Flexible OECT non-enzymatic glucose sensor and preparation method thereof

Abstract

The invention aims to provide a flexible OECT non-enzymatic glucose sensor and a preparation method thereof, and belongs to the technical field of glucose electrochemical sensors. According to the sensor, a chemically modified laser-induced graphene (LIG) electrode is used as a grid electrode, and gold nanoclusters (AuNC) are modified on LIG through controllable in-situ electrochemical reduction; meanwhile, the pH-dependent conductivity of PEDOT: PSS is utilized to realize the self-calibration of the sensor to the pH change of human sweat, so that the interference is reduced, and the measurement accuracy is improved. The flexible non-enzymatic glucose sensor designed by the invention has the advantages of high detection sensitivity, low detection limit, excellent long-term stability, no influence of pH value fluctuation and the like, and can be widely applied to the field of wearable electronics.

Description

Technical Field

[0001] The invention belongs to the technical field of glucose electrochemical sensors, and in particular relates to a flexible OECT non-enzymatic glucose sensor and a preparation method thereof. Background Art

[0002] With the rapid development of flexible wearable devices, smart medical and other fields, continuous monitoring of glucose concentration in different biological fluids is of great significance in the evaluation of chronic diseases and metabolic processes. Although isotope dilution gas chromatography-mass spectrometry has been widely adopted as the gold standard measurement method, achieving high-precision, continuous and real-time glucose concentration monitoring that meets clinical requirements is still a major challenge in health management and chronic disease treatment. In order to solve this problem, wearable sensors based on optical or electrochemical principles have been widely innovated and developed in recent years. Optical sensors use the interaction between light and glucose molecules to achieve detection, but optical sensors usually require the use of precise, expensive and bulky optical systems, which are severely limited in clinical point-to-point applications.

[0003] Electrochemical sensors have become an alternative because of their simple structure and low cost. According to the working principle, electrochemical glucose sensors can be divided into enzyme-based sensors (based on glucose oxidase or glucose dehydrogenase) and non-enzyme-based sensors (based on precious metals or transition metal oxides). At the same time, since the content of glucose in biological fluids is extremely trace, both electrochemical glucose sensors are required to have high sensitivity. One of the common strategies is to increase the electrochemical reaction sites by introducing micro-nano structures to improve sensitivity. The nanostructured gold electrode prepared by etching can significantly increase the enzyme loading, thereby increasing the sensitivity of glucose by three times. However, enzyme-based sensors usually have problems such as poor stability and susceptibility to temperature fluctuations. Non-enzymatic glucose sensors have higher sensitivity and long-term stability because there is no problem of enzyme degradation. However, current non-enzymatic glucose sensors are all based on the traditional three-electrode system and have the defect that their sensitivity increases with the increase of pH value. For example, the sensitivity is highest in a strongly alkaline environment. The pH value of human body fluids is generally neutral, but the sensitivity of non-enzymatic glucose sensors under neutral conditions is still insufficient, which limits their practical application in the human body.

[0004] In addition, the use of organic electrochemical transistors (OECTs) to reduce the detection limit of electrochemical sensors has also been reported. For example, carbon nanotubes and platinum nanoparticles are used to increase the gate surface area of ​​OECTs, thereby reducing the detection limit. However, the detection range of this solution is not enough to fully cover the glucose level in the human body, and it is not suitable for neutral or slightly acidic human body fluid pH environments.

[0005] How to design a flexible OECT non-enzymatic sensor to achieve highly sensitive detection of glucose in human body fluids. Summary of the Invention

[0006] Aiming at the problems existing in the background technology, the purpose of the present invention is to provide a flexible OECT non-enzymatic glucose sensor and its preparation method. This sensor uses a chemically modified laser-induced graphene (LIG) electrode as the gate, and gold nanoclusters (AuNC) are modified on the LIG through controllable in-situ electrochemical reduction; at the same time, by utilizing the pH-dependent conductivity of PEDOT:PSS (poly(3,4-ethylenedioxythiophene)-polystyrenesulfonic acid), the self-calibration of the sensor of the present invention to the pH change of human body fluids is realized, thereby reducing interference and improving the accuracy of measurement. The flexible non-enzymatic glucose sensor designed by the present invention has the advantages of high detection sensitivity, low detection limit, excellent long-term stability, and self-calibration in a dynamic pH environment, and can be widely applied to the field of wearable electronics.

[0007] To achieve the above purpose, the technical solution of the present invention is as follows:

[0008] A flexible OECT non-enzymatic glucose sensor includes a substrate, as well as a source electrode, a drain electrode, a channel, and a gate disposed on the surface of the substrate;

[0009] The source electrode is composed of an arc, a first rectangular strip connected to one end of the arc, a second rectangular strip connected to the other end of the arc, and a first rectangular block connected to the first rectangular strip; the drain electrode has the same structure as the source electrode and is symmetric about the gate left and right; the gate is disposed in the middle of the source electrode and the drain electrode, and is composed of a circular sensing area, a third rectangular strip connected to the circular sensing area, and a second rectangular block connected to the third rectangular strip, wherein the circular sensing area is disposed at the center of the arcs of the source electrode and the drain electrode;

[0010] The two second rectangular strips are adjacent but do not touch, and the area between the two second rectangular strips is a rectangular channel area; the channel is obtained by covering the channel area with a PEDOT:PSS thin film;

[0011] The material of the circular sensing area of the gate is laser-induced graphene modified with gold nanoclusters (AuNC / LIG), on which a perfluorosulfonic acid-based polymer layer is attached.

[0012] Furthermore, the smaller the ratio of the length to the width of the channel, the better the detection performance of the flexible OECT non-enzymatic glucose sensor.

[0013] Furthermore, the flexible OECT non-enzymatic glucose sensor preferably includes a microchannel, and the microchannel is used to concentrate body fluids to flow through the circular sensing area and the channel of the sensor to achieve non-invasive glucose detection.

[0014] Further, the method for pH self-calibration of the flexible OECT non-enzymatic glucose sensor of the present invention includes the following steps:

[0015] Step 1. Test the transfer curve of the OECT for phosphate buffer solutions (PBS) with different pH values, where the gate voltage V gs is set to the third voltage value, and the voltage between the source and drain (V ds ) is set to the second voltage value, to obtain the channel current I ds ,

[0016] Step 2. At different pH values, set the gate voltage to the third voltage value, add glucose at different concentrations, and measure the current I ds有葡萄糖 under different glucose concentrations. Define the normalized current response NCR, NCR = (I ds有葡萄糖 -I ds无葡萄糖 ) / I ds无葡萄糖 , and based on this, obtain the calibration curve between the glucose concentration and NCR;

[0017] Step 3. Set the gate voltage to the first voltage value and the voltage between the source and drain to the second voltage value, measure the channel current at this time, and corresponding to the transfer curve obtained in Step 1, obtain the pH value at this time;

[0018] Then set the gate voltage to the third voltage value, drop the glucose solution to be measured, measure the channel current at this time, and calculate the NCR at this time;

[0019] According to the pH value and the normalized current response NCR, search the calibration curve to obtain the concentration of the glucose solution to be measured.

[0020] Further, the first voltage value is preferably 0 V, the second voltage value is preferably -0.5 V, and the third voltage value is 0 - 1 V, preferably 0.5 V.

[0021] A method for preparing a flexible OECT non-enzymatic glucose sensor includes the following steps:

[0022] S1. Prepare a polyimide substrate film;

[0023] S2. Prepare patterned source and drain electrodes by photolithography combined with magnetron sputtering;

[0024] S3. Spin-coat PEDOT:PSS solution on the entire PI surface, then heat and cure it. Then prepare a photoresist in the channel area between the source and drain electrodes, and then perform photolithography and plasma etching, leaving only the photoresist covering the PEDOT:PSS film on the channel, and finally remove the photoresist to obtain a PEDOT:PSS channel;

[0025] S4. Use CO2 A patterned laser-induced graphene (LIG) electrode is prepared by laser processing, and then gold nanoclusters (AuNC) are deposited on the LIG by in-situ electrochemical reduction to obtain an AuNC / LIG gate;

[0026] S5. Drop a perfluorosulfonic acid-based polymer (Nafion) solution on the AuNC / LIG gate and allow it to dry naturally into a film as an anti-interference layer, then the required sensor can be prepared.

[0027] Furthermore, the specific process of step S4 is as follows:

[0028] S4.1. Use CO 2 laser to process the polyimide to obtain a patterned LIG gate;

[0029] S4.2. Connect the copper foil and the rectangular area of the LIG gate with a silver wire, and then encapsulate the LIG gate with PDMS, only exposing the circular sensing area of the LIG;

[0030] S4.3. Immerse the sample in a solution of chloroauric acid trihydrate (HAuCl 4 ·3H 2 O) for a period of time to deposit a layer of gold nanoclusters on the circular sensing area of the LIG gate, then take out the sample, wash it, and then remove the silver wire and copper foil to obtain the AuNC / LIG gate.

[0031] Furthermore, the power of the CO 2 laser is 8.5% - 12.5% of the maximum power, the maximum power is preferably 30 W, the speed is 8% - 20% of the maximum speed, the maximum speed is preferably 1270 mm / s; when depositing gold nanoclusters, the concentration of chloroauric acid trihydrate (HAuCl 4 ·3H 2 O) in the chloroauric acid trihydrate (HAuCl 4 ·3H 2 O) solution is 1 mg / mL - 36 mg / mL, and the deposition time is controlled to be 0.5 - 1 hour.

[0032] The mechanism of the present invention is as follows:

[0033] The redox peaks of the AuNC / LIG electrode of the present invention are significantly higher than those of the planar Au and the original LIG electrodes, and it has better electrochemical reaction activity. At the same time, after the LIG electrode is modified with AuNC, the charge transfer resistance (Rct) in EIS decreases from 310 Ω to 250 Ω. Due to the porous structure of LIG and the modification of AuNC, its diffusion rate is further increased. This enables the sensor of the present invention to have excellent sensing performance for glucose and can quickly and accurately detect the presence and concentration changes of glucose. At the same time, the sensor of the present invention is constructed based on OECT. OECT can amplify the current signal. At the same pH value, OECT can amplify the current signal by 2-3 times compared with the traditional three-electrode system. Therefore, it also has high sensitivity under acidic or neutral conditions. Thus, the non-enzymatic glucose sensor based on OECT can be detected at different pH values. Combined with pH self-calibration, the accuracy of the measurement results is greatly improved.

[0034] The present invention uses in-situ electrochemical reduction method to modify the sensor. Compared with the traditional electroplating method, in-situ electrochemical reduction can provide a more uniform and stable metal coating on a porous substrate with poor conductivity (such as the LIG electrode). The AuNC modified by this method in the present invention is evenly distributed on the LIG electrode. This uniformity ensures the consistency of glucose detection in different regions, avoids detection deviation caused by uneven coating, and improves the reliability of the sensor. At the same time, during the in-situ electrochemical reduction process, the loading amount and morphology of AuNC can be precisely controlled by the concentration and deposition time of the HAuCl 4 ・3H 2 O precursor solution, enabling researchers to optimize the sensor performance according to actual needs and regulate performance such as the detection sensitivity of glucose by adjusting the state of AuNC.

[0035] In summary, due to the adoption of the above technical solutions, the beneficial effects of the present invention are as follows:

[0036] 1. The present invention uses the in-situ electrochemical reduction method to achieve highly uniform, stable and precisely controllable AuNC modification on the miniaturized LIG gate. By combining the AuNC / LIG gate with OECT, a highly sensitive flexible non-enzymatic glucose sensor is obtained, with a detection limit as low as 0.08 µM at pH = 7.4, and its performance is superior to all previously reported non-enzymatic glucose sensors.

[0037] 2. The non-enzymatic glucose sensor of the present invention based on OECT has the ability of self-pH calibration and can accurately measure the glucose concentration under the dynamic change of the pH of human sweat. At the same time, the non-enzymatic glucose sensor based on OECT shows excellent long-term stability, and its performance only decreases by about 10% within 180 days.

[0038] 3. The preparation method of the present invention is suitable for large-scale production. The modification method based on in-situ electrochemical reduction performs excellently in large-scale production, with the advantages of high consistency and low inter-sample difference. The electrochemical impedance spectra (EIS) of ten AuNC / LIG electrodes prepared in the same batch show that the inter-sample difference is about 2.2%, and the sensing performance difference between different batches of samples is only 3.4%. This high degree of consistency provides a strong guarantee for the large-scale production and application of the sensor, reduces the production cost, and improves the production efficiency. Brief Description of the Drawings

[0039] Figure 1 It is a schematic structural diagram of the OECT-based non-enzymatic glucose sensor of the invention.

[0040] Figure 2 It is a preparation flow chart of the OECT-based non-enzymatic glucose sensor of the present invention.

[0041] Figure 3 It is an optical picture of mass-producing AuNC / LIG electrodes.

[0042] Figure 4 It is an SEM image of LIG and AuNC / LIG electrodes.

[0043] Figure 5 It is for [Fe(CN) 6 ³⁻ / 4 ⁻ cyclic voltammetry (CV) curves of LIG electrodes, AuNC / LIG electrodes and planar gold (p_Au) electrodes measured in the redox solution.

[0044] Figure 6 It is a graph of current response data of the OECT-based non-enzymatic glucose sensor to gradually added glucose.

[0045] Figure 7 It is a normalized current response (NCR) graph of the OECT-based non-enzymatic glucose sensor in different pH solutions.

[0046] Figure 8 It is a comparison graph of the long-term stability of the OECT-based non-enzymatic glucose sensor and the OECT-based enzymatic glucose sensor during 180 days of measurement.

[0047] Figure 9 It is a curve of the channel current of OECT varying with the solution pH value when the gate voltage is 0 V.

[0048] Figure 10 It is the anti-interference test of the OECT-based non-enzymatic glucose sensor.

[0049] Figure 11Optical pictures of the microchannel and the integrated microchannel and OECT-based non-enzymatic glucose sensor. Detailed implementation mode

[0050] To make the objectives, technical solutions and advantages of the present invention clearer, the present invention will be further described in detail below in conjunction with the implementation modes and the drawings.

[0051] A flexible OECT non-enzymatic glucose sensor, the structural schematic diagram of which is as Figure 1 shown, including a substrate, and a source electrode (S), a drain electrode (D), a channel and a gate electrode (G) arranged on the surface of the substrate;

[0052] The source electrode is composed of an arc, a first rectangular strip connected to one end of the arc, a second rectangular strip connected to the other end of the arc, and a first rectangular block connected to the first rectangular strip; the drain electrode has the same structure as the source electrode and is symmetric about the gate electrode; the gate electrode is arranged in the middle of the source electrode and the drain electrode, and is composed of a circular sensing area, a third rectangular strip connected to the circular sensing area, and a second rectangular block connected to the third rectangular strip. Among them, the circular sensing area is arranged at the center of the arcs of the source electrode and the drain electrode, and the diameter of the circular sensing area is d;

[0053] The two second rectangular strips are adjacent but not in contact, and the area between the two second rectangular strips is a rectangular channel area. The length of the channel is l and the width is w; the channel is obtained by covering the channel area with a PEDOT:PSS thin film;

[0054] The material of the circular sensing area of the gate electrode is laser-induced graphene modified with gold nanoclusters (AuNC / LIG), on which a Nafion layer is attached.

[0055] The present invention can further combine the device with a microchannel to obtain a sweat patch. The microchannel is used to concentrate sweat to flow through the circular sensing area and the channel of the sensor for non-invasive glucose detection.

[0056] Example 1

[0057] A preparation method of a flexible OECT non-enzymatic glucose sensor, the preparation flow chart of which is as Figure 2 shown, including the following steps:

[0058] (a) Coat a layer of polydimethylsiloxane (PDMS) on a glass slide as an adhesive, then place a polyimide film layer (PI) on the PDMS, and then heat at 120 °C for 10 minutes to cure the PDMS, so that the PI film is fixed on the glass slide;

[0059] (b)Prepare a patterned mask on the surface of the PI film by photolithography: Spin-coat a photoresist on the PI film. The photoresist model is AZ5214. After spin-coating, place the glass slide on a hot plate at 100 °C and heat for 60 seconds. After heating, perform the first exposure with an exposure time of 4 seconds. After the exposure, place it on a hot plate at 120 °C and heat for 90 seconds. After heating, perform the second exposure with an exposure time of 40 seconds. After the exposure, soak it in the developer for 45 seconds to remove the excess photoresist, and a patterned mask can be obtained;

[0060] (c)Sputter a layer of Au film on the mask by magnetron sputtering;

[0061] (d)Soak the glass slide in acetone to remove the photoresist, and a patterned gold source electrode and gold drain electrode can be obtained;

[0062] (e)Spin-coat the PEDOT:PSS solution on the entire PI surface and cure it by heating at 120 °C for 1 hour; Then use photolithography to prepare a rectangular photoresist in the channel area of the source and drain electrodes. The photoresist model is AZ6112. After spin-coating the photoresist, heat it on a hot plate at 100 °C for 60 seconds, expose it for 4 seconds, and finally soak it in the developer for 45 seconds to remove the excess photoresist, leaving only a rectangular photoresist covering the PEDOT:PSS film on the channel as a mask; Then use oxygen plasma etching to etch away all the PEDOT:PSS not covered by the photoresist. Finally, soak the device in acetone to remove the photoresist on the channel, and a rectangular PEDOT:PSS channel can be obtained;

[0063] (f)Use a CO 2 laser to sinter the blank area of the polyimide. After sintering, the surface layer of the polyimide film will become graphene, and an LIG pattern can be obtained;

[0064] (g)Connect the copper foil and the rectangular area of the LIG gate with a silver wire, and then encapsulate the LIG gate with PDMS, leaving only the circular sensing area of the LIG exposed;

[0065] (h)Soak the device in a solution of chloroauric acid trihydrate (HAuCl 4 ·3H 2 O) for one hour to deposit gold nanoclusters on the LIG electrode. After rinsing with ultrapure water, remove the silver wire and copper foil to obtain an AuNC / LIG gate;

[0066] The present invention adopts the in-situ electrochemical reduction method and can also achieve highly uniform, stable and precisely controllable AuNC modification on miniaturized LIG electrodes in batches, as Figure 3 shown; The SEM images of the circular areas of the LIG electrodes before and after modification are as Figure 4As shown; it can be seen from the figure that the LIG electrode exhibits a typical porous structure with an average pore size of about 4 μm, and this structure provides a large surface area for electrochemical sensing and energy storage. After the AuNC modification, a dense nanosheet structure is obtained.

[0067] (i)Drop Nafion solution on the AuNC / LIG gate as an anti-interference layer;

[0068] (j)Use SU-8 photoresist to encapsulate the entire device, only exposing the channel, the gate to be processed, and the contact points of the source, drain, and gate respectively; the purpose of encapsulation is to avoid conduction between the source and the gate, and the reason for exposing the latter part is that the working principle of the OECT is to apply a voltage between the gate and the source to drive the directional movement of ions in the solution into the PEDOT:PSS inside the channel, changing the doping state of PEDOT:PSS, thereby causing a change in the current at the channel (i.e., the conductivity of PEDOT:PSS). The contact points refer to Figure 1 the three small rectangles marked D, G, and S in the figure;

[0069] The preparation of the OECT-based non-enzymatic glucose sensor is completed.

[0070] In order to verify the performance and indicators of the OECT-based non-enzymatic glucose sensor prepared by the present invention, various different electrochemical detections are used to explore its sensing performance.

[0071] Figure 5 For the cyclic voltammetry (CV) curves of the LIG electrode, AuNC / LIG electrode, and planar gold (p_Au) electrode measured in a 5 mM [Fe(CN) 6 ³⁻ / 4 ⁻ redox solution. The CV test conditions are: the scanning range is -0.2 V to 0.6 V, the scanning rate is 50 mV / s, and the sensor of the present invention is used as the working electrode (WE), and together with the counter electrode (CE) and reference electrode (RE) to form a three-electrode working system. In the CV test, due to the presence of the Fe²⁺ / Fe³⁺ redox couple, the redox peaks corresponding to the AuNC / LIG electrode are significantly higher than those of the planar gold (p_Au) electrode and the original LIG electrode, which indicates that the modification of gold nanoclusters can provide a larger specific surface area of the electrode, significantly increasing the electrochemistry reaction activity, that is, indicating that the AuNC / LIG electrode of the present invention has high redox characteristics, which is beneficial to the occurrence of the oxidation reaction of glucose and can better realize the detection of glucose.

[0072] Figure 6 For the current response of the OECT-based non-enzymatic glucose sensor to the gradually added glucose, the glucose concentration ranges from 10 -8M to 30 mM (pH = 7.4). The OECT-based non-enzymatic glucose sensor starts to show a current response upon the addition of glucose at a concentration of 10⁻ 7 M and tends to saturate when the concentration reaches approximately 30 mM, that is, the experimental lowest detection limit of the sensor of the present invention is 10⁻ 7 M.

[0073] Figure 7 is the normalized current response (NCR) of the OECT-based non-enzymatic glucose sensor in different pH solutions. The OECT-based non-enzymatic glucose sensor exhibits significant pH-dependence. In an alkaline environment (such as pH 8.5 - 9.5), its NCR increases with the increase of glucose concentration until saturation, and within the low concentration range (10⁻ 8 -10⁻ 6 M), the NCR increases significantly with the increase of pH. This is due to the improvement of the electrochemical activity of the non-enzymatic AuNC / LIG electrode under alkaline conditions (under alkaline conditions, there are more hydroxide ions (OH⁻), and OH⁻ will adsorb on the surface of AuNC to form AuOH ads (ads means adsorbed), and AuOH ads this substance will oxidize glucose into gluconate. In addition, the detection limit of the sensor decreases with the increase of pH (the detection limit is 0.01 μM at pH 9.5 and 10 μM at pH 6.5), but the maximum measurable concentration shows the opposite trend (the maximum measurable concentration is 0.1 mM at pH 9.5 and greater than 5 mM at pH 6.5). It should be noted that under near-neutral conditions close to human body fluids (pH = 7.4), the OECT-based non-enzymatic glucose sensor of the present invention still maintains an ultra-low detection limit (about 0.08 μM), which is lower than the trace glucose concentration in human body fluids.

[0074] Figure 8 shows the comparison of the long-term stability of the OECT-based non-enzymatic glucose sensor and the OECT-based enzymatic glucose sensor when stored in air and phosphate buffered saline solution (PBS) during a 180-day measurement period, which is reflected by the ratio of the normalized current response to the initial normalized current response (NCR / NCR 0 ). The OECT-based non-enzymatic glucose sensor exhibits excellent long-term stability, with only about 10% performance decay during a 180-day test. Its performance is significantly better than that of the enzymatic OECT sensor and much higher than the long-term stability of existing non-enzymatic glucose sensors. The comparison results are shown in Table 1.

[0075] Table 1. Comparison of the long-term stability of non-enzymatic glucose sensors

[0076] Electrode material Long-term stability References Nanoporous platinum / glassy carbon electrode Maintain 90% of the initial response after one month Xu, Q. et al., Facile fabrication of nanoporous platinum by alloying–dealloying process and its application in glucose sensing. Sensors and Actuators B: Chemical 2012, 173, 716-723 Gold foam / silicon electrode Maintain 90% of the initial response after 30 days Shen, N. et al., Highly Responsive and Ultrasensitive Non-Enzymatic Electrochemical Glucose Sensor Based on Au Foam. Sensors (Basel, Switzerland) 2019, 19 (5) <![CDATA[MOF(Ni-BDC-NH 2 ) / Glassy carbon electrode]]> Maintain 77% of the initial response after 180 days Daud, A. D. et al., An effective metal-organic framework-based electrochemical non-enzymatic glucose sensor. Journal of Electroanalytical Chemistry 2022, 921, 116676 AuNC / LIG electrode Maintain 90% of the initial response after 180 days The present invention

[0077] Figure 9 The variation curve of the channel current of the OECT with the solution pH value when the gate voltage is 0 V (drain-source voltage V ds = -0.5 V). The present invention also proposes a pH self-calibration method, specifically based on the pH response characteristics of the PEDOT:PSS channel material in the OECT. Its core principle lies in that the conductivity of PEDOT:PSS will change with the environmental pH value. Under alkaline conditions, OH⁻ will destroy the π-π stacking structure of the PEDOT⁺ chain, resulting in a decrease in conductivity. The pH calibration process adopts a two-step method: In the first step, set the voltage between the source and the drain (V ds = 0 V) to -0.5 V, and measure the channel current (I gs = 0 V) under zero gate voltage (V ds ). Determine the pH value of the current solution through the pre-calibrated curve of pH versus channel current ( Figure 7 ). In the second step, set the gate voltage to 0.5 V (V gs = 0.5 V), drop glucose, and obtain the calibration curve of the accurate glucose concentration through the NCR corresponding to this pH value. Finally, determine the concentration of the glucose solution to be measured based on the calibration curve. This pH self-calibration method can significantly reduce the measurement error caused by different pH values of sweat and obtain a more accurate glucose concentration.

[0078] Figure 10 Anti-interference test of the non-enzymatic glucose sensor based on OECT. The numbers on the bar graph represent the relative response values of each interferent signal to the glucose signal. The results show that due to the addition of the Nafion film, the non-enzymatic glucose sensor based on OECT only has a significant response to glucose and has good anti-interference characteristics.

[0079] Figure 11 Optical pictures of the microchannel and the integration of the microchannel and the non-enzymatic glucose sensor based on OECT. The flow of red ink was used to simulate and demonstrate the microchannel. It can be seen from the figure that the microchannel can make the liquid to be measured contact the sensing unit better and faster, so as to realize the detection of the liquid to be measured.

[0080] As described above, it is only the specific implementation manner of the present invention. Any feature disclosed in this specification, unless specifically stated, can be replaced by other equivalent or features with similar purposes; all the features disclosed, or all the steps in any method or process, except for mutually exclusive features and / or steps, can be combined in any way.

Claims

1. A flexible OECT non-enzymatic glucose sensor, characterized in that: It includes a substrate, and a source, a drain, a channel and a gate arranged on the surface of the substrate; The source electrode is composed of a circular arc, a first rectangular strip connected to one end of the circular arc, a second rectangular strip connected to the other end of the circular arc, and a first rectangular block connected to the first rectangular strip; the drain electrode has the same structure as the source electrode and is symmetrical about the gate electrode; The gate is arranged between the source and the drain, and is composed of a circular sensing area, a third rectangular strip connected to the circular sensing area, and a second rectangular block connected to the third rectangular strip, wherein the circular sensing area is arranged at the center of the two arcs of the source and the drain; The two second rectangular strips are adjacent to each other but not in contact, and the area between the two second rectangular strips is a rectangular channel area; the channel is obtained by covering the channel area with a PEDOT:PSS film; The circular sensing area material of the gate is laser-induced graphene modified by gold nanoclusters, on which a perfluorosulfonic acid-based polymer layer is attached.

2. The flexible OECT non-enzymatic glucose sensor according to claim 1, characterized in that The smaller the ratio of the length to the width of the channel, the better the detection performance of the flexible OECT non-enzymatic glucose sensor.

3. The flexible OECT non-enzymatic glucose sensor according to claim 1, characterized in that: The flexible OECT non-enzymatic glucose sensor comprises a microchannel, and the microchannel is used to concentrate body fluid to flow in the circular sensing area and channel of the sensor, thereby realizing non-invasive glucose detection.

4. The method for pH self-calibration of a flexible OECT non-enzymatic glucose sensor according to any one of claims 1 to 3, characterized in that: The following steps are involved: Step 1. Test the transfer curves of OECTs for phosphate buffers with different pH values, where the gate voltage V gs Set to a third voltage value, the voltage between the source and drain V ds Set to the second voltage value, and get the channel current I ds , Step 2. At different pH values, set the gate voltage to the third voltage value, add different concentrations of glucose, and measure the current I at different concentrations of glucose ds有葡萄糖 , define the normalized current response NCR, NCR = (I ds有葡萄糖 -I ds无葡萄糖 ) / I ds无葡萄糖 , based on which a calibration curve between glucose concentration and NCR is obtained; Step 3. Set the gate voltage to a first voltage value, the voltage between the source and the drain to a second voltage value, measure the channel current at this time, and obtain the pH value at this time corresponding to the transfer curve obtained in step 1; Then, the gate voltage is set to a third voltage value, the glucose solution to be tested is added dropwise, the channel current at this time is measured, and the NCR at this time is calculated; According to the pH value and the normalized current response NCR, the concentration of the glucose solution to be tested can be obtained by looking up the calibration curve.

5. The method for pH self-calibration of a flexible OECT non-enzymatic glucose sensor according to claim 4, characterized in that: The first voltage value is 0V, the second voltage value is -0.5V, and the third voltage value is 0.5V.

6. A method for preparing a flexible OECT non-enzymatic glucose sensor according to any one of claims 1 to 3, characterized in that: The following steps are involved: S1. Preparing a polyimide substrate film; S2. Patterned source and drain electrodes are prepared by photolithography combined with magnetron sputtering; S3. Spin-coat the PEDOT:PSS solution on the entire PI surface, heat and solidify it, prepare a photoresist in the channel region between the source and drain electrodes, perform photolithography and plasma etching, and only leave the photoresist covering the PEDOT:PSS film on the channel. Finally, remove the photoresist to obtain a PEDOT:PSS channel. S4. A patterned laser-induced graphene LIG electrode was prepared by CO2 laser processing, and then gold nanoclusters AuNC were deposited on LIG by in situ electrochemical reduction to obtain AuNC / LIG gate; S5. Add perfluorosulfonic acid polymer solution onto the AuNC / LIG gate and allow it to dry naturally into a thin film as an anti-interference layer to prepare the desired sensor.

7. The preparation method according to claim 6, characterized in that: The specific process of S4 is: S4.

1. Processing polyimide using CO2 laser to obtain patterned LIG gates; S4.

2. Connect the copper foil and the rectangular area of ​​the LIG gate with silver wire, and then encapsulate the LIG gate with PDMS, leaving only the circular sensing area of ​​the LIG exposed; S4.

3. Soak the sample in a tetrachloroauric acid trihydrate solution for a period of time to deposit a layer of gold nanoclusters on the circular sensing area of ​​the LIG gate. Then take out the sample, clean it, and remove the silver wire and copper foil to obtain the AuNC / LIG gate.

8. The preparation method according to claim 7, characterized in that: The power of CO2 laser in S4.1 is 2.55-3.75 W, and the speed is 101.6-254 mm / s.

9. The preparation method according to claim 7, characterized in that: When gold nanoclusters are deposited in S4.3, the concentration of tetrachloroauric acid trihydrate in the tetrachloroauric acid trihydrate solution is 1 mg / mL to 36 mg / mL, and the deposition time is controlled to be 0.5 hour to 1 hour.

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