invention realizes good linear correlation of testing under low potential, obviously reduces interference of acetaminophen and other interfering substances on test signals under the condition of low-potential testing, and improves the accuracy of detection.

Description

Technical Field

[0001] This invention belongs to the field of sensor technology, specifically relating to the electrode structure of implantable biosensors and their applications. Background Technology

[0002] Blood glucose monitoring is an important component of diabetes management, and its results help assess the degree of glucose metabolism disorder in diabetic patients and develop appropriate blood glucose control strategies. There are various methods of blood glucose monitoring, such as blood glucose monitoring (BGM) at specific times, such as fasting or 2 hours after a meal, and continuous glucose monitoring (CGM) that tracks blood glucose levels around the clock.

[0003] Continuous glucose monitoring (CGM) systems consist of a needle applicator, sensor, transmitter, and receiver, providing continuous and comprehensive blood glucose information. The needle applicator is used to implant electrodes of an implantable biosensor subcutaneously in the patient's abdomen or arm. The system detects glucose concentration in the interstitial fluid based on the electrical signal generated by the electrochemical reaction catalyzed by glucose oxidase. Since the glucose concentration in the interstitial fluid is correlated with blood glucose concentration, an algorithm converts this into a blood glucose reading, which is then transmitted by the transmitter to the terminal device (receiver).

[0004] Implantable biosensors in CGM systems typically include a conductive layer on a flexible insulating substrate, forming a working electrode, a counter electrode, and / or a reference electrode. The conductive layer of the working electrode is often made of platinum. An enzyme layer and a glucose restriction membrane layer are sequentially arranged on the conductive layer of the working electrode, sometimes also including a separate biocompatible membrane layer. The enzyme layer of the glucose biosensor contains glucose oxidase (Gox), and the glucose restriction membrane layer is made of materials such as polyurethane. Oxygen molecules dissolved in the interstitial fluid and glucose molecules pass through the glucose restriction membrane layer into the enzyme layer. If a biocompatible membrane layer is present, glucose molecules pass through the biocompatible membrane layer and the glucose restriction membrane layer before entering the enzyme layer, where they are catalyzed by glucose oxidase to produce H₂O₂ and gluconolactone. The glucose oxidation reaction catalyzed by glucose oxidase can be divided into two steps: first, the coenzyme group FAD in Gox is reduced to its reduced form FADH₂ by glucose, and then oxidized to its oxidized form FAD by oxygen, which acts as a mediator (Medox), thus completing a catalytic cycle.

[0005] Currently, implantable biosensors offer advantages such as simple structure and easy miniaturization of electrodes. Existing implantable biosensors for glucose detection typically require relatively high potentials, such as +0.55V, to detect hydrogen peroxide, a product of the glucose reaction, as this facilitates the redox reaction of hydrogen peroxide. However, some electroactive substances in the sample, such as acetaminophen, ascorbic acid, and uric acid, are prone to redox reactions at high potentials, thus interfering with glucose concentration detection and leading to inaccurate results. Summary of the Invention

[0006] To address the interference problems existing in the prior art, the present invention aims to provide an electrode for an implantable biosensor, a biosensor having the electrode structure, and its application in continuous blood component (e.g., blood glucose) monitoring.

[0007] The present invention provides an electrode for an implantable biosensor. The conductive layer of the electrode is located on the insulating substrate of the biosensor. A sensing layer and a confinement layer are sequentially disposed on the conductive layer of the electrode. The sensing layer contains an enzyme required for the reaction and also contains a cationic polymer.

[0008] The sensing layer may include a two-layer structure of a cationic layer and an enzyme layer, with the cationic polymer located inside the cationic layer and the enzyme located inside the enzyme layer.

[0009] Alternatively, the sensing layer may comprise a hybrid layer composed of a cationic polymer and an enzyme.

[0010] Furthermore, an optimization layer may be provided above the restriction layer. The optimization layer includes an enzyme layer and a polyurethane layer, with the enzyme layer located on the restriction layer and the polyurethane layer located on the enzyme layer.

[0011] The cationic polymer is selected from one of polydiallyldimethylammonium chloride, polydiallyldimethylammonium chloride derivatives, polyethyleneimine, polyethyleneimine derivatives, poly(allylamine hydrochloride), poly(allylamine hydrochloride) derivatives, or a combination of two or more of them.

[0012] The molecular weight of the cationic polymer is selected from, but not limited to, 15,000-700,000.

[0013] The biosensor is used for glucose detection, and the enzyme on the electrode includes glucose oxidase, and the restriction layer includes polyurethane.

[0014] In different combinations, the concentrations of solutions such as GOx solution, cationic polymer solution, and polyurethane solution can be adjusted according to the actual design. For example, the concentration of GOx solution can be selected from, but is not limited to, 60-120 mg / mL, the concentration of cationic polymer solution can be selected from, but is not limited to, 0.5%-5%, and the concentration of polyurethane solution can be selected from, but is not limited to, 1%-7%. The soaking, drying, and crosslinking times can also be adjusted according to the reagents used.

[0015] The present invention provides a biosensor utilizing the electrode described herein, a flexible insulating substrate, an electrode located on the flexible insulating substrate, leads, and pins.

[0016] The electrodes on a biosensor include a working electrode and a counter electrode.

[0017] The biosensor may also include a reference electrode, a temperature electrode, etc.

[0018] The working electrode, counter electrode, reference electrode, and other electrodes can be disposed on the front and back sides of the biosensor, respectively. For example, the working electrode and counter electrode can be located on the same side of the flexible insulating substrate of the biosensor, or on the front and back sides, respectively.

[0019] The present invention provides an electrode for an implantable biosensor. The conductive layer of the electrode is located on an insulating substrate, and a confinement layer is provided on the conductive layer of the electrode. The electrode includes a first electrode and a second electrode. At least the first electrode has a sensing layer located between the conductive layer and the confinement layer. The sensing layer contains an enzyme required for the reaction and also contains a cationic polymer.

[0020] Furthermore, the sensing layer includes a cationic layer and an enzyme layer, with the cationic layer located above the conductive layer and containing a cationic polymer, and the enzyme layer containing the enzyme required for the reaction; or the sensing layer includes a mixed layer composed of a cationic polymer and an enzyme.

[0021] Furthermore, an optimization layer is provided above the electrode's confinement layer. The optimization layer includes an enzyme layer and a polyurethane layer, with the enzyme layer located on the confinement layer and the polyurethane layer located on the enzyme layer.

[0022] Furthermore, the cationic polymer is selected from one of polydiallyldimethylammonium chloride, polydiallyldimethylammonium chloride derivatives, polyethyleneimine, polyethyleneimine derivatives, poly(allylamine hydrochloride), poly(allylamine hydrochloride) derivatives, or a combination of two or more thereof.

[0023] Furthermore, the molecular weight of the cationic polymer is 15,000-700,000.

[0024] Furthermore, the biosensor is used for glucose detection, and the enzyme is glucose oxidase.

[0025] Furthermore, the first electrode is the working electrode. The second electrode is selected from either the counter electrode or the reference electrode.

[0026] The sensing layer, located on the conductive layer of the electrode, contains the enzyme required for the reaction, which reacts with the analyte in the biological sample. For example, when the electrode is used in a biosensor for glucose detection, the enzyme on the sensing layer is glucose oxidase, which reacts with glucose in the sample to generate an electrical signal.

[0027] A confinement layer on the electrode controls the amount or rate at which specific components in a biological sample reach the sensing layer. In implantable continuous glucose monitoring applications, the confinement layer ensures that glucose and oxygen in the sample pass through it in a specific ratio, maintaining the stability of the glucose oxidase reaction on the sensing layer. The confinement layer also acts as a biocompatible membrane, effectively preventing or reducing inflammation in the surrounding tissues after the electrode is inserted into the body. The material of the confinement layer can be selected from, but is not limited to, polyurethane, polyurethanepolyuria, poly(ethylene oxide), polydimethyl siloxane monomethacrylate, and a branched acrylate polymer.

[0028] Biosensors with electrodes described in this invention can be used in continuous detection systems, such as, but not limited to, the application of biosensors described in this invention in implantable continuous blood glucose monitoring.

[0029] Compared to existing technologies, this invention achieves the following advantages. By using a cationic polymer to prepare the outer membrane of the biosensor electrode, this invention achieves good linear correlation between the biosensor and biological samples such as glucose at low potentials. Furthermore, under low-potential testing conditions, it significantly reduces interference from substances such as acetaminophen, ascorbic acid, and uric acid on the glucose test signal, thereby improving detection accuracy. Attached Figure Description

[0030] Figure 1 This is a schematic diagram of a membrane structure on the platinum electrode layer of the biosensor of the present invention. The membrane layer includes a sensor layer, a confinement layer, and an optimization layer.

[0031] Figure 2 This is a schematic diagram of another membrane structure on the platinum electrode layer of the biosensor of the present invention, wherein the membrane layer includes a sensing layer and a confinement layer.

[0032] Figure 3 These are schematic diagrams of the front and back of an implantable glucose sensor.

[0033] Figure 4 This is a schematic diagram of the current response of the sensors to glucose after the completion of the four formulations (A, B, C, and D) in Example 2.

[0034] Figure 5 This is a schematic diagram showing the sensitivity of the sensors to glucose after the completion of the four formulations (A, B, C, and D) in Example 2.

[0035] Figure 6 A schematic diagram of the current response of a sensor made from a mixture of cationic polymer and enzyme to glucose.

[0036] Figure 7 A schematic diagram of the glucose sensor's response to glucose, fabricated using three cations (polydiallyldimethylammonium chloride, polyethyleneimine, and poly(allylamine hydrochloride)).

[0037] Figure 8 A schematic diagram illustrating the sensitivity of glucose sensors fabricated with different concentrations of polyethyleneimine.

[0038] Figure 9 Schematic diagram of the anti-interference performance evaluation of biosensors. Detailed Implementation

[0039] The present invention will be further described below with reference to specific embodiments in order to better understand the technical solution.

[0040] Example 1 Electrode structure of implantable biosensor like Figure 1 As shown, the film layer covering the conductive material of the biosensor electrode in the electrode structure includes at least a sensing layer and a confinement layer. The conductive material of the electrode is platinum, and the platinum layer is located on the insulating substrate of the sensor. The sensing layer includes a cation layer 20 and a first enzyme layer 30 covering the platinum layer 10 of the electrode. The confinement layer includes a first polyurethane layer 40 covering the sensing layer. In a preferred embodiment, the cation layer 20 and the first enzyme layer 30 of the sensing layer can be alternately arranged several times to form a multilayer sensing layer structure. For example, alternating arrangement twice means that the sensing layer on the platinum layer is sequentially arranged with a cation layer, a first enzyme layer, another cation layer, and another first enzyme layer.

[0041] In addition to the sensing layer and the confinement layer, the outer membrane layer may further include an optimization layer, which covers the confinement layer. The optimization layer includes a second enzyme layer 50 covering the confinement layer and a second polyurethane layer 60 covering the enzyme layer. The second polyurethane layer may be one or more layers.

[0042] In some embodiments, such as Figure 2In the electrode structure shown, the film layer covering the conductive material of the biosensor electrode includes at least a sensing layer and a confinement layer. The conductive material of the electrode is selected from platinum, and the platinum layer is located on the insulating substrate of the sensor. The sensing layer includes a cation-enzyme mixture layer 200 covering the platinum layer 100 of the electrode, and the confinement layer includes a polyurethane layer 300 covering the sensing layer.

[0043] The cationic layer 20 or the mixture layer 200 of cationic and enzyme layers contains a cationic polymer selected from, but not limited to, polydiallyldimethylammonium chloride (PDDA), polyethyleneimine (PEI), and poly(allylamine hydrochloride) (PAAH). The concentration of the PDDA solution is selected from, but not limited to, 0.5-1% (wt, mass percentage), the concentration of the polyethyleneimine solution is selected from, but not limited to, 1-4% (wt), and the concentration of the PAAH solution is selected from, but not limited to, 1% (wt). The solvent is pure water. The first enzyme layer 30 and the second enzyme layer 50 contain glucose oxidase and can be formed from the same enzyme solution or different enzyme solutions. In one specific embodiment, the concentration of the enzyme solution is selected from, but not limited to, 60-120 mg / ml, and the solvent is 0.1M PBS buffer. The first polyurethane layer 40 and the second polyurethane layer 60 comprise polyurethane, formed from the same or different polyurethane solutions, and in one specific embodiment, the polyurethane concentration is selected from, but not limited to, 3% (w / v), and the solvent composition is selected from, but not limited to, 98% tetrahydrofuran and 2% N,N-dimethylformamide.

[0044] Example 2: The sensing layer on the electrode conductive material adopts a combination of independent cation layer and enzyme layer. like Figure 1 and Figure 3 The implantable glucose sensor 1000 shown can be tested at low potentials. The sensor 1000 employs a dual-electrode system with a double-sided design, including a flexible insulating substrate 1. A working electrode 2 is located on one side (front) of the substrate, connected to a first pin 71 via a first lead 61. A counter electrode 3 is located on the other side (back) of the substrate, connected to a second pin 72 via a second lead 62. The electrode terminals of the sensor are in contact with the analyte, and the pins are connected to the circuitry of the analyzer. A film layer is coated on the platinum layer of the working electrode, and the film layer sequentially comprises a sensing layer and a confinement layer. In some embodiments, an optimization layer may also be included. The sensing layer includes at least a cationic layer containing a cationic polymer and an enzyme layer containing glucose oxidase, with the enzyme layer located above the cationic layer. The confinement layer comprises polyurethane. The optimization layer includes an enzyme layer containing glucose oxidation and a polyurethane layer containing polyurethane.

[0045] The insulating substrate material of the sensor is a flexible substrate, which can be selected from, but is not limited to, polyimide, polyphenylsulfone, methyl methacrylate, polycarbonate, or mixtures thereof. The working electrode, first lead, second lead, first pin, and second pin of the sensor can be made of, but is not limited to, conductive metallic materials such as platinum, and can be formed on the substrate by sputtering or screen printing. In this example, the working electrode is approximately 0.3 mm wide and 5.0 mm long, and a sputtered platinum layer is formed on the surface of the flexible substrate using a sputtering process. The counter electrode is made of, but is not limited to, silver / silver chloride, and is formed on the substrate by sputtering or screen printing; the counter electrode is approximately 0.3 mm wide and 5.0 mm long. In use, the working electrode and counter electrode are inserted into the human body.

[0046] In this embodiment, polydiallyldimethylammonium chloride is used as the cationic layer material of the cationic polymer. The following four schemes are used to prepare the film layer (i.e., sensing layer, confinement layer and / or optimization layer) of the sensor electrode. The polydiallyldimethylammonium chloride is purchased from McLean Company and has a molecular weight of 250,000-350,000.

[0047] The fabrication process of the glucose sensor includes the following steps: Step 1: Set the working electrode, counter electrode, lead wire and contact on both sides of the insulating substrate; Step 2: Immerse the electrode terminals in a 1 M sulfuric acid solution for electrochemical cleaning; Step 3: After drying, prepare the electrode film layer according to the following four methods.

[0048] Group A The electrode was immersed in 90 mg / mL glucose oxidase (GOx) solution for 30 min, and then air-dried at room temperature (approximately 18-26℃). The treated electrode was then fixed in a brown cross-linking bottle, and 50 μL of 25% glutaraldehyde was added. The bottle was sealed and placed in a 37℃ incubator for 15 min for cross-linking. A restriction layer was then applied to the enzyme layer. The electrode covered with the enzyme layer was immersed in 3% (w / v) polyurethane solution for 2 s using a dip-coating method. This was recorded as one application, and a total of one application was performed for this group. The electrode was then air-dried for later use.

[0049] Group B The electrode was immersed in a 1% cationic solution of polydiallyldimethylammonium chloride for 30 min, then removed and dried in a 37°C incubator to obtain a platinum flexible electrode with a cationized surface. The electrode was then immersed in a 90 mg / mL GOx solution for 30 min, and air-dried at room temperature after immersion. Each step of coating with a cationic layer and an enzyme layer was counted as one cycle; this group completed one cycle, forming an electrode coated with a sensing layer. The electrode coated with the sensing layer was fixed in a brown cross-linking bottle, and 50 μL of 25% glutaraldehyde was added. The bottle was sealed and placed in a 37°C incubator for cross-linking for 15 min. A confinement layer was then applied to the sensing layer. The electrode coated with the sensing layer was immersed in a 3% (w / v) polyurethane solution for 2 s using a dip-coating method; this was counted as one cycle; this group completed one cycle. The electrode was then air-dried for later use.

[0050] Group C The electrode was immersed in a 1% cationic solution of polydiallyldimethylammonium chloride for 30 min, then removed and dried in an incubator to obtain a platinum flexible electrode with a cationic surface. The electrode was then immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature. Each step of coating with a cationic and enzyme layer was counted as one cycle. This group repeated the above steps once more, for a total of two cycles, to form an electrode coated with a sensing layer. The electrode coated with the sensing layer was fixed in a brown cross-linking bottle, and 50 μL of 25% glutaraldehyde was added. After sealing, it was placed in an incubator at 37°C for 15 min for cross-linking. A confinement layer was then applied to the sensing layer. The electrode coated with the sensing layer was immersed in a 3% (w / v) polyurethane solution for 2 s using a dip-coating method, and then removed. This was counted as one cycle, and this group completed one cycle. The electrode was then air-dried for later use.

[0051] Group D The electrode was immersed in a 1% cationic solution of polydiallyldimethylammonium chloride for 30 min, then removed and dried in an incubator to obtain a platinum flexible electrode with a cationic surface. The electrode was then immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature. Each step of coating with a cationic layer and an enzyme layer was considered one cycle. This group repeated the above steps once more, for a total of two cycles, to form an electrode coated with a sensing layer. The electrode coated with the sensing layer was fixed in a brown cross-linking bottle, and 50 μL of 25% glutaraldehyde was added. After sealing, it was placed in an incubator at 37°C for 15 min for cross-linking. A confinement layer was then applied to the sensing layer. Using a dip-coating method, the electrode coated with the sensing layer was immersed in a 3% (w / v) polyurethane solution for 2 seconds, then removed and dried. The electrode was then immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature. A 3% (w / v) polyurethane solution was then dip-coated twice onto the enzyme layer to form an optimization layer. The layer was then dried and ready for use.

[0052] Sensors prepared in groups A, B, C, and D were used to test glucose samples of different concentrations at a low potential of -0.1V. The glucose in the test samples was dissolved in 5 mL of 0.01 M PBS buffer (pH 7.2-7.4) at concentrations of 0 mM, 5 mM, 10 mM, 15 mM, 20 mM, 25 mM, 30 mM, and 35 mM, respectively. The results are shown in Table 1. Figure 4 and Figure 5 As shown, the sensor prepared with formulation A did not exhibit good linear response to glucose samples of different concentrations at a low potential of -0.1V. Sensors prepared with formulations B, C, and D showed good linear response and sensitivity to glucose samples of different concentrations at a low potential of -0.1V, with a linear range ≥30 mM and R0. 2 ≥0.99.

[0053]

[0054] Note: 3%PU*1 means covering with one layer of 3%PU, (PDDA+enzyme)*2 means covering with two layers of PDDA and enzyme alternately, and so on.

[0055] Example 3: The sensing layer on the electrode conductive material is a hybrid layer formed by pre-mixing cations and enzymes. like Figure 2 As shown, this embodiment uses a mixture layer formed by pre-mixing cations and enzymes as the sensing layer to prepare a glucose sensor membrane and glucose sensor capable of low-potential testing. The cationic polymer in this example is poly(acrylamide-co-diallyldimethylammonium chloride) (Sigma, 409081). The sensing layer is a mixture layer containing the cationic polymer and glucose oxidase, and the confinement layer includes a polyurethane layer covering the sensing layer.

[0056] like Figure 2 As shown, in this embodiment, poly(acrylamide-co-diallyldimethylammonium chloride) is used as the cationic polymer material, and the film layer (i.e., sensing layer, confinement layer) of the sensor electrode is prepared by the following scheme.

[0057] The fabrication process of the glucose sensor includes the following steps: Step 1: Set the working electrode, counter electrode, lead wire, and pin on both sides of the insulating substrate; Step 2: Immerse the electrode terminals in a 1 M sulfuric acid solution for electrochemical cleaning; Step 3: After drying, prepare the electrode film layer according to the following method.

[0058] Reagents for preparing the sensing layer: a mixed solution containing 1% cationic polymer and 90 mg / mL GOx (the solvent is water; the cations are dissolved in water in proportion to form a solution, and then GOx is added to it in proportion).

[0059] The electrode was dipped in a mixed solution of 1% cationic polymer and 90 mg / mL GOx for 3 seconds, dried at room temperature for 10 minutes, and then fixed in a brown cross-linking bottle with a mixed cationic and enzyme layer. 50 μL of 25% glutaraldehyde was added, and the bottle was sealed and placed in a 37°C incubator for 20 minutes for cross-linking. This step of coating with a mixed cationic and enzyme layer was recorded as one cycle; in this example, a total of 4 cycles were performed to form an electrode covered with a sensing layer. A confinement layer was then coated on the sensing layer. The electrode covered with the sensing layer was immersed in a 6.5% (w / v) polyurethane solution for 2 seconds using a dip-coating method; this was recorded as one cycle, and a total of 3 cycles were performed in this example. The electrode was then air-dried for later use.

[0060] The prepared sensor was tested with glucose samples of different concentrations at a low potential of -0.1V. For example... Figure 6 As shown, the sensor prepared according to the above formula exhibits good linear response to samples with glucose concentrations in the range of 0-35 mM at a low potential of -0.1V, with a sensitivity of 1.60 nA / mM and a linear range ≥35 mM. 2 ≥0.99.

[0061] Example 4: Fabrication of a glucose sensor using different cationic polymers as sensing layer components In this embodiment, a glucose sensor was prepared using cationic polymer components, namely polydiallyldimethylammonium chloride, polyethyleneimine, and poly(allylamine hydrochloride), as the sensing layer. The process includes the following steps: Step 1: Set the working electrode, counter electrode, lead wire, and pin on both sides of the insulating substrate; Step 2: Immerse the electrode terminals in a 1 M sulfuric acid solution for electrochemical cleaning; Step 3: Prepare 1% (wt) aqueous solutions of polydiallyldimethylammonium chloride, polyethyleneimine, and poly(allylamine hydrochloride) using ultrapure water. After preparation, remove air bubbles by ultrasonication. The molecular weight of polydiallyldimethylammonium chloride (PDDA) is 250,000-350,000, the molecular weight of polyethyleneimine (PEI) is 25,000, and the molecular weight of poly(allylamine hydrochloride) (PAH) is 15,000. All were purchased from McLean Company.

[0062] The pretreated electrodes from step two were immersed in the three cationic polymer solutions mentioned above for 30 minutes at room temperature. After immersion, they were removed and dried in a constant temperature oven to obtain platinum electrodes with cationized surfaces.

[0063] The platinum electrode with cationized surface was immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature. The dried platinum flexible electrode was then immersed in a 1% cation solution for 30 min, removed and air-dried at room temperature. Finally, the flexible electrode was immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature.

[0064] Step 4: Fix the electrode covered with the sensing layer in Step 3 into a brown crosslinking bottle, add 50 μL of 25% glutaraldehyde, seal it, and place it in a 37℃ constant temperature oven for crosslinking for 15 min.

[0065] Step 5: Immerse the electrode treated in Step 4 in a 3% (w / v) polyurethane solution for 2 seconds and then remove it.

[0066] Step Six: Immerse the electrode treated in Step Five in a 90 mg / mL GOx solution for 30 min. After immersion, air dry at room temperature. Dip the enzyme layer twice with 3% (w / v) polyurethane to form an optimized layer. Air dry and set aside.

[0067] Sensors prepared using the three different cationic polymers described above were used to test glucose samples of different concentrations at a low potential of -0.1V. The test results are as follows: Figure 7 As shown, sensors prepared from three different cationic polymers exhibited good linear response to glucose concentration samples in the range of 0-35 mM at a low potential of -0.1 V, with sensitivities of 3.5793, 2.041, and 5.4162 nA / mM, respectively, and linear range ≥35 mM. R0 2 ≥0.99.

[0068] Example 5: Performance evaluation of sensor sensing layer using cationic polymers of different concentrations In this embodiment, the sensing layer of the sensor was prepared using cationic polymers of different concentrations. Polyethyleneimine was used as the cationic polymer component in this embodiment, and the preparation process included the following steps: Step 1: Set the working electrode, counter electrode, lead wire and contact on both sides of the insulating substrate; Step 2: Immerse the platinum electrode in a 1 M sulfuric acid solution for electrochemical cleaning; Step 3: Prepare 1%, 2% and 4% aqueous solutions of polyethyleneimine using ultrapure water, and remove air bubbles by sonication after preparation.

[0069] The electrodes pretreated in step two were immersed in the above-mentioned polyethyleneimine solution at room temperature for 30 minutes. After immersion, they were removed and dried in a constant temperature oven for 20 minutes to obtain platinum electrodes with cationized surfaces.

[0070] The platinum electrode with cationized surface was immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature. The dried platinum flexible electrode was then immersed in a 1% cation solution for 30 min, and then air-dried at room temperature. Finally, the flexible electrode was immersed in a 90 mg / mL GOx solution for 30 min, and then air-dried at room temperature.

[0071] Step 4: Fix the electrode covered with the sensing layer in Step 3 into a brown crosslinking bottle, add 50 μL of 25% glutaraldehyde, seal it, and place it in a 37℃ constant temperature oven for crosslinking for 15 min.

[0072] Step 5: Immerse the working electrode portion treated in Step 4 in a 3% (w / v) polyurethane solution for 2 seconds and then remove it.

[0073] Step Six: Immerse the electrode treated in Step Five in a 90 mg / mL GOx solution for 30 min, and then air dry at room temperature. Coat the enzyme layer twice with 3% (w / v) polyurethane to form an optimized layer. Air dry for later use.

[0074] Sensors prepared using the aforementioned cationic polymers at different concentrations were used to test glucose samples of different concentrations at a low potential of -0.1V. Experimental results are shown below. Figure 8 As shown, 1-4% (w / v) of polyethyleneimine used to prepare glucose sensors exhibits good detection sensitivity.

[0075] Example 6: Evaluation of the anti-interference performance of biosensors against various interfering factors The difference between the sensor fabrication in this embodiment and that in Embodiment 5 is that the cationic polymer used is 1% (w / v) polydiallyldimethylammonium chloride. Acetaminophen, ascorbic acid, and uric acid were selected as interfering agents for glucose testing, and tests were conducted at -0.1 V and +0.55 V, respectively.

[0076] The testing procedure is as follows: Taking acetaminophen as an example, the test was performed in 5 mL of 0.01 M PBS buffer (pH 7.2-7.4) at room temperature. The electrode was first stabilized in PBS for 5-10 min. The experimental group's test voltage was -0.1 V, and the control group's test voltage was +0.55 V. After the baseline stabilized, glucose (GLU) was added to a final concentration of 6.6 mM. After the signal stabilized, acetaminophen (ACE) was added to a final concentration of 4 mg / dL. Similarly, ascorbic acid (VC) and uric acid (UA) to a final concentration of 6 mg / dL were tested in the same manner. The test results are shown in Table 2 and... Figure 9As shown, under the test condition of -0.1V, the current interference of acetaminophen, uric acid, and ascorbic acid on glucose (MARD, calculated as (interfering current value - glucose current value) / glucose current value × 100%, used to reflect the degree of influence of the interfering substance on the test current) is 0.97%, 10.14%, and 4.83%, respectively, which is much smaller than that of 48.30%, 88.26%, and 17.47% under the test condition of +0.5V. This indicates that the sensor made of cationic polymer has better anti-interference performance at low potential.

[0077] Other cationic polymers, such as polyethyleneimine and poly(allylamine hydrochloride), also exhibit excellent anti-interference properties.

[0078]

Claims

1. An electrode for an implantable biosensor, the conductive layer of the electrode being located on an insulating substrate, a limiting layer being provided on the conductive layer of the electrode, the electrode comprising a first electrode and a second electrode, a sensing layer being provided on at least the first electrode, the sensing layer being located between the conductive layer and the limiting layer, the sensing layer comprising an enzyme required for a reaction, characterized in that, The sensing layer further comprises a cationic polymer.

2. The electrode of claim 1, wherein The sensing layer comprises a cationic layer and an enzyme layer, the cationic layer is on the conductive layer, the cationic layer comprises a cationic polymer, and the enzyme layer comprises an enzyme required for reaction; or the sensing layer comprises a mixed layer mixed by the cationic polymer and the enzyme.

3. The electrode of claim 1, wherein An optimization layer is further provided on the limiting layer of the electrode, the optimization layer comprises an enzyme layer and a polyurethane layer, the enzyme layer is on the limiting layer, and the polyurethane layer is on the enzyme layer.

4. The electrode of claim 1, wherein The cationic polymer is selected from one of polydiallyldimethylammonium chloride, a polydiallyldimethylammonium chloride derivative, polyethyleneimine, a polyethyleneimine derivative, poly(allylamine hydrochloride), a poly(allylamine hydrochloride) derivative, or a combination of two or more thereof.

5. The electrode of claim 1, wherein The molecular weight of the cationic polymer is 15000-700000.

6. The electrode of claim 1, wherein The biosensor is used for glucose detection, and the enzyme is glucose oxidase.

7. The electrode of claim 1, wherein The first electrode is a working electrode.

8. The electrode of claim 1, wherein The second electrode is selected from a counter electrode or a reference electrode.

9. A biosensor comprising a flexible insulating substrate, an electrode on the flexible insulating substrate, a lead and a pin, characterized in that, The electrode is the electrode according to any one of claims 1 to 8.

10. Use of the biosensor according to claim 9 in continuous blood glucose monitoring.

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