Electrochemical-field effect transistor dual-mode biosensor and its preparation method and application

Abstract

The application provides an electrochemical-field effect transistor dual-mode biosensor, which comprises, from bottom to top, an insulating substrate, a carbon nanotube layer, a source electrode and a drain electrode, a gate medium layer, an extended gate electrode layer and a passivation layer; the extended gate electrode layer comprises a channel gate metal electrode arranged in a channel layer region and an electrode sensitive region arranged away from the channel layer region, the channel gate metal electrode and the electrode sensitive region are connected, and a biological probe is arranged on the surface of the electrode sensitive region; the sensor has both EC sensing mode and FET sensing mode detection systems by adding a test solution and an external electrode; the two detection systems can be independently detected, and the detection ranges are complementary, so that direct detection of targets in a large concentration span range is realized; meanwhile, the same sample is detected by using the two detection systems to verify each other, and the accuracy and reliability of the detection sample, especially the complex sample matrix result, are improved.

Description

Technical Field

[0001] The present invention relates to the field of biosensor technology, and in particular to an electrochemical-field effect transistor dual-mode biosensor and its preparation method and application. Background Art

[0002] Biosensors are cutting-edge innovative technologies that fuse biology and electronics. They can detect and quantify a wide range of biomolecules and chemicals, playing an irreplaceable and important role in key areas such as medical diagnosis, environmental monitoring, food safety, and biotechnology research. They possess extremely broad development prospects and profound social significance. However, the detection capabilities of biosensors face complex challenges due to the diverse types of biomolecules and their extremely wide concentration distribution across different samples.

[0003] The concentration of biomolecules in different samples may extend from the picomolar (pM) level to the millimolar (mM) level, spanning up to nine orders of magnitude. Such a concentration span places extremely high demands on the detection sensitivity and dynamic range of biosensors. Traditional biosensors usually rely on a single physical or chemical signal conversion pathway, such as electrochemical, optical, acoustic, thermal or magnetic signals. Although this single mechanism can provide effective detection in specific application scenarios, it has significant limitations when faced with a wide concentration range and complex sample environment.

[0004] Single-signal response mechanisms often maintain high sensitivity and linearity only within a narrow concentration range. When the concentration of the target biomolecule exceeds this range, the sensor's response may saturate or become unstable, rendering it ineffective for quantifying the target molecule. Furthermore, in complex sample matrices, background noise and nonspecific interferences can mask the target molecule's signal, further reducing detection accuracy and reliability. Summary of the Invention

[0005] The present invention aims to address the problems that existing biosensors cannot effectively quantify target molecules when detecting samples with a wide concentration range on the same platform, or the poor accuracy and reliability of detection results when detecting samples with complex components. The present invention provides a dual-mode biosensor with complementary advantages. The dual-mode biosensor combines the electrochemical (EC) sensing mode and the field-effect transistor (FET) sensing mode with two different signal transduction mechanisms. It can independently collect signals through different test structure configurations on the same platform to meet the needs of high sensitivity, wide dynamic range and high specificity detection in practical applications.

[0006] According to a first aspect of the present invention, there is provided an electrochemical-field effect transistor dual-mode biosensor, comprising an insulating substrate, a carbon nanotube layer, a source electrode and a drain electrode, a gate dielectric layer, an extended gate electrode layer, and a passivation layer stacked sequentially from bottom to top;

[0007] The carbon nanotube layer is located on one surface of the insulating substrate;

[0008] The source electrode and the drain electrode are located above the carbon nanotube layer, the source electrode and the drain electrode are arranged opposite to each other and spaced apart, and the carbon nanotube layer between the source electrode and the drain electrode forms a channel layer;

[0009] The gate dielectric layer is located above the channel layer and covers the entire channel layer;

[0010] The extended gate electrode layer includes a channel gate metal electrode and an electrode sensitive region, wherein the channel gate metal electrode is located in the channel layer region between the source electrode and the drain electrode and is disposed above the gate dielectric layer;

[0011] The electrode sensitive region is located at one end of the insulating substrate away from the channel layer region and is connected to the channel gate metal electrode via a metal lead; the surface of the electrode sensitive region is modified with a biological probe for specifically capturing target molecules;

[0012] The passivation layer is configured to cover all areas except the electrode sensitive area and the output electrode (pad), so that during the target molecule incubation, washing, and detection processes of the biosensor, only the electrode sensitive area is exposed to the solution environment;

[0013] A test electrolyte and an external electrode are provided on the surface of the electrode sensitive area, and the external electrode is in contact with the test electrolyte;

[0014] Among them, the external electrode is used as the reference electrode, the gate electrode is used as the working electrode, and the reference electrode and the working electrode form an electrochemical (EC) sensing system through the test electrolyte;

[0015] The external electrode is used as the gate to apply gate voltage, and the transfer characteristic curve between the source electrode and the drain electrode is recorded to form a field effect transistor (FET) sensing system.

[0016] As an optional embodiment, the test electrolyte is a PBS solution containing 10 μM [Ru(NH 3 ) 6 ]Cl 3 .

[0017] As an optional implementation, the external electrode is an Ag wire or an Ag / AgCl wire.

[0018] As an optional embodiment, the gate electrode is a Ti / Au or Cr / Au film.

[0019] As an optional embodiment, the carbon nanotube layer is a semiconductor single-walled carbon nanotube film.

[0020] As an optional embodiment, the source electrode and the drain electrode are Ti / Pd / Au films.

[0021] As an optional implementation, the gate dielectric layer is a composite film of Y2O3 and HfO2.

[0022] As an optional embodiment, the source electrode, the drain electrode and the extended gate electrode are all configured with independent leads and are connected to the respective output electrodes through the leads.

[0023] According to a second aspect of the present invention, there is provided a method for preparing the aforementioned electrochemical-field effect transistor dual-mode biosensor, comprising the following steps:

[0024] S1: depositing a carbon nanotube layer on the surface of an insulating substrate;

[0025] S2: Applying a photoresist on the surface of the carbon nanotube layer, etching the corresponding pattern using a photolithography process, and evaporating metal to form source and drain electrodes, as well as corresponding leads and output electrodes, and then removing the photoresist and peeling off;

[0026] S3: Spread the film again, use oxygen plasma etching to remove excess carbon nanotubes, and remove the film to obtain the channel area connected to the source and drain electrodes;

[0027] S4: evaporating metallic yttrium on the entire surface of the device obtained in step S3 and oxidizing it to form Y2O3, and then depositing HfO2 on the Y2O3 by an atomic layer deposition system;

[0028] S5: Spreading and exposure are performed again, and metal is evaporated on the channel area after step S4 to form a channel gate metal electrode. Metal is also evaporated on one end of the insulating substrate away from the channel layer area to form an electrode sensitive area, which serves as the modification area of ​​the biological probe. The channel gate metal electrode and the electrode sensitive area are connected by metal wires;

[0029] S6: encapsulating all areas except the electrode sensitive area and the output electrode (pad) using photolithography and development technology again to form a passivation layer, so that during the target molecule incubation, washing, and detection processes of the biosensor, only the electrode sensitive area is exposed to the solution environment;

[0030] S7: Modify the desired biological probe on the surface of the electrode sensitive area to obtain the desired electrochemical-field effect transistor dual-mode biosensor.

[0031] According to a third aspect of the present invention, there is provided a use of the aforementioned electrochemical-field effect transistor dual-mode biosensor in detecting samples with a wide concentration range and complex components.

[0032] As can be seen from the technical solutions of the present invention described above, the electrochemical-field effect transistor dual-mode biosensor proposed in the present invention has two detection systems: an electrochemical (EC) sensing mode and a field effect transistor (FET) sensing mode. Both detection systems can perform independent detection. The electrochemical (EC) sensing mode can be used to detect samples with a high concentration detection range, and the field effect transistor (FET) sensing mode can be used to detect samples with a low concentration range. The detection ranges of the two detection modes complement each other, thereby achieving direct detection of targets within a large concentration span.

[0033] At the same time, the electrochemical (EC) sensing mode and the field effect transistor (FET) sensing mode can be used to detect the same sample respectively to verify each other and improve the accuracy and reliability of the detection results of samples, especially complex sample matrices.

[0034] During testing of the electrochemical-field effect transistor dual-mode biosensor of the present invention, the sensing area modified with the biorecognition probe is the only part of the device exposed to the liquid environment, effectively avoiding direct contact between the complex solution and the sensor channel surface, ensuring its stability and repeatability.

[0035] The electrochemical-field effect transistor dual-mode biosensor of the present invention has simple operation, no labeling, rapid detection, low detection limit and high sensitivity. BRIEF DESCRIPTION OF THE DRAWINGS

[0036] Figure 1 A cross-sectional view of the structure of the electrochemical-field effect transistor dual-mode biosensor of the present invention.

[0037] Figure 2 It is a partial structural schematic diagram of the electrochemical-field effect transistor dual-mode biosensor of the present invention.

[0038] Figure 3 4 is a top view of the electrochemical-field effect transistor dual-mode biosensor of the present invention.

[0039] Figure 4 Schematic diagram of the test structure configuration and signal response of the EC sensing mode of the electrochemical-field effect transistor dual-mode biosensor of the present invention.

[0040] Figure 5 It is a schematic diagram of the test structure configuration and signal response of the FET sensing mode of the electrochemical-field effect transistor dual-mode biosensor of the present invention.

[0041] Figure 6 The differential pulse voltammetry curves (A) recorded in the EC sensing mode at different IFN-γ concentrations in Example 2 of the present invention and the linear fitting curve (B) of the peak current increment and the logarithmic concentration of IFN-γ.

[0042] Figure 7 The transfer characteristic curve (A) and the threshold voltage shift value ΔV recorded in the FET sensing mode under different concentrations of IFN-γ in Example 3 of the present invention are shown in FIG. th Linear fitting curve of IFN-γ logarithmic concentration (B).

[0043] Figure 8 This is the response of different cytokines in Example 4 of the present invention under two modes.

[0044] Explanation of the accompanying reference numerals: 10, insulating substrate; 20, carbon nanotube layer; 21, channel layer; 30, source electrode; 40, drain electrode; 50, gate dielectric layer; 50-1, metal oxide insulating dielectric layer; 60, extended gate electrode; 61, channel gate metal electrode; 62, electrode sensitive area; 62-1, biological probe; 63, metal lead; 70, passivation layer; 80, test electrolyte; 90, external electrode. DETAILED DESCRIPTION

[0045] In order to better understand the technical content of the present invention, specific embodiments are given and described below with reference to the accompanying drawings.

[0046] Various aspects of the present invention are described in this disclosure with reference to the accompanying drawings, in which a number of illustrative embodiments are shown. The embodiments of the present disclosure are not necessarily intended to be comprehensive. It should be understood that the various concepts and embodiments described above, as well as those described in more detail below, can be implemented in any of a number of ways.

[0047] Electrochemical-field-effect transistor dual-mode biosensor

[0048] Combine Figure 1-3 The electrochemical-field effect transistor dual-mode biosensor of the illustrated example includes an insulating substrate 10, a carbon nanotube layer 20, a source electrode 30 and a drain electrode 40, a gate dielectric layer 50, an extended gate electrode 60 and a passivation layer 70 stacked sequentially from bottom to top.

[0049] like Figure 1 As shown, the carbon nanotube layer 20 is located on one surface of the insulating substrate 10 ; for ease of description, the surface of the carbon nanotube layer away from the insulating substrate is defined as the upper side of the carbon nanotube layer.

[0050] As an optional example, the insulating substrate can be obtained by coating a PI film on a silicon wafer.

[0051] Combine Figure 1 and Figure 2 In the example of FIG. 1 , the source electrode 30 and the drain electrode 40 are located above the carbon nanotube layer 20 . The source electrode 30 and the drain electrode 40 are both configured with independent leads and are connected to their respective output electrodes through the leads.

[0052] The source electrode 30 and the drain electrode 40 are disposed opposite to each other and spaced apart. The carbon nanotube layer between the source electrode 30 and the drain electrode 40 forms a channel layer 21 . The carbon nanotubes in other areas of the insulating substrate 10 are removed by etching.

[0053] In the embodiment of the present invention, the carbon nanotube layer is a network-shaped semiconductor single-walled carbon nanotube film, and the formed channel layer 21 has a width of 600 μm and a length of 40-48 μm.

[0054] As an optional example, the source electrode and the drain electrode are strip-shaped electrodes obtained by sequentially depositing 0.6 nm of Ti, 20 nm of Pd, and 60 nm of Au, with a length of 600 μm.

[0055] like Figure 1 、 Figure 3 As shown in FIG. 2 , the gate dielectric layer 50 is located above the channel layer 21 and covers the entire channel layer 21 .

[0056] In an optional example, the gate dielectric layer 50 is a high-к gate dielectric, which covers the channel layer 21 and improves the control capability of the channel gate metal electrode over the conductive channel.

[0057] In another optional example, the gate dielectric layer 50 is a composite film of Y 2 O 3 and HfO 2 , wherein the thickness of Y 2 O 3 is 6 nm, and the thickness of HfO 2 is 6 nm.

[0058] In one example, when preparing the gate dielectric layer, the dielectric can be covered on the entire device. The dielectric layer on the channel layer is the gate dielectric layer, and the dielectric layer in other areas, namely the surface of the source and drain electrodes, and above the insulating substrate 10, and avoiding the dielectric layer covering the area where the channel layer 21, the source electrode 30 and the drain electrode 40 are located, is defined as a metal oxide insulating dielectric layer 50-1.

[0059] It is understandable that in the example of the present invention, the metal oxide insulating dielectric layer 50 - 1 is retained to save steps. Alternatively, the metal oxide insulating dielectric layer 50 - 1 may be removed by etching it before preparing the next step of the extended gate electrode.

[0060] Combine Figure 1 、 Figure 2 As shown, the extended gate electrode 60 includes a channel gate metal electrode 61 and an electrode sensitive area 62. The channel gate metal electrode 61 is located in the channel layer area between the source electrode 30 and the drain electrode 40, and is arranged above the gate dielectric layer 50. The channel gate metal electrode 61 is connected to its separately configured output electrode through an independent lead.

[0061] In the example of the present invention, the channel gate metal electrode 61 is a long strip electrode with a length of 600 μm, a width of 44 μm, and a thickness of 60 nm. The spacing between the gate source and the gate drain electrodes is 2-4 μm.

[0062] As an optional example, the gate electrode is made of Ti / Au or Cr / Au. In one preferred example, the gate electrode is formed by sequentially depositing 0.6 nm of Ti and 60 nm of Au.

[0063] Combine Figure 2 、 Figure 3 As shown, the electrode sensitive region 62 is located at one end of the insulating substrate away from the channel layer region, and is connected to the channel gate metal electrode 61 through a metal lead 63 .

[0064] In the example of the present invention, the electrode sensitive area 62 can be set to any shape, and the area size is not specifically limited. It can be set according to actual needs. For example, when a pipette is used to drop 5 μL of test liquid on the electrode sensitive area, the radius of the droplet is about 1250 μm. In preparation, the electrode sensitive area can be set to a circle with a radius of 1250 μm.

[0065] The surface of the electrode sensitive area 62 is modified with a biological probe 62-1 for specifically capturing the target molecule. It is understandable that a corresponding aptamer probe can be set as the biological probe according to the target molecule.

[0066] The passivation layer 70 is configured to cover all areas except the electrode sensitive area and the output electrode (pad), so that during the target molecule incubation, washing and detection processes of the biosensor, only the electrode sensitive area is exposed to the solution environment.

[0067] In an example of the present invention, the passivation layer is prepared using 1813 photoresist.

[0068] like Figure 3 As shown, a test electrolyte 80 and an external electrode 90 are provided on a surface of the electrode sensitive region 62 away from the third gate dielectric layer 53 , and the external electrode 90 is in contact with the test electrolyte 80 .

[0069] The biosensor exemplified in the present invention has two detection systems: electrochemical (EC) sensing mode and transistor (FET) sensing mode. Both detection systems can perform independent detection. The test method is as follows:

[0070] {Electrochemical (EC) Sensing Mode}

[0071] like Figure 4As shown, the test structure configuration of the EC sensing mode is to use the external electrode 90 as the reference electrode, the gate electrode 60 as the working electrode, and the reference electrode and the working electrode form a two-electrode detection system through the test electrolyte 80, and record the current signal of the two-electrode system composed of the reference electrode, the working electrode and the test electrolyte.

[0072] The signal response of the EC sensing mode comes from the specific binding of the biological probe to the target molecule, which changes the difficulty of electron exchange between the electroactive molecules in the electrolyte solution and the electrode, thereby changing the redox peak current of the electroactive molecules.

[0073] {Transistor (FET) sensing mode}

[0074] like Figure 5 As shown, the FET sensing mode test structure configuration is to use the external electrode 90 as the gate to apply a gate voltage and record the transfer characteristic curve between the source electrode 30 and the drain electrode 40.

[0075] The signal response of FET sensing mode originates from the changes in gate surface potential and interface capacitance caused by the specific binding of biological probes to target molecules, thereby changing the electrical characteristics of the transistor.

[0076] As an optional example, the test electrolyte is a PBS solution containing 10 μM [Ru(NH 3 ) 6 ]Cl 3 .

[0077] It can be understood that the PBS solution is a standard solution and can be purchased directly, specifically a PBS solution with a pH of 7.2 and 0.1 mM.

[0078] As an optional example, the external electrode 90 is an Ag wire or an Ag / AgCl wire.

[0079] It can be understood that the respective output electrodes corresponding to the source electrode, the drain electrode and the extended gate electrode, as well as the connected leads, are made of the same material as the respective electrodes.

[0080] Preparation method of electrochemical-field effect transistor dual-mode biosensor

[0081] In another exemplary embodiment of the present invention, a method for preparing the aforementioned electrochemical-field effect transistor dual-mode biosensor is provided, comprising the following steps:

[0082] S1: Using a polyamic acid solution (pyromellitic dianhydride-co-4,4'-diaminodiphenyl ether), a film is formed on the surface of a silicon wafer by spin coating and thermal curing to serve as an insulating substrate. The insulating substrate is then immersed in a culture dish containing a carbon nanotube dispersion solution to deposit a carbon nanotube layer on the surface of the insulating substrate.

[0083] S2: Applying a photoresist on the surface of the carbon nanotube layer, etching the corresponding pattern using a photolithography process, and evaporating metal to form source and drain electrodes, as well as corresponding leads and output electrodes, and then removing the photoresist and peeling off;

[0084] S3: Spread the film again, use oxygen plasma etching to remove excess carbon nanotubes, and remove the film to obtain the channel area connected to the source and drain electrodes;

[0085] S4: evaporating metallic yttrium on the entire surface of the device obtained in step S3 and oxidizing it to form Y2O3, and then depositing HfO2 on the Y2O3 by an atomic layer deposition system;

[0086] S5: Spreading and exposure are performed again, and metal is evaporated on the channel area after step S4 to form a channel gate metal electrode. Metal is also evaporated on one end of the insulating substrate away from the channel layer area to form an electrode sensitive area, which serves as the modification area of ​​the biological probe. The channel gate metal electrode and the electrode sensitive area are connected by metal wires;

[0087] S6: encapsulating all areas except the electrode sensitive area and the output electrode (pad) using photolithography and development technology again to form a passivation layer, so that during the target molecule incubation, washing, and detection processes of the biosensor, only the electrode sensitive area is exposed to the solution environment;

[0088] S7: Modify the desired biological probe on the surface of the electrode sensitive area to obtain the desired electrochemical-field effect transistor dual-mode biosensor.

[0089] In another exemplary embodiment of the present invention, there is provided an application of the aforementioned electrochemical-field effect transistor dual-mode biosensor in detecting samples with a wide concentration range and complex components.

[0090] The electrochemical (EC) sensing mode can be used to detect samples with a high concentration detection range, and the field effect transistor (FET) sensing mode can be used to detect samples with a low concentration range. The detection ranges of the two detection modes are complementary, thus enabling direct detection of targets within a large concentration span.

[0091] The electrochemical (EC) sensing mode and the field effect transistor (FET) sensing mode were used to detect the same sample respectively to verify each other and improve the accuracy and reliability of the detection results, especially for complex sample matrices.

[0092] It is understandable that the complex sample in the present invention refers to a sample to be tested that contains many interfering substances in addition to the target molecule, so that background noise and nonspecific interference signals exist during the test, resulting in the masking of the signal of the target molecule.

[0093] For better understanding, the present invention is further described below with reference to several specific examples, but the preparation process is not limited thereto, and the content of the present invention is not limited thereto.

[0094] Unless otherwise specified, the materials in the examples were prepared according to existing methods or purchased directly from the market.

[0095] Example 1

[0096] Preparation of electrochemical-field effect transistor dual-mode biosensor

[0097] (1) Using polyamic acid solution (pyromellitic dianhydride-co-4,4′-diaminodiphenyl ether), a clean silicon wafer (2 cm × 2 cm) was placed in a sizing machine. 2 mL of polyamic acid solution was dripped onto the silicon wafer and allowed to stand for 10 seconds to avoid uneven solution spreading due to solution viscosity.

[0098] The coating machine was controlled to perform two steps of spin coating, wherein the first step was a rotation speed of 1000 rpm / min for 10 seconds, and the second step was a rotation speed of 6000 rpm / min for 60 seconds.

[0099] The silicon wafer with the polyamic acid solution spin-coated was heated at 200°C for 1 hour on a hot plate, and then heated to 350°C for 2 hours to complete the imidization process of the solution to form a polyimide (PI) film as an insulating substrate.

[0100] (2) The silicon wafer prepared with the PI film was immersed in a dispersion containing high-purity carbon tubes for 2 hours to obtain a uniform and dense network of carbon tube films.

[0101] (3) Photolithography, evaporation of source and drain electrodes, and stripping to form the required pattern. The source and drain electrodes are composed of 0.6nm Ti, 20nm Pd, and 60nm Au deposited in sequence.

[0102] (4) Photolithography: Use oxygen plasma etching for 60s to remove excess carbon nanotubes and perform debonding and peeling.

[0103] (5) 3 nm of metallic yttrium was evaporated in the channel area of ​​the device and baked on a hot plate at 270 °C for 30 min for oxidation. After oxidation, Y2O3 was about 6 nm. Then, about 6 nm of HfO2 was deposited on the Y2O3-coated carbon nanotubes by an atomic layer deposition (ALD) system.

[0104] (6) Photolithography, evaporation of gate electrode, and stripping to form the required pattern; the gate electrode is composed of 0.6nm Ti and 60nm Au deposited in sequence.

[0105] (7) Photolithography: Use 1813 photoresist to isolate the electrodes and expose the source, drain, gate pad and extended gate area for biosensing detection.

[0106] (8) Surface modification of the electrode sensitive region with biological probes: In order to block the excess probe binding sites, 6-Hydroxy-1-hexanethiol (MCH) was diluted to 2 mM and used immediately, and then dropped onto the surface of the electrode sensitive region for 30 minutes, followed by deionized water cleaning and nitrogen gun drying to reduce non-specific adsorption.

[0107] The probe solution for surface modification of the electrode sensitive region was composed of IFN-γ aptamer probes with 5-end modification (5'-SH C6) and Tris-(2-carboxyethyl)-phosphine hydrochloride (TCEP) for activating the probes. The sequence of the aptamer probe was: 5'-SH C6 / GGG GTT GGT TGT GTT GGG TGTTGT GT.

[0108] The probe solution was prepared by mixing 10 μL of aptamer with a concentration of 20 μM and 10 μL of TCEP with a concentration of 10 mM, and then adding 180 μL of PBS to dilute to 200 μL. The mixed solution should be used after waiting for 1 hour after preparation.

[0109] The mixed solution was dropped onto the surface of the electrode sensitive region, and the probe solution was incubated in a humid, closed, and low-temperature environment at 4°C for 4 hours. Then, the biosensor for detecting the cytokine interferon-γ (IFN-γ) was obtained by using ultrapure water to rinse and nitrogen to dry.

[0110] During the test, 5 μL of PBS test solution (pH = 7.2, 0.1 mM) containing 10 μM [Ru(NH3)6]Cl3 was dropped onto the surface of the electrode sensitive region, and one end of the external electrode was placed in the test solution.

[0111] The size of the biosensor in this embodiment is as follows:

[0112] Source electrode: 600 μm long, 50 μm wide, and 60.6 nm thick

[0113] Drain electrode: 600 μm long, 50 μm wide, and 60.6 nm thick

[0114] Source electrode PAD: 800 μm long, 295 μm wide, and 60.6 nm thick

[0115] Drain electrode PAD: 800 μm long, 295 μm wide, and 60.6 nm thick

[0116] Channel layer region: channel length of 48 μm and channel width of 600 μm

[0117] Channel gate metal electrode: 600 μm long, 44 μm wide, 60 nm thick, the distance between the gate source and the gate drain electrode is 2 μm

[0118] Electrode sensitive area: radius 1250 μm, thickness 60 nm.

[0119] Example 2

[0120] {EC sensing mode for cytokine IFN-γ analysis}

[0121] Prepare IFN-γ solutions with different concentrations (0, 0.1, 0.4, 4, 50, 200, 800, 2000, 5000 ng / mL)

[0122] Use a pipette to drop 5 μL of the IFN-γ solution onto the surface of the electrode sensitive area of the sensor (prepared according to the method of Example 1) and incubate for 7 min; after rinsing with pure water, drop 5 μL of a PBS test solution (pH = 7.2, 0.1 mM) containing 10 μM [Ru(NH3)6]Cl3 onto the surface of the electrode sensitive area, and place one end of the Ag / AgCl wire external electrode in the test solution, and perform differential pulse voltammetry testing (0 V to -0.65 V) using the structure as shown in Figure 4 . Each concentration is tested according to the above method, and the results are shown in Figure 6 .

[0123] As can be seen from the figure, when the concentration of IFN-γ is less than 0.1 ng / mL, there is essentially no significant signal change; as the concentration of IFN-γ increases, the reduction peak current of the device increases sharply, and when it reaches 800 ng / mL, the response signal gradually tends to be flat. In the concentration range of 0.4 ng / mL to 2000 ng / mL, the peak current increment (ΔI) has a good linear relationship with the logarithm of the concentration of IFN-γ (C [IFN-γ] ) ( Figure 6 B ), and the linear regression equation is ΔI = 0.0659 × lgC [IFN-γ] + 0.0597 (ng / mL, R 2 = 0.9926).

[0124] The test data results show that the binding between the aptamer and the IFN-γ on the surface of the electrode sensitive area follows the Langmuir adsorption model, and the dissociation constant Kd is about 5.51 nM, proving that the EC sensing mode of the sensor of the present application can effectively bind the aptamer and the IFN-γ in the detection range of 0.4-2000 ng / mL, and the EC sensing mode has high reliability and can detect samples in the high concentration (0.4-2000 ng / mL) detection range.

[0125] Example 3

[0126] {FET sensing mode for cytokine IFN-γ analysis}

[0127] Prepare IFN-γ solutions of different concentrations (0, 0.02, 0.1, 1, 20, 200, 1000, 4000, 20000, 40000 pg / mL)

[0128] 5 μL of IFN-γ solution was pipetted onto the surface of the electrode sensitive area of ​​the sensor (prepared according to the method of Example 1) and allowed to stand for 7 minutes for incubation. After rinsing with pure water, 5 μL of PBS test solution (pH = 7.2, 0.1 mM) containing 10 μM [Ru(NH3)6]Cl3 was added onto the surface of the electrode sensitive area, and one end of the external electrode using Ag / AgCl wire was placed in the test solution. Figure 5 The transfer characteristic curve test was performed using the structure shown in the figure. When the transistor was given Vds = -0.1V, a linear voltage sweep was performed to obtain the transfer characteristic curve, and the test window was fixed between Vg = -1.2V and 1V. Each concentration was tested according to the above method, and the results are shown in the figure below. Figure 7 As shown in A.

[0129] As can be seen from the figure, as the concentration of interferon-γ increases, the transfer characteristic curve of the device shifts parallel to the left. In the concentration range of 0.1pg / mL to 4000pg / mL, the threshold voltage shift value (ΔV th ) and C [IFN-γ] The logarithm of Figure 7 B), its linear regression equation is ΔV th (V) = 0.0340 × 1gC [IFN-γ] +0.0788(pg / mL,R 2 =0.997, n=3), the calculated maximum standard deviation was 8.1%, and the detection limit was 24.1 fg / mL.

[0130] The test shows that the FET sensing mode of the sensor of the present invention can detect samples in a low concentration range (0.1-4000 pg / mL) with a detection limit of 24.1 fg / mL.

[0131] Combining Example 2 and Example 3, it can be demonstrated that the two detection modes of the sensor of the present invention have complementary detection ranges and can verify each other, and can achieve direct detection of targets within a wide concentration range on the same platform.

[0132] Example 4

[0133] {Detection of response to distractors}

[0134] In order to verify that the device does not respond to other interfering substances, the responses of EC and FET to some inflammatory cytokines closely related to IFN-γ, including interleukin-6 (IL-6), interleukin-2 (IL-2) and tumor necrosis factor-α (TNF-α), were measured under the same conditions as in Examples 3 and 4. The results are shown in Figure 2. Figure 8 shown.

[0135] Figure 8 The results show that the degree and trend of the impact of interfering species on the measurement signals of the two sensing modes are different. For interfering substances with a concentration 5 times higher than the target, the normalized signal (defined as the ratio of the response to the interfering substance to the response to the target) ranges from -5.4% to 8.5%. The results show that the response of EC and FET to the target is more than 10 times greater than the response to the interfering substance. When the target is mixed with other interfering substances, the normalized signals of EC and FET remain at 94% and 108%, respectively, indicating that the sensor of the present invention has good anti-interference ability.

[0136] The above demonstrates that the EC and FET dual-mode biosensor of the present invention can selectively detect IFN-γ with negligible interference from other inflammatory cytokines. Specifically, the sensor of the present invention can detect the same sample in both the EC and FET sensing modes, enabling mutual verification and improving the accuracy and reliability of detection results, particularly in complex sample matrices.

[0137] While the present invention has been disclosed above with reference to preferred embodiments, this is not intended to limit the present invention. Persons skilled in the art will readily appreciate that various modifications and variations can be made without departing from the spirit and scope of the present invention. Therefore, the scope of protection of the present invention shall be determined by the claims.

Claims

1. An electrochemical-field effect transistor dual-mode biosensor, characterized in that: The device comprises an insulating substrate (10), a carbon nanotube layer (20), a source electrode (30), a drain electrode (40), a gate dielectric layer (50), a gate electrode (60) layer and a passivation layer (70) stacked in sequence from bottom to top; The carbon nanotube layer (20) is located on one side surface of the insulating substrate (10); The source electrode (30) and the drain electrode (40) are located above the carbon nanotube layer (20), the source electrode (30) and the drain electrode (40) are arranged opposite to each other and spaced apart, and the carbon nanotube layer (20) between the source electrode (30) and the drain electrode (40) forms a channel layer (21); The gate dielectric layer (50) is located above the channel layer (21) and covers the entire channel layer (21); The extended gate electrode (60) layer includes a channel gate metal electrode (61) and an electrode sensitive region (62), wherein the channel gate metal electrode (61) is located in the channel layer (21) region between the source electrode (30) and the drain electrode (40), and is disposed above the gate dielectric layer (50); The electrode sensitive region (62) is located at one end of the insulating substrate (10) away from the channel layer (21) and is connected to the channel gate metal electrode (61) via a metal lead (63); the surface of the electrode sensitive region (62) is modified with a biological probe (62-1) for specifically capturing target molecules; The passivation layer (70) is arranged to cover all areas except the electrode sensitive area (62) and the output electrode (pad), so that during the target molecule incubation, washing and detection process of the biosensor, only the electrode sensitive area (62) is exposed to the solution environment; A test electrolyte (80) and an external electrode (90) are provided on the surface of the electrode sensitive area (62), and the external electrode (90) is in contact with the test electrolyte (80); The external electrode (90) is used as a reference electrode, the gate electrode (60) is used as a working electrode, and the reference electrode and the working electrode form an electrochemical (EC) sensing system through a test electrolyte (80); The external electrode (90) is used as a gate to apply a gate voltage, and the transfer characteristic curve between the source electrode (30) and the drain electrode (40) is recorded to form a field effect transistor (FET) sensing system.

2. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The test electrolyte (80) is a PBS solution containing 10 μM [Ru(NH3)6]Cl3.

3. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The external electrode (90) is an Ag wire or an Ag / AgCl wire.

4. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The gate electrode (60) is a Ti / Au or Cr / Au film.

5. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The carbon nanotube layer (20) is a semiconductor single-walled carbon nanotube film.

6. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The source electrode (30) and the drain electrode (40) are Ti / Pd / Au films.

7. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The gate dielectric layer (50) is a composite film of Y2O3 and HfO2.

8. The electrochemical-field effect transistor dual-mode biosensor according to claim 1, characterized in that: The source electrode (30), the drain electrode (40) and the extended gate electrode (60) are all provided with independent leads and are connected to respective output electrodes via the leads.

9. A method for preparing an electrochemical-field effect transistor dual-mode biosensor according to any one of claims 1 to 8, characterized in that: The following steps are involved: S1: depositing a carbon nanotube layer on the surface of an insulating substrate; S2: Applying a photoresist on the surface of the carbon nanotube layer, etching the corresponding pattern using a photolithography process, and evaporating metal to form source and drain electrodes, as well as corresponding leads and output electrodes, and then removing the photoresist and peeling off; S3: Spread the film again, use oxygen plasma etching to remove excess carbon nanotubes, and remove the film to obtain the channel area connected to the source and drain electrodes; S4: evaporating metallic yttrium on the entire surface of the device obtained in step S3 and oxidizing it to form Y2O3, and then depositing HfO2 on the Y2O3 by an atomic layer deposition system; S5: Spreading and exposure are performed again, and metal is evaporated on the channel area after step S4 to form a channel gate metal electrode. Metal is also evaporated on one end of the insulating substrate away from the channel layer area to form an electrode sensitive area, which serves as the modification area of ​​the biological probe. The channel gate metal electrode and the electrode sensitive area are connected by metal wires; S6: encapsulating all areas except the electrode sensitive area and the output electrode (pad) using photolithography and development technology again to form a passivation layer, so that during the target molecule incubation, washing, and detection processes of the biosensor, only the electrode sensitive area is exposed to the solution environment; S7: Modify the desired biological probe on the surface of the electrode sensitive area to obtain the desired electrochemical-field effect transistor dual-mode biosensor.

10. Use of the electrochemical-field effect transistor dual-mode biosensor according to any one of claims 1 to 8 in detecting samples with a wide concentration range and complex components.

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