A biochemical sensor based on an epitaxial gate field effect transistor and its application

Abstract

The present invention relates to the field of biochemical sensors, and discloses a biochemical sensor based on an epitaxial gate type field effect transistor and its application. The biochemical sensor of the present invention includes a substrate, a field effect transistor, a sensitive electrode, and a reference electrode. Among them, the sensitive electrode is electrically connected to the gate of the field effect transistor in the form of an epitaxial gate. The sensitive electrode includes a conductive substrate, and an ordered mesoporous carbon layer, a hydrogen peroxide catalyst, and an oxidoreductase sequentially modified on the surface of the conductive substrate. The sensitive electrode of the present invention exists in the form of an epitaxial gate, which is convenient for replacement, and appropriate oxidoreductases can be selected for different analytes to be detected, so that the biochemical sensor can achieve high-sensitivity detection of a variety of analytes with high redox potentials; since only the sensitive electrode as the epitaxial gate and the reference electrode contact the liquid to be detected during detection, and the field effect transistor does not contact the liquid to be detected, the influence of the solution on the performance of the field effect transistor can be reduced and the device stability can be improved.

Description

Technical Field

[0001] The present invention relates to the field of biochemical sensors, and particularly to a biochemical sensor based on an epitaxial gate field effect transistor and its application. Background Art

[0002] Biochemical small molecules in various human body fluids (such as blood, sweat, saliva, urine, tissue fluid, etc.), such as glucose, uric acid, lactic acid, cholesterol, etc., have a close relationship with the health of a person. Therefore, the accurate detection of these molecules is of great significance for monitoring the health status of the human body and providing a basis for diagnosis.

[0003] Currently, commonly used detection methods include chromatography, electrochemical methods, etc. These methods rely on bulky detection equipment and cannot achieve rapid and convenient point-of-care testing. Electrochemical sensor devices based on transistors have excellent research prospects due to their high detection sensitivity and easy integration characteristics.

[0004] For existing electrochemical sensor devices based on transistors, their detection ability depends on the gate electrode or solution gate that is sensitive to the analyte. Therefore, the inherent characteristics of the transistor impose certain constraints on this detection ability. First, the redox potentials of different molecules are different. Due to the differences in the interface capacitances in the gate-source loop of the transistor, it is difficult for the gate and solution interface to reach the potential required for the reaction. Therefore, the types of detectable substances are limited, generally only including ascorbic acid, dopamine, hydrogen peroxide, etc. with low redox potentials. Second, the top gate of the transistor is in direct contact with the solution, which places high requirements on the waterproof performance of the transistor in the solution. Devices with improper packaging are easily damaged by moisture, affecting the detection results.

[0005] Therefore, researching new biochemical sensor devices based on field effect transistors to achieve highly sensitive detection of various analytes with high redox potentials, reducing the impact of the solution on the transistor performance, and improving the device stability is an urgent problem to be solved. Summary of the Invention

[0006] To solve the above technical problems, firstly, the present invention provides a biochemical sensor based on an epitaxial gate type field effect transistor; secondly, the present invention provides a biochemical small molecule detection system and a detection method. The biochemical sensor of the present invention is provided with a sensitive electrode in the form of an epitaxial gate. The advantages of this design are as follows: on the one hand, since the sensitive electrode exists in the form of an epitaxial gate, it is convenient to replace, and appropriate types of redox enzymes can be selected according to different analytes to be detected, with strong universality, enabling the biochemical sensor of the present invention to achieve highly sensitive detection of a variety of analytes with high redox potentials; on the other hand, since only the sensitive electrode as the epitaxial gate and the reference electrode are in contact with the liquid to be detected during detection, while the field effect transistor is not in contact with the liquid to be detected, the influence of the solution on the performance of the field effect transistor can be reduced and the device stability can be improved.

[0007] The specific technical solution of the present invention is as follows:

[0008] In the first aspect, the present invention provides a biochemical sensor based on an epitaxial gate type field effect transistor, including:

[0009] A substrate; the substrate has various selections according to the forms of the field effect transistor and the sensitive electrode, and can be but not limited to a silicon substrate, calcium sodium glass, a flexible substrate such as polyethylene naphthalate (PEN), etc.

[0010] A field effect transistor; the field effect transistor is fixed on the surface of the substrate and includes a drain, a gate, and a source; the field effect transistor can be but not limited to a single crystal silicon field effect transistor, a low temperature polycrystalline silicon field effect transistor, an oxide field effect transistor, an organic field effect transistor, etc.

[0011] A sensitive electrode, the sensitive electrode is fixed on the surface of the substrate as an epitaxial gate and is electrically connected to the gate of the field effect transistor; during detection, the sensitive electrode is immersed in the solution to be detected and is sensitive to the analyte. According to the form of the sensitive electrode, the connection method of the sensitive electrode to the gate of the field effect transistor includes: direct wiring connection on the surface of the substrate for a planar conductive substrate and silver paste connection of the fiber and the exposed part of the gate for a fibrous conductive substrate. The sensitive electrode includes:

[0012] A conductive substrate;

[0013] An ordered mesoporous carbon layer, the ordered mesoporous carbon layer covers the surface of the conductive substrate;

[0014] A hydrogen peroxide catalyst, the hydrogen peroxide catalyst is loaded in the pores of the ordered mesoporous carbon layer, which is used to catalyze the product hydrogen peroxide of the enzyme reaction and transfer the electrons after oxidizing hydrogen peroxide to the ordered mesoporous carbon layer and the conductive substrate;

[0015] An oxidoreductase, which is loaded in the pores of the ordered mesoporous carbon layer and is used to selectively catalyze the substrate to be hydrogen peroxide, so as to realize the conversion of the detection of various analytes into the detection of hydrogen peroxide.

[0016] A reference electrode, which is fixed on the surface of the substrate. The reference electrode is immersed in the solution to be detected during detection to maintain a stable potential. Together with the sensitive electrode, it constitutes a part of the solution potential circuit between the gate and the source of the field effect transistor. The gate-source voltage of the field effect transistor is adjusted by the bias voltage at the reference electrode terminal. The reference electrode can be, but is not limited to, a silver / silver chloride reference electrode, a mercury / mercuric chloride reference electrode, a standard hydrogen electrode, a gold electrode, etc., corresponding to the form of the sensitive electrode. The reference electrode can also be made into a fibrous or planar shape as required.

[0017] As can be seen from the above, the biochemical sensor of the present invention mainly includes a substrate, a field effect transistor, a sensitive electrode and a reference electrode. Specifically, an ordered mesoporous carbon layer is covered on the conductive substrate of the sensitive electrode of the present invention, and a hydrogen peroxide catalyst and an oxidoreductase are loaded in the ordered mesoporous carbon layer. Under the action of oxygen, the oxidoreductase can oxidize the analyte to produce hydrogen peroxide. The hydrogen peroxide diffuses to the surface of the hydrogen peroxide catalyst and is oxidized to oxygen, while transferring the generated electrons to the surface covered with the ordered mesoporous carbon layer. Due to the good conductivity of the mesoporous carbon, the electrons are further transmitted to the conductive substrate. The sensitive electrode exists in the form of an epitaxial gate. The hydrogen peroxide catalyst and oxidoreductase fixed on its surface can convert the concentration information of the analyte in the solution to be detected into the potential change at the electrode-solution interface, while the potential difference between the reference electrode and the solution remains unchanged. The field effect transistor converts the voltage drop information between the gate and the channel into the output drain current. The advantages of the above structural design are as follows: (1) Since the sensitive electrode exists in the form of an epitaxial gate, it is convenient to replace, and appropriate types of oxidoreductases can be selected for different analytes to be detected, with strong universality, so that the biochemical sensor of the present invention can achieve highly sensitive detection of a variety of analytes with high redox potentials; (2) Since only the sensitive electrode as the epitaxial gate and the reference electrode are in contact with the solution to be detected during detection, and the field effect transistor is not in contact with the solution to be detected, the influence of the solution on the performance of the field effect transistor can be reduced and the device stability can be improved.

[0018] Preferably, the biochemical sensor based on the epitaxial gate type field effect transistor further includes a packaging layer that only exposes the sensitive gate and reference electrode regions. Further, the packaging layer material can be, but is not limited to, polyethylene.

[0019] The packaging layer plays a protective role and can reduce the influence of the solution to be detected on the device.

[0020] Preferably, in the sensitive electrode: the pore size of the ordered mesoporous carbon layer is 5-20 nm. The particle size of the hydrogen peroxide catalyst is less than 5 nm. The size of the redox enzyme is less than 20 nm.

[0021] In addition to good electrical conductivity, the advantages of the ordered mesoporous carbon layer of the present invention also include that it can significantly increase the specific surface area of the electrode, enabling the electrode to load more catalysts; for enzyme loading, the ordered mesoporous carbon also provides a confined space for enzyme loading, making the enzyme loading more stable.

[0022] Furthermore, the research team of the present invention found that the ordered mesoporous carbon layer with the above pore size can not only achieve the transfer of electrons from the hydrogen peroxide catalyst to the conductive substrate, but also achieve the confined loading of the hydrogen peroxide catalyst and the redox enzyme. If the pore size is too small, the loading of the hydrogen peroxide catalyst and the redox enzyme cannot be achieved; if the pore size is too large, multiple enzyme molecules will be adsorbed in the pores, reducing the catalytic efficiency.

[0023] Preferably, the conductive substrate is a fibrous conductive substrate or a planar conductive substrate; the fibrous conductive substrate can be, but is not limited to, a carbon fiber electrode, a carbon nanotube fiber electrode, etc.; the planar conductive substrate can be, but is not limited to, a gold electrode, a silver electrode, an indium tin oxide electrode, etc.

[0024] Preferably, the hydrogen peroxide catalyst can be, but is not limited to, noble metal nanoparticles such as platinum and gold, Prussian blue nanoparticles, etc.

[0025] Preferably, the redox enzyme can be, but is not limited to, glucose oxidase that catalyzes glucose, uricase that catalyzes uric acid, cholesterol oxidase that catalyzes cholesterol, lactate oxidase that catalyzes lactate, etc.

[0026] As a further preference, for the fibrous conductive substrate, the ordered mesoporous carbon layer is formed by in-situ growth of an ordered mesoporous carbon material on the surface of the fibrous conductive substrate through a hydrothermal reaction followed by high-temperature calcination; for the planar conductive substrate, the ordered mesoporous carbon layer is formed by drop-coating an ordered mesoporous carbon material suspension on the planar conductive substrate.

[0027] Preferably, the preparation method of the sensitive electrode includes the following steps:

[0028] Step 1, modify an ordered mesoporous carbon layer on the surface of the conductive substrate.

[0029] Step 2, modify a hydrogen peroxide catalyst on the surface of the ordered mesoporous carbon layer.

[0030] Step 3: Further modify the surface of the ordered mesoporous carbon layer modified with the hydrogen peroxide catalyst with redox enzyme; the modification method of the redox enzyme is physical adsorption using the confinement effect of the ordered mesoporous carbon layer or chemical adsorption between the redox enzyme and the ordered mesoporous carbon layer.

[0031] As a further preference, in Step 1:

[0032] For the fibrous conductive substrate, directly modify and grow the single micelle composed of organic small molecules / surfactants on the surface of the fibrous conductive substrate by means of interfacial assembly, and form an ordered mesoporous carbon layer after calcination.

[0033] For the planar conductive substrate, directly drop and cover the surface of the planar conductive substrate with the pre-prepared ordered mesoporous carbon material suspension, and form an ordered mesoporous carbon layer after drying, with a surface mass density of 1-3 μg / mm 2 .

[0034] More specifically, for the sensitive electrode using a fibrous conductive substrate, its preparation method includes the following steps:

[0035] Step A: Prepare a low-order phenolic resin solution and a triblock copolymer solution respectively, then mix the low-order phenolic resin solution and the triblock copolymer solution, heat and stir to obtain a spherical micelle solution.

[0036] Step B: Take the above spherical micelle solution, dilute it with water, and at the same time immerse the fibrous conductive substrate in the obtained solution, and carry out a hydrothermal reaction to obtain a fibrous conductive substrate with ordered micelles grown on its surface.

[0037] Step C: Calcinate the fibrous conductive substrate with ordered micelles grown on its surface in an inert gas atmosphere to obtain a fibrous conductive substrate with an ordered mesoporous carbon layer modified on its surface.

[0038] Step D: Prepare a precursor solution of the hydrogen peroxide catalyst, immerse the fibrous conductive substrate with an ordered mesoporous carbon layer modified on its surface in the precursor solution, and then add a reducing agent. After the reduction reaction, obtain a fibrous conductive substrate loaded with the hydrogen peroxide catalyst.

[0039] Step E: Immerse the fibrous conductive substrate loaded with the hydrogen peroxide catalyst in the redox enzyme solution, let it stand, take it out, and prepare the sensitive electrode.

[0040] In the second aspect, the present invention provides a biochemical small molecule detection system, including:

[0041] A biochemical sensor based on an epitaxial gate type field effect transistor;

[0042] A voltage bias and current reading unit; the voltage bias and current reading unit is electrically connected to the biochemical sensor, and is used to supply power to the biochemical sensor and read the output leakage current;

[0043] A signal processing unit; the signal processing unit is electrically connected to the voltage bias and current reading unit, and is used to analyze the read leakage current signal and convert it into the concentration information of the analyte.

[0044] In a third aspect, the present invention provides a method for detecting biochemical small molecules, and the above-mentioned biochemical small molecule detection system is used to detect the analyte. When detecting, the sensitive electrode and the reference electrode are immersed in the solution to be detected, a constant bias voltage is set at the reference electrode end to control the gate voltage of the field effect transistor, and at the same time, a constant source-drain voltage is set. After the sensitive electrode is in full contact with the solution, because the enzyme molecule catalyzes the production of hydrogen peroxide for a specific analyte, and the surface potential of the electrode loaded with the hydrogen peroxide catalyst is affected by the concentration of hydrogen peroxide, the voltage drop at the interface between the sensitive electrode and the solution is indirectly affected by the concentration of the target molecule. In this way, when the gate-source voltage is fixed, the voltage drop between the gate electrode and the channel of the transistor is also affected by the concentration of the target molecule in the solution to be analyzed. According to the transconductance characteristics of the transistor, the output leakage current is correlated with the concentration of the target molecule in the solution to be analyzed. Therefore, the concentration of the target molecule in the solution can be judged according to the magnitude of the leakage current.

[0045] Preferably, in order to improve the sensitivity of the device, the operating region of the field effect transistor is preferably the subthreshold region. Therefore, it is necessary to first scan the transfer characteristic curve in a solution without target biomolecules, and then select a suitable gate-source operating voltage to make the field effect transistor operate in the subthreshold region.

[0046] Preferably, in order to convert the read leakage current value into the concentration of the target molecule in the solution, it is necessary to pre-measure the current values of the transistor device in solutions with different target molecule concentrations as the calibration curve.

[0047] Compared with the prior art, the beneficial effects of the present invention are:

[0048] (1) The sensitive electrode in the biochemical sensor of the present invention exists in the form of an epitaxial gate, so it is easy to replace, and the appropriate type of redox enzyme can be selected according to different analytes to be detected, with strong universality, making the biochemical sensor of the present invention capable of achieving highly sensitive detection of a variety of analytes with high redox potentials.

[0049] (2) When the biochemical sensor of the present invention is detecting, only the sensitive electrode as the epitaxial gate and the reference electrode are in contact with the test solution, while the field effect transistor is not in contact with the test solution. Therefore, the contact between the channel layer and the test solution can be avoided, improving the stability and reusability of the field effect transistor, and also making the detection accuracy higher.

[0050] (3) The sensitive electrode of the present invention can be processed by solution method / printing process, with low cost and strong customization. Description of the Drawings

[0051] Figure 1 is a schematic diagram of the preparation process of the fibrous glucose-sensitive electrode in Example 1 of the present invention;

[0052] Figure 2 is a scanning electron microscope image of the fibrous glucose-sensitive electrode prepared in Example 1 of the present invention;

[0053] Figure 3 is a schematic diagram of the structure of the biochemical sensor in Example 2 of the present invention;

[0054] Figure 4 is a schematic diagram of the structure of the biochemical sensor in Example 3 of the present invention;

[0055] Figure 5 is a schematic diagram of the structure of the biochemical small molecule detection system in Example 4 of the present invention;

[0056] Figure 6 is a flowchart of the biochemical small molecule detection method in Example 5 of the present invention;

[0057] Figure 7 is the calibration curve of the output current and glucose concentration in Example 5 of the present invention.

[0058] The reference numerals are: substrate 1, field effect transistor 2, fibrous glucose-sensitive electrode 3, fibrous reference electrode 4, encapsulation layer 5, silver paste 6, copper electrode 7, planar reference electrode 8, planar glucose-sensitive electrode 9, planar gold electrode 10, sensitive electrode 11, reference electrode 12, voltage bias and current reading unit 13, signal processing unit 14, polymer tank 15, drain 21, gate 22, source 23, sensitive electrode carbon fiber 31, ordered mesoporous carbon layer 32, platinum nanoparticles 33, glucose oxidase 34, reference electrode carbon fiber 41, silver / silver chloride layer 42, polyvinyl butyral layer containing saturated sodium chloride 43. Detailed Embodiments

[0059] The present invention will be further described below in conjunction with the embodiments.

[0060] General Embodiment

[0061] A biochemical sensor based on an epitaxial gate type field effect transistor, comprising:

[0062] A substrate; the substrate has various selections according to the forms of the field effect transistor and the sensitive electrode, and can be but not limited to a silicon substrate, calcium sodium glass, a flexible substrate such as polyethylene naphthalate (PEN), etc.;

[0063] A field-effect transistor; the field-effect transistor is fixed on the surface of a substrate and includes a drain, a gate, and a source; the field-effect transistor may be, but is not limited to, a single-crystalline silicon field-effect transistor, a low-temperature polycrystalline silicon field-effect transistor, an oxide field-effect transistor, an organic field-effect transistor, etc.;

[0064] A sensitive electrode, the sensitive electrode is fixed on the surface of the substrate as an epitaxial gate and is electrically connected to the gate of the field-effect transistor; during detection, the sensitive electrode is immersed in the solution to be detected and is sensitive to the analyte. According to the form of the sensitive electrode, the connection method of the sensitive electrode to the gate of the field-effect transistor includes: direct wiring connection on the surface of the substrate for a planar conductive substrate and silver paste connection of the fiber and the exposed part of the gate for a fibrous conductive substrate;

[0065] The sensitive electrode includes:

[0066] A conductive substrate;

[0067] An ordered mesoporous carbon layer, the ordered mesoporous carbon layer covers the surface of the conductive substrate; it is used to provide a confined space for the physical adsorption of enzymes; a hydrogen peroxide catalyst, the hydrogen peroxide catalyst is loaded in the pores of the ordered mesoporous carbon layer, and it is used to catalyze the product hydrogen peroxide of the enzyme reaction and transfer the electrons after oxidizing hydrogen peroxide to the ordered mesoporous carbon layer and the conductive substrate;

[0068] An oxidoreductase, the oxidoreductase is loaded in the pores of the ordered mesoporous carbon layer, and it is used to selectively catalyze the substrate to be hydrogen peroxide to realize the conversion of the detection of various analytes into the detection of hydrogen peroxide.

[0069] A reference electrode, the reference electrode is fixed on the surface of the substrate. During detection, it is immersed in the solution to be detected to maintain a stable potential, and it forms a part of the solution potential circuit between the gate and the source of the field-effect transistor together with the sensitive electrode. The gate-source voltage of the field-effect transistor is adjusted by the bias voltage at the reference electrode end. The reference electrode may be, but is not limited to, a silver / silver chloride reference electrode, a mercury / mercury chloride reference electrode, a standard hydrogen electrode, a gold electrode, etc., corresponding to the form of the epitaxial gate. The reference electrode can also be made into a fibrous or planar shape as required.

[0070] Preferably, the biochemical sensor based on the epitaxial gate type field-effect transistor further includes a packaging layer that only exposes the sensitive gate and the reference electrode regions. Further, the packaging layer material may be, but is not limited to, polyethylene.

[0071] Preferably, in the sensitive electrode: the pore diameter of the ordered mesoporous carbon layer is 5 - 20 nm; the particle size of the hydrogen peroxide catalyst is less than 5 nm; the size of the oxidoreductase is less than 20 nm.

[0072] Preferably, the conductive substrate is a fibrous conductive substrate or a planar conductive substrate; the fibrous conductive substrate can be, but is not limited to, a carbon fiber electrode, a carbon nanotube fiber electrode, etc.; the planar conductive substrate can be, but is not limited to, a gold electrode, a silver electrode, an indium tin oxide electrode, etc. The hydrogen peroxide catalyst can be, but is not limited to, noble metal nanoparticles such as platinum and gold, Prussian blue nanoparticles, etc. The oxidoreductase can be, but is not limited to, glucose oxidase that catalyzes glucose, uricase that catalyzes uric acid, cholesterol oxidase that catalyzes cholesterol, lactate oxidase that catalyzes lactate, etc.

[0073] More preferably, for the fibrous conductive substrate, the ordered mesoporous carbon layer is formed by in-situ growth of an ordered mesoporous carbon material on the surface of the fibrous conductive substrate through a hydrothermal reaction followed by high-temperature calcination; for the planar conductive substrate, the ordered mesoporous carbon layer is formed by drop-coating an ordered mesoporous carbon material suspension on the planar conductive substrate.

[0074] Preferably, the preparation method of the sensitive electrode includes the following steps:

[0075] Step 1, modify an ordered mesoporous carbon layer on the surface of the conductive substrate.

[0076] Step 2, modify a hydrogen peroxide catalyst on the surface of the ordered mesoporous carbon layer.

[0077] Step 3, continue to modify an oxidoreductase on the surface of the ordered mesoporous carbon layer modified with the hydrogen peroxide catalyst; the modification method of the oxidoreductase is physical adsorption using the confinement effect of the ordered mesoporous carbon layer or chemical adsorption between the oxidoreductase and the ordered mesoporous carbon layer.

[0078] More preferably, in Step 1:

[0079] For the fibrous conductive substrate, a single micelle composed of an organic small molecule / surfactant is directly modified and grown on the surface of the fibrous conductive substrate by interfacial assembly, and an ordered mesoporous carbon layer is formed after calcination.

[0080] For the planar conductive substrate, a pre-prepared ordered mesoporous carbon material suspension is directly dropped and covered on the surface of the planar conductive substrate, and an ordered mesoporous carbon layer is formed after drying, with a surface mass density of 1-3 μg / mm 2 .

[0081] More specifically, for the sensitive electrode using a fibrous conductive substrate, its preparation method includes the following steps:

[0082] Step A: Prepare a low-order phenolic resin solution and a triblock copolymer solution respectively, then mix the low-order phenolic resin solution and the triblock copolymer solution, heat and stir to obtain a spherical micelle solution.

[0083] Step B: Take the above spherical micelle solution, dilute it with water, and at the same time immerse the fibrous conductive substrate in the obtained solution, and perform a hydrothermal reaction to obtain a fibrous conductive substrate with ordered micelles grown on its surface.

[0084] Step C: Calcinate the fibrous conductive substrate with ordered micelles grown on its surface in an inert gas atmosphere to obtain a fibrous conductive substrate with an ordered mesoporous carbon layer modified on its surface.

[0085] Step D: Prepare a precursor solution of a hydrogen peroxide catalyst, immerse the fibrous conductive substrate with an ordered mesoporous carbon layer modified on its surface in the precursor solution, and then add a reducing agent. After the reduction reaction, a fibrous conductive substrate loaded with a hydrogen peroxide catalyst is obtained.

[0086] Step E: Immerse the fibrous conductive substrate loaded with a hydrogen peroxide catalyst in an oxidoreductase solution, let it stand, and take it out to prepare a sensitive electrode.

[0087] A biochemical small molecule detection system, comprising:

[0088] A biochemical sensor based on an epitaxial gate type field effect transistor;

[0089] A voltage bias and current reading unit; the voltage bias and current reading unit is electrically connected to the biochemical sensor for supplying power to the biochemical sensor and reading the output drain current;

[0090] A signal processing unit; the signal processing unit is electrically connected to the voltage bias and current reading unit for analyzing the read drain current signal and converting it into the concentration information of the analyte.

[0091] A biochemical small molecule detection method, using the above biochemical small molecule detection system to detect the analyte. When detecting, immerse the sensitive electrode and the reference electrode in the solution to be detected, set a constant bias voltage at the reference electrode end to control the gate voltage of the field effect transistor, and at the same time set a constant source-drain voltage. After the sensitive electrode is in full contact with the solution, because the enzyme molecule catalyzes the production of hydrogen peroxide for a specific analyte, and the surface potential of the electrode loaded with the hydrogen peroxide catalyst is affected by the concentration of hydrogen peroxide, so the voltage drop at the interface between the sensitive electrode and the solution is indirectly affected by the concentration of the target molecule. In this way, under the condition of a fixed gate-source voltage, the voltage drop between the gate electrode and the channel of the transistor is also affected by the concentration of the target molecule in the solution to be analyzed. According to the transconductance characteristic of the transistor, the output drain current is correlated with the concentration of the target molecule in the solution to be analyzed. Therefore, the concentration of the target molecule in the solution can be judged according to the magnitude of the drain current.

[0092] Preferably, in order to improve the sensitivity of the device, the working region of the field-effect transistor is preferably the subthreshold region. Therefore, it is necessary to first scan the transfer characteristic curve in a solution without the target biomolecule, and then select an appropriate gate-source operating voltage to make the field-effect transistor operate in the subthreshold region.

[0093] Preferably, in order to convert the readout drain current value into the concentration of the target molecule in the solution, it is necessary to pre-measure the current values of the transistor device in solutions with different target molecule concentrations as the calibration curve. Specific embodiments

[0095] Example 1

[0096] A fibrous glucose-sensitive electrode uses carbon fiber as the substrate, and modifies an ordered mesoporous carbon material on its surface by hydrothermal and calcination methods, and then modifies platinum metal nanoparticles and glucose oxidase on the surface of the ordered mesoporous carbon by chemical reduction and physical adsorption methods in sequence. Specifically, as Figure 1 shown, it includes the following steps:

[0097] Step 1, prepare low-order phenolic resin: Dissolve 0.6 g of phenol and 2.1 mL of 37% formaldehyde aqueous solution in 15 mL of 0.1 mol / L sodium hydroxide aqueous solution, and stir in a water bath at 67 - 70 °C for 30 minutes to obtain a light red low-order phenolic resin.

[0098] Step 2, dissolve 0.96 g of triblock copolymer F127 in 65 mL of water to obtain a clear aqueous solution, and then mix the light red low-order phenolic resin solution obtained in Step 1 with the above triblock copolymer aqueous solution, and stir in a water bath at 67 - 70 °C for 10 - 14 hours to obtain a dark red spherical micelle solution.

[0099] Step 3, take 10 g of the above spherical micelle solution, add 20 g of water for dilution, and at the same time immerse less than 0.5 g of carbon fiber in the obtained solution, and hydrothermally treat it in an oven at 130 °C for 20 hours to obtain carbon fiber with ordered micelles grown on the surface, and take out the fiber, wash it and dry it.

[0100] Step 4, calcine the carbon fiber obtained in Step 3 in an inert gas atmosphere at 750 °C for 2 hours to obtain carbon fiber with ordered mesoporous carbon modified on the surface, and the pore diameter of the ordered mesoporous carbon layer is about 10 nm.

[0101] In other embodiments, there may also be no carbon fiber in Steps 3 and 4, and directly calcine the small balls formed by self-polymerization of micelles to obtain ordered mesoporous carbon spheres, or replace the carbon fiber with other substrate materials to obtain hollow ordered mesoporous carbon spheres, ordered mesoporous carbon sheets, etc. The substrate materials include but are not limited to silica nanospheres, double metal hydroxide nanosheets, etc.

[0102] Step 5: Prepare 3 mL of 0.3 mmol / L aqueous chloroplatinic acid solution. Immerse the carbon fiber with surface-modified ordered mesoporous carbon obtained in Step 4 in the above-mentioned aqueous chloroplatinic acid solution, then add dropwise 180 μL of aqueous sodium borohydride solution with a mass fraction of 0.075%, and let it stand in a refrigerator at 4 °C for 8 hours to obtain ordered mesoporous carbon fiber loaded with platinum metal nanoparticles. The average size of the platinum metal nanoparticles is 4.3 nm (To prepare nanoparticles with a specific particle size, it is necessary to control the concentration and type of reactants. First, the concentration of chloroplatinic acid should be dilute. Second, the reducing agent should have strong reducibility to achieve explosive nucleation. At the same time, the amount of the reducing agent should exceed the amount that can reduce all chloroplatinic acid). Take out the fiber, wash it, and dry it in air.

[0103] In other embodiments, it is also possible to modify Prussian blue nanoparticles on the surface of ordered mesoporous carbon, etc. It is required that the size of the nanoparticles is less than 5 nm, which does not affect the subsequent confined loading of redox enzymes in the mesopores.

[0104] Step 6: Prepare 200 μL of 1 mg / mL glucose oxidase solution. The lattice parameters of the glucose oxidase are 5.0 nm ( 6.4 nm ( 7.6 nm. Then immerse the fiber obtained in Step 5 in the above-mentioned glucose oxidase solution, let it stand in a refrigerator at 4 °C for 8 hours, take it out, and still dry it in a refrigerator at 4 °C to obtain the fiber-shaped glucose-sensitive electrode.

[0105] In other embodiments, the corresponding redox enzyme can be selected according to the analyte to be detected. If a uric acid oxidase solution is used in Step 6, a uric acid-sensitive electrode is obtained; if a cholesterol oxidase solution is used in Step 6, a cholesterol-sensitive electrode is obtained; if lactic acid oxidase is used in Step 6, a lactic acid-sensitive electrode is obtained.

[0106] Figure 2 The transmission electron micrograph of the fiber-shaped glucose-sensitive electrode obtained according to the preparation method of Example 1 is shown. This figure shows the ordered mesoporous carbon layer on the surface of the carbon fiber, the loading of platinum metal nanoparticles on the surface of the ordered mesoporous carbon, and the physical adsorption of glucose oxidase on the ordered mesoporous carbon.

[0107] The working principle of the glucose-sensitive electrode is as follows: Under the action of oxygen, glucose oxidase oxidizes glucose to gluconolactone, and at the same time generates hydrogen peroxide. The intermediate product hydrogen peroxide is oxidized to oxygen under the action of platinum metal nanoparticles, and two electrons are transferred to the ordered mesoporous carbon. This step reaction of hydrogen peroxide causes a change in the potential of the electrode interface. Therefore, the interface potential of the glucose electrode is correlated with the electrode concentration, and subsequent detection is carried out based on this correlation.

[0108] Example 2

[0109] A biochemical sensor based on an epitaxial gate type field effect transistor (assembled from the fibrous glucose sensitive electrode of Example 1 and the field effect transistor), such as Figure 3 shown, comprising: a substrate 1, a fibrous reference electrode 4, a fibrous glucose sensitive electrode 3, and a field effect transistor 2.

[0110] The field effect transistor 2 includes a drain 21, a gate 22, and a source 23.

[0111] The fibrous reference electrode 4 includes:

[0112] A reference electrode carbon fiber 41 disposed on the substrate 1;

[0113] A silver / silver chloride layer 42 disposed on the surface of the reference electrode carbon fiber;

[0114] A polyvinyl butyral layer 43 containing saturated sodium chloride disposed on the surface of the silver / silver chloride layer 42;

[0115] Silver paste 6 for connecting the reference electrode carbon fiber and the copper wire.

[0116] The fibrous glucose sensitive electrode 3 includes:

[0117] A sensitive electrode carbon fiber 31;

[0118] An ordered mesoporous carbon layer 32 disposed on the surface of the sensitive electrode carbon fiber;

[0119] Platinum nanoparticles 33 and glucose oxidase 32 disposed on the surface of the ordered mesoporous carbon layer 32;

[0120] Silver paste 6 for connecting the sensitive electrode carbon fiber 31 and the copper wire.

[0121] This biochemical sensor further includes:

[0122] A copper electrode 7 pre-exposed on the surface of the substrate 1;

[0123] An encapsulation layer 5 disposed on the substrate 1 but covering the wire portion, and the encapsulation layer 5 exposes the sensitive areas of the fibrous reference electrode 4 and the fibrous glucose sensitive electrode 3, and the material can be selected as polymethyl methacrylate (PMMA).

[0124] Figure 3 The three dashed boxes in [] represent the fibrous reference electrode 4, the fibrous glucose sensitive electrode 3, and the field effect transistor 2 respectively. The fibrous reference electrode 4 provides a stable potential in the solution, the fibrous glucose sensitive electrode 3 provides an interfacial potential related to the glucose concentration, and the field effect transistor 2 converts the change in the interfacial potential into a change in the output drain current.

[0125] Example 3

[0126] A biochemical sensor based on an epitaxial gate type field effect transistor (assembled from a planar glucose sensitive electrode and a field effect transistor), as Figure 4 shown, comprising: a substrate 1, a planar reference electrode 8, a planar glucose sensitive electrode 9, and a field effect transistor 2.

[0127] The field effect transistor 2 includes a drain 21, a gate 22, and a source 23.

[0128] The planar reference electrode 8 includes:

[0129] A planar gold electrode 10 disposed on the substrate 1;

[0130] A silver / silver chloride layer 42 disposed on the planar gold electrode 10;

[0131] A polyvinyl butyral layer 43 containing saturated sodium chloride disposed on the silver / silver chloride layer 42.

[0132] The planar glucose sensitive electrode 9 includes:

[0133] A planar gold electrode 10 disposed on the substrate 1;

[0134] An ordered mesoporous carbon layer 32 disposed on the planar gold electrode 10;

[0135] Platinum nanoparticles 33 and glucose oxidase 34 disposed on the surface of the ordered mesoporous carbon layer 32.

[0136] The biochemical sensor further includes:

[0137] A wire connecting the planar glucose sensitive electrode 42 and the gate 22 of the field effect transistor 2;

[0138] A packaging layer 5 disposed on the substrate 1 but covering the wire portion, and the packaging layer 5 exposes the sensitive regions of the planar reference electrode 8 and the planar glucose sensitive electrode 9, and the material can be selected as polymethyl methacrylate (PMMA).

[0139] Figure 4 The three dashed boxes in

[0140] Example 4

[0141] A biochemical small molecule detection system, such as Figure 5 shown, includes:

[0142] A biochemical sensor based on an epitaxial gate field effect transistor; it mainly includes a substrate 1, a field effect transistor 2, a sensitive electrode 11 and a reference electrode 12;

[0143] A voltage bias and current reading unit 13; the voltage bias and current reading unit is electrically connected to the biochemical sensor for supplying power to the biochemical sensor and reading the output drain current;

[0144] A signal processing unit 14; the signal processing unit is electrically connected to the voltage bias and current reading unit for analyzing the read drain current signal and converting it into the concentration information of the analyte;

[0145] A polymer tank 15 for placing the solution to be detected disposed on the substrate 1, and the polymer tank 15 surrounds the sensitive parts of the sensitive electrode 11 and the reference electrode 12, and the material can be selected as polydimethylsiloxane (PDMS).

[0146] In the actual detection process, the solution to be detected is placed in the polymer tank, and a constant voltage is applied to the reference electrode end, the source end and the drain end of the field effect transistor respectively, and the concentration of the molecule to be detected in the solution is judged according to the magnitude of the output drain current.

[0147] Example 5

[0148] Figure 6 Show a method for biochemical small molecule detection using the detection system of Example 4, including calibration and detection steps, specifically:

[0149] The calibration steps include:

[0150] Step 1: Add a solution with a known concentration of the analyte to the polymer tank, and the solution completely covers the sensitive electrode and the reference electrode;

[0151] Step 2: The voltage bias unit applies a bias voltage to the source-drain end and the reference electrode end of the field effect transistor, and the current reading unit reads the stable value of the drain current;

[0152] Step 3: Aspirate the solution in the polymer tank, wash it with deionized water, replace it with a second solution with a known concentration of the analyte, repeat the operations of the above Step 1 and Step 2, and then replace it with a third solution with a known concentration of the analyte, and repeat the operations of the above Step 1 and Step 2;

[0153] Step 4: Fit the linear relationship between "drain current - analyte concentration" based on the obtained data as the calibration curve.

[0154] The detection steps include:

[0155] Step 1: Add the solution containing the analyte to the polymer cell, and the solution completely covers the sensitive electrode and the reference electrode;

[0156] Step 2: The voltage biasing unit applies a bias voltage to the source-drain terminals and the reference electrode terminal of the field-effect transistor, and the current reading unit reads the stable value of the drain current;

[0157] Step 3: According to the calibration curve obtained from the calibration step, convert the drain current value obtained from the above test into the concentration of the analyte.

[0158] It should be noted that when measuring for the first time, according to the instructions in Figure 6 , it is necessary to first perform the calibration step to obtain the calibration curve, and then perform the detection of the real sample. It is also possible to perform calibration in advance after the device is fabricated, store the calibration curve in the signal analysis unit, and then directly perform the detection after obtaining the sample solution to be detected. This method requires regular updating of the calibration curve to reduce the signal drift of the electrode over time.

[0159] Figure 7 shows the calibration curve for glucose detection in a specific embodiment of the present invention.

[0160] The above embodiments are a specific and preferred implementation manner of the present invention, and the specific features, structures, and materials described therein can be combined in a suitable manner in any implementation manner. Therefore, without departing from the essence and principle of the present invention, those skilled in the art can make appropriate modifications, substitutions, simplifications, and combinations to the embodiments described in this specification, and these should all be included within the protection scope of the present invention.

Claims

1. A biochemical sensor based on an epitaxial gate field effect transistor, characterized in that Comprising: A substrate; A field-effect transistor; The field-effect transistor is fixed on the surface of the substrate and includes a drain, a gate, and a source; A sensitive electrode, which is fixed on the surface of the substrate as an epitaxial gate and is electrically connected to the gate of the field-effect transistor; The sensitive electrode includes: A conductive substrate; An ordered mesoporous carbon layer covering the surface of the conductive substrate; A hydrogen peroxide catalyst loaded in the pores of the ordered mesoporous carbon layer, which is used to catalyze the product hydrogen peroxide of the enzymatic reaction and transfer the electrons after oxidizing hydrogen peroxide to the ordered mesoporous carbon layer and the conductive substrate; An oxidoreductase loaded in the pores of the ordered mesoporous carbon layer, which is used to selectively catalyze the substrate to be hydrogen peroxide to realize the conversion of the detection of various analytes into the detection of hydrogen peroxide; A reference electrode fixed on the surface of the substrate.

2. The biochemical sensor based on an epitaxial gate type field effect transistor according to claim 1, wherein: It further includes a packaging layer that only exposes the sensitive gate and reference electrode regions.

3. The biochemical sensor based on an epitaxial gate type field effect transistor according to claim 1, characterized in that: In the sensitive electrode: The pore diameter of the ordered mesoporous carbon layer is 5 - 20 nm; The particle size of the hydrogen peroxide catalyst is less than 5 nm; The size of the oxidoreductase is less than 20 nm.

4. The biochemical sensor based on an epitaxial gate type field effect transistor according to claim 1, wherein: The conductive substrate is a fibrous conductive substrate or a planar conductive substrate; The fibrous conductive substrate is selected from carbon fiber electrodes and carbon nanotube fiber electrodes; the planar conductive substrate is selected from gold electrodes, silver electrodes, and indium tin oxide electrodes; The hydrogen peroxide catalyst is selected from noble metal nanoparticles and Prussian blue nanoparticles; The oxidoreductase is selected from glucose oxidase that catalyzes glucose, uricase that catalyzes uric acid, cholesterol oxidase that catalyzes cholesterol, and lactate oxidase that catalyzes lactate.

5. The biochemical sensor based on an epitaxial gate type field-effect transistor according to claim 4, wherein: For the fibrous conductive substrate, the ordered mesoporous carbon layer is formed by in-situ growth of an ordered mesoporous carbon material on the surface of the fibrous conductive substrate through a hydrothermal reaction and then high-temperature calcination; For the planar conductive substrate, the ordered mesoporous carbon layer is formed by drop-coating an ordered mesoporous carbon material suspension on the planar conductive substrate.

6. The biochemical sensor based on an epitaxial gate type field effect transistor according to claim 5, characterized in that: The preparation method of the sensitive electrode includes the following steps: Step 1, modifying an ordered mesoporous carbon layer on the surface of the conductive substrate; Step 2, modifying a hydrogen peroxide catalyst on the surface of the ordered mesoporous carbon layer; Step 3, continuously modifying an oxidoreductase on the surface of the ordered mesoporous carbon layer modified with the hydrogen peroxide catalyst; the modification method of the oxidoreductase is physical adsorption using the confinement effect of the ordered mesoporous carbon layer or chemical adsorption between the oxidoreductase and the ordered mesoporous carbon layer.

7. The biochemical sensor based on an epitaxial-gate field-effect transistor according to claim 6, characterized in that: In step 1: For the fibrous conductive substrate, directly modify and grow a single micelle containing an organic small molecule on the surface of the fibrous conductive substrate through interfacial assembly, and form an ordered mesoporous carbon layer after calcination; For a planar conductive substrate, a pre-prepared suspension of ordered mesoporous carbon material is directly dropped and covered on the surface of the planar conductive substrate, and an ordered mesoporous carbon layer is formed after drying, with a surface mass density of 1-3 μg / mm 2 .

8. The biochemical sensor based on an epitaxial gate type field effect transistor according to claim 7, wherein: For the sensitive electrode using a fibrous conductive substrate, its preparation method includes the following steps: Step A: Prepare a low-order phenolic resin solution and a triblock copolymer solution respectively, then mix the low-order phenolic resin solution and the triblock copolymer solution, heat and stir to obtain a spherical micelle solution; Step B: Take the above spherical micelle solution, dilute it with water, and at the same time immerse the fibrous conductive substrate in the obtained solution, and carry out a hydrothermal reaction to obtain a fibrous conductive substrate with ordered micelles grown on the surface; Step C: Calcinate the fibrous conductive substrate with ordered micelles grown on the surface in an inert gas atmosphere to obtain a fibrous conductive substrate with an ordered mesoporous carbon layer modified on the surface; Step D: Prepare a precursor solution of a hydrogen peroxide catalyst, immerse the fibrous conductive substrate with an ordered mesoporous carbon layer modified on the surface in the precursor solution, and then add a reducing agent. After the reduction reaction, a fibrous conductive substrate loaded with a hydrogen peroxide catalyst is obtained; Step E: Immerse the fibrous conductive substrate loaded with a hydrogen peroxide catalyst in an oxidoreductase solution, let it stand, and take it out to prepare a sensitive electrode.

9. A biochemical small molecule detection system, characterized in that Comprising: The biochemical sensor based on an epitaxial gate type field effect transistor according to any one of claims 1-8; A voltage bias and current reading unit; the voltage bias and current reading unit is electrically connected to the biochemical sensor for supplying power to the biochemical sensor and reading the output drain current; A signal processing unit; The signal processing unit is electrically connected to the voltage bias and current reading unit for analyzing the read drain current signal and converting it into the concentration information of the analyte.

10. A method for detecting biochemical small molecules, characterized in that: Use the biochemical small molecule detection system according to claim 9 to detect the analyte; when detecting, immerse the sensitive electrode and the reference electrode in the solution to be detected, set a constant bias voltage at the reference electrode end to control the gate voltage of the field effect transistor, and at the same time set a constant source-drain voltage; since the interfacial potential between the sensitive electrode and the solution to be detected depends on the concentration of the analyte in the solution to be detected, this interfacial potential affects the gate voltage of the field effect transistor, thereby changing the output drain current; and then detect the analyte according to the relationship between the drain current and the concentration of the analyte.

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