Flexible injectable organic electrochemical transistor sensor for detecting protein immune markers as well as preparation method and application of flexible injectable organic electrochemical transistor sensor

Abstract

The invention discloses a flexible injectable organic electrochemical transistor sensor for protein immune marker detection and a preparation method and application thereof, and relates to the technical field of biochemical sensors, the flexible injectable organic electrochemical transistor sensor comprises a bottom flexible substrate, a metal source electrode, a metal drain electrode and a metal grid electrode which are arranged on the surface of the flexible substrate, and a top packaging layer, an organic semiconductor conductive channel is arranged between the metal source electrode and the metal drain electrode, the metal grid electrode and the organic semiconductor conductive channel are arranged in a coplanar manner, the top packaging layer plays a role in packaging, and the organic semiconductor conductive channel and the metal grid electrode are exposed outside after packaging; the metal gate is modified with a monomolecular layer formed by self-assembly of 11-mercaptoundecanoic acid, and the monomolecular layer is covalently coupled with an amino-modified single-stranded DNA molecule aptamer. The sensor can be used for detecting various protein molecules including IFN-gamma tumor markers.

Description

Technical Field

[0001] This invention relates to the field of biochemical sensor technology, specifically to a flexible injectable organic electrochemical transistor sensor for the detection of protein immunomarkers, its preparation method, and its application. Background Technology

[0002] With the development of precision medicine, the clinical demand for in-situ, continuous, and highly sensitive monitoring of biomolecules in vivo continues to rise. Traditional in vitro detection methods (such as ELISA and chemiluminescence) rely on repeated sampling and laboratory analysis, which have limitations such as response delay and unsuitability for dynamic processes. In recent years, some studies have proposed injectable field-effect transistor (FET) sensors for in vivo implantable biosignal detection, but several core technical bottlenecks remain, such as limited signal amplification capabilities, high material and structural rigidity, and poor biocompatibility.

[0003] Chinese patent document CN118533936A discloses an injectable needle-type field-effect transistor biosensor, its fabrication method, and its application. This invention designs an injectable needle-type FET detection and sensing microneedle device with a hollow chamber, integrating drug injection and FET detection and sensing functions. However, the FET architecture lacks the advantages of organic electrochemical transistors in amplifying electrical signals and enhancing ion coupling detection sensitivity. Furthermore, this sensor is not suitable for long-term in vivo operation.

[0004] Chinese patent document CN119881036A discloses a malaria-related biomarker detection chip based on an integrated organic electrochemical transistor. The sensor uses an electrochemical deposition method to obtain a three-dimensional gold electrode on the surface of a two-dimensional gold electrode. The surface of the three-dimensional gold electrode is modified with single-stranded DNA aptamers for recognizing malaria-related biomarkers. The three-dimensional structure of the gold electrode can increase the modification density of aptamer molecules, which can improve the transconductance of the coplanar organic electrochemical transistor and achieve the purpose of detecting target molecules at extremely low concentrations. However, the three-dimensional structure on its surface increases the surface degrees of freedom and roughness of the metal gate, reduces the electron mobility of the transistor, increases the operating voltage, and the high affinity recognition molecules have a slow dissociation rate after binding to the target, which is not conducive to continuous tracking.

[0005] Current sensors based on organic electrochemical transistors generally suffer from drawbacks such as large size and high implantation difficulty. Most existing designs cannot be implanted in vivo via injection and are limited to patch or surgical fixation, thus restricting clinical applications. In addition, the high-affinity recognition molecules (such as aptamers) in these sensors have a slow dissociation rate after binding to the target, and existing devices generally lack auxiliary dissociation mechanisms, making signal self-regeneration impossible and hindering continuous tracking.

[0006] In addition, currently available organic electrochemical transistor (OECT) or field-effect transistor (FET) sensors are mostly focused on the detection of electrolytes or metabolites, and have not yet achieved highly sensitive detection of protein-level immunomarkers such as interferon-γ (IFN-γ). Although existing in vitro detection technologies (such as immunochromatographic test strips) are specific, their detection limits are often around 1 μg / mL, which cannot meet the dynamic monitoring requirements at the pg / mL level.

[0007] Therefore, how to construct a transistor sensor with high sensitivity, miniaturization, flexible injectability, dynamic response capability, and suitability for in-situ detection of protein-level biomarkers is one of the key technical problems that urgently need to be solved in this field. Summary of the Invention

[0008] To address the aforementioned technical problems, this invention provides a flexible injectable organic electrochemical transistor sensor for the detection of protein immunomarkers. By introducing an organic channel material (PEDOT:PSS) with an ion-electron coupling mechanism, the device possesses a natural current amplification capability, making it particularly suitable for the detection of low-concentration target substances. The device is constructed entirely of a flexible polymer material, is small in size, can be rolled up, and can be injected subcutaneously or into tissues via a fine needle. It is stable and reusable over a long period and adaptable to complex biological environments.

[0009] The specific technical solution adopted is as follows: A flexible injectable organic electrochemical transistor sensor for detecting protein immunomarkers includes a bottom flexible substrate, a metal source, a metal drain, and a metal gate disposed on the surface of the flexible substrate, and a top encapsulation layer. An organic semiconductor conductive channel is disposed between the metal source and the metal drain. The metal gate is coplanar with the organic semiconductor conductive channel. The top encapsulation layer encapsulates the sensor, exposing the organic semiconductor conductive channel and the metal gate after encapsulation. The metal gate is modified with a monolayer formed by the self-assembly of 11-mercaptoundecanoic acid, and an amino-modified single-stranded DNA aptamer is covalently coupled to the monolayer.

[0010] This invention further enhances the device's specific recognition capability for target molecules and signal response performance by introducing a self-assembled monolayer (SAM) and aptamer recognition elements onto the gate electrode surface. Specifically, 11-mercaptoundecanoic acid (11-MUA) is used to form an ordered SAM structure with carboxyl ends on the metal gate surface, providing a stable and controllable chemical interface for the subsequent covalent fixation of functional molecules. Based on this, a single-stranded DNA aptamer with a 5′ amino-terminal modification is amidated with the carboxyl groups on the SAM layer via EDC / NHS chemical cross-linking, thereby constructing a high-affinity molecular recognition layer on the gate surface that specifically recognizes target proteins (such as IFN-γ).

[0011] Preferably, the flexible substrate is polyimide, SU-8, polyethylene terephthalate (PET), or polydimethylsiloxane (PDMS), providing overall mechanical flexibility and optical transparency of the device.

[0012] Preferably, the resistance of the organic semiconductor conductive channel is no greater than the kiloohm level to ensure that the flexible injectable organic electrochemical transistor sensor has good sensitivity.

[0013] Preferably, the organic semiconductor conductive channel is obtained by curing an organic semiconductor composite solution composed of PEDOT:PSS, ethylene glycol, 3-glycidoxypropyltrimethoxysilane and dodecylbenzenesulfonic acid.

[0014] More preferably, in the organic semiconductor composite solution, the content of PEDOT:PSS is 90-96 wt%, the content of ethylene glycol is 2-8 wt%, the content of 3-glycidoxypropyltrimethoxysilane is 0.1-2 wt%, and the content of dodecylbenzenesulfonic acid is 0.01-0.50 wt%.

[0015] Specifically, after the flexible injectable organic electrochemical transistor sensor binds to the target protein-based immunomarker, the conjugate is dissociated by an electrical pulse or buffer elution strategy (the electrical pulse is applied to the gate and drain), thereby regenerating the signal after sensor detection and allowing for multiple uses.

[0016] Preferably, the protein-based immunomarker is IFN-γ, and the amino-modified single-stranded DNA aptamer is 5′-NH2-TAATTCCCAATCCATGTGTTGTGGGTTGTGTTGGTTGGGG-3′ (nucleotide sequence as shown in SEQ ID NO. 1) with 5′ amino-modified end.

[0017] Preferably, the flexible substrate has a thickness of 3-4 μm; the metal source, metal drain, and metal gate all have a thickness of 80-140 nm; the encapsulation layer has a thickness of 1-3 μm; and the organic semiconductor conductive channel has a thickness of 90-110 nm, a width of 50-700 μm, and a width-to-length ratio of 1-3:1.

[0018] Furthermore, the metal source, metal drain, and metal gate are all composed of a 5-15 nm thick bottom chromium metal layer and an 80-120 nm thick top gold metal layer. The bottom chromium metal layer can form a good adhesion with the flexible substrate, making the flexible injectable organic electrochemical transistor sensor more robust and more durable.

[0019] Preferably, the encapsulation layer is made of SU-8, polydimethylsiloxane (PDMS), or parylene-C.

[0020] The above parameter ranges were optimized and determined based on a full consideration of the device's sensitivity, response speed, and miniaturization requirements. In organic electrochemical transistors (OECTs), the channel current output mainly depends on the channel geometry and the regulation efficiency of carrier concentration. Increasing the channel width helps to improve the drain current (Id) per unit gate voltage. DS This enhances signal strength and increases the detection limit; while reducing the channel length effectively reduces the device's channel resistance, accelerates the response rate, and increases transconductance (g). m Therefore, under certain electrode size and voltage driving conditions, a larger aspect ratio (W / L) helps improve the transconductance output and signal amplification capability of the device, enhancing the detection sensitivity to slight concentration changes. However, an excessively high aspect ratio may lead to increased device power consumption, structural instability, or increased packaging difficulty, especially in injectable and miniaturized applications where a balance between size and performance must be struck. Experiments show that in the device structure of this invention, when the width is controlled within 50-700 μm and the aspect ratio is 1-3:1, the device maintains a small size and good mechanical flexibility while ensuring sensitivity, making it suitable for in vivo implantation and continuous monitoring applications.

[0021] More preferably, the organic semiconductor conductive channel has a width of 50 μm and a width-to-length ratio of 2:1. Under the same device size, the larger the channel width-to-length ratio, the higher the transconductance performance of the device. Under the same channel width-to-length ratio, the smaller the overall device size, the better its performance and the better its signal amplification capability.

[0022] Preferably, the flexible injectable organic electrochemical transistor sensor has a maximum size of 350-4900 μm. The flexible injectable organic electrochemical transistor sensor is rolled up and then injected using a syringe. The flexible injectable organic electrochemical transistor sensor has outwardly led signal lines.

[0023] The present invention also provides a method for fabricating the aforementioned flexible injectable organic electrochemical transistor sensor, comprising: (1) The metal source, metal drain and metal gate are deposited on the surface of a flexible substrate by evaporation and photolithography. (2) The organic semiconductor composite solution is spin-coated, cured and patterned to obtain an organic semiconductor channel layer, which is then placed between the metal source and the metal drain to obtain a sensor component; (3) Encapsulate the sensor components to expose the encapsulated organic semiconductor conductive channels and metal gates, clean them, and incubate them in an ethanol solution of 11-mercaptoundecanoic acid for ≥16 h. After complete incubation, activate them for 1-5 h, then immerse them in a solution of amino-modified single-stranded DNA aptamers for ≥12 h, and dry them to obtain the flexible injectable organic electrochemical transistor sensor.

[0024] This invention employs amino-modified single-stranded DNA aptamers covalently linked to self-assembled monolayers to ensure the aptamer's immobilization efficiency and recognition stability on the surface of flexible devices. Simultaneously, by using electro-pulse or buffer elution strategies to assist in the dissociation of the conjugates, signal regeneration after detection is achieved, enabling continuous dynamic monitoring.

[0025] Preferably, the flexible substrate is prepared by spin-coating a polyimide solution onto a substrate and then drying it.

[0026] Preferably, sulfuric acid is used to clean the metal gate region by cyclic voltammetry at -0.4 to 1.4V for 1 to 50 cycles.

[0027] Preferably, the concentration of the 11-mercaptoundecanoic acid ethanol solution is 1-1.5 mg / mL.

[0028] Preferably, the activation treatment uses an EDC / NHS activation buffer, wherein the concentration of EDC in the EDC / NHS activation buffer is 150-250 mM, the concentration of NHS is 40-60 mM, and the concentration of the amino-modified single-stranded DNA aptamer solution is 5-10 μM.

[0029] The present invention also provides a method for detecting IFN-γ for non-disease diagnosis or non-treatment purposes, utilizing the aforementioned flexible injectable organic electrochemical transistor sensor.

[0030] When the flexible injectable organic electrochemical transistor sensor is used in vitro, an electrolyte solution is added to it. The electrolyte solution can be a buffer solution (such as phosphate buffer), an inorganic salt ion solution (such as NaCl solution, KCl solution, or Na2SO4 solution), or an organic / ionic liquid electrolyte solution (such as EMIM:TFSI solution). When the flexible injectable organic electrochemical transistor sensor is used in vivo, human body fluids are used as the electrolyte solution.

[0031] Compared with the prior art, the beneficial effects of the present invention are as follows: (1) The method of the present invention can control the overall thickness of the device within tens of micrometers, achieving high mechanical compliance and miniaturized size of the device, enabling it to be implanted through conventional syringe needles, breaking through the limitation of existing devices being limited to adhesive or surgical implantation, and possessing true injectable capability.

[0032] (2) This invention optimizes the gate modification chemistry and channel structure parameters for the electrical response characteristics of protein-level biomarker INF-γ, significantly improves the device's current modulation capability for low-concentration proteins (pg / mL level), realizes the detection function of highly specific and highly sensitive immunomarkers, realizes the construction of a highly sensitive protein detection mechanism, and overcomes the problem that traditional OECT devices are mainly used for small molecule or ion detection.

[0033] (3) This invention uses amino-modified single-stranded DNA aptamers covalently linked to the SAM layer to ensure the immobilization efficiency and recognition stability of the aptamers on the surface of flexible devices. At the same time, the recognition complex is dissociated by means of electrical pulse or buffer elution strategies, which realizes signal regeneration after detection and supports continuous dynamic monitoring, which is superior to the existing disposable detection architecture.

[0034] (4) This invention can be used for the detection of a variety of protein molecules, including IFN-γ tumor markers, and has good versatility, biocompatibility and future wearable / wireless expansion capabilities, providing key technical support for the next generation of implantable diagnostic systems. Attached Figure Description

[0035] Figure 1 Photolithography process for patterning organic semiconductor channels.

[0036] Figure 2 This is a schematic diagram of the structure of a flexible injectable organic electrochemical transistor sensor.

[0037] Figure 3 The flowchart shows the biofunctionalization modification process of a metal gate, where (a) represents cleaning the gold electrode, (b) represents modifying the self-assembled monolayer, (c) represents activating the self-assembled monolayer, and (d) represents the binding of an amino-modified single-stranded DNA aptamer to the self-assembled monolayer via covalent bonds.

[0038] Figure 4 This is a top-view photograph of the flexible injectable organic electrochemical transistor sensor prepared in Example 1.

[0039] Figure 5 The graph shows the maximum transconductance values ​​of the flexible injectable organic electrochemical transistor sensors prepared in Examples 1-6.

[0040] Figure 6 The transfer curve (a) and transconductance curve (b) of the flexible injectable organic electrochemical transistor sensor prepared in Example 1 were obtained by testing it in INF-γ solutions with concentrations of 10 pg / mL, 1 ng / mL, 100 ng / mL, and 10 μg / mL.

[0041] Figure 7 The protein response curves of the flexible injectable organic electrochemical transistor sensor prepared in Example 1 were obtained by testing it in INF-γ solutions with concentrations of 10 pg / mL, 1 ng / mL, 100 ng / mL, and 10 μg / mL.

[0042] Figure 8 The transmission curve of the flexible injectable organic electrochemical transistor sensor prepared in Example 1 was obtained by testing in PBS and 10 μg / mL INF-γ solution.

[0043] Figure 9 The results are for injectability testing, including (a) the integrated release sensor syringe, (b) the release and performance testing of the sensor before and after injection from the silicon substrate, (c) the release sensor injected into an in vitro simulated subcutaneous environment using an indwelling needle, and (d) a microscopic image of the sensor released subcutaneously by the syringe needle. Detailed Implementation

[0044] To make the objectives, features, and advantages of this invention more apparent and understandable, a detailed description is provided below through specific embodiments. Many specific details are set forth in the following description to provide a thorough understanding of the invention. However, the invention can be practiced in many other ways different from those described herein, and those skilled in the art can make similar modifications without departing from the spirit of the invention. Therefore, the invention is not limited to the specific embodiments disclosed below. Technical features in various embodiments of the invention can be combined appropriately without mutual conflict.

[0045] Unless otherwise specified, the operating methods in the following examples are generally performed under conventional conditions or as recommended by the manufacturer. Contents not described in detail in this specification are prior art known to those skilled in the art. Unless otherwise specified, the experimental materials used in the examples below can be purchased from conventional biochemical reagent companies.

[0046] Example 1 (1) Polyimide liquid was spin-coated onto a silicon oxide wafer and vacuum dried overnight at 200°C to obtain a flexible substrate with a thickness of 3.5 μm.

[0047] (2) Electrode patterning adopts the Lift-off process. Negative photoresist is spin-coated on the surface of a flexible substrate at a speed of 2.5krpm for 30s, baked and cured, and exposed and developed to form a photoresist structure with a preset pattern. Then, 5nm Cr and 80nm Au metal are deposited sequentially on its surface by thermal evaporation. After that, the device is immersed in acetone solution for 30-120 min to clean it, and then peeled off to obtain a polyimide substrate with metal source, metal drain and metal gate.

[0048] (3) Prepare a PEDOT:PSS composite solution, wherein the content of PEDOT:PSS is 93.75 wt%, the content of ethylene glycol is 5 wt%, the content of 3-glycidoxypropyltrimethoxysilane is 1 wt%, and the content of dodecylbenzenesulfonic acid is 0.25 wt%.

[0049] (4) Organic semiconductor channel patterning adopts the lift-off process, and the processing is as follows: Figure 1 As shown, a 1.6-1.7 μm thick AZ series photoresist is spin-coated onto the surface of a polyimide substrate with completed electrode patterning, leaving only the channel area uncovered by the photoresist. After pre-baking, exposure and development, a semiconductor channel patterned substrate is obtained. Subsequently, the PEDOT:PSS composite solution prepared in step (3) is spin-coated onto the entire semiconductor channel patterned substrate, heated and cured, and the device is immersed in acetone to remove the photoresist, resulting in an organic semiconductor channel layer with a thickness of 90-110 nm, a channel width of 50 μm, a channel width-to-length ratio of 2:1, and a resistance of no more than kiloohms.

[0050] (5) Prepare a 2 μm thick encapsulation layer by spin coating, pre-baking, exposure, post-baking and development of SU-8 solution to ensure that the organic semiconductor conductive channel region, the metal gate region and the lead-out port connecting the electrode to the external circuit are kept exposed.

[0051] (6) Figure 2 The diagram shows the structure of the sensor device, which includes a bottom flexible substrate, a metal source, a metal drain, and a metal gate disposed on the surface of the flexible substrate, and a top encapsulation layer. An organic semiconductor conductive channel is disposed between the metal source and the metal drain, and the metal gate is arranged coplanarly with the organic semiconductor conductive channel.

[0052] (7) Biofunctionalization modification process as follows Figure 3As shown, the sensor device obtained in step (6) is immersed in 0.5M sulfuric acid and chemically cleaned by 20 cycles of cyclic voltammetry with a voltage range of -0.4-1.4V, while avoiding the organic semiconductor channel region.

[0053] (8) After cleaning, the device was immersed in an ethanol solution of 11-mercaptoundecanoic acid at a concentration of 1.1 mg / mL and incubated at room temperature for 16 hours to allow 11-MUA molecules to form a self-assembled monolayer on the metal gate surface. Subsequently, the device was transferred to EDC / NHS activation buffer (EDC 200mM, NHS 50mM) and reacted for 2 hours.

[0054] (9) The aptamer molecule solution was prepared using Tris-HCl buffer at a concentration of 10 mM and a pH of 7.4-8.0 to provide a stable, weakly alkaline environment to maintain the conformational stability of the aptamer molecule. Specifically, the lyophilized aptamer molecule (5′-NH2-TAATTCCCAATCCATGTGTTGTGGGTTGTGTTGGTTGGGG-3′, SEQ ID NO.1) was dissolved in the above Tris-HCl buffer at a concentration of 5 μM. After preparation, the aptamer solution can be stored for a short period at 4°C or aliquoted and frozen at -20°C to avoid repeated freeze-thaw cycles affecting its activity.

[0055] (10) Immerse the activated sensor device from step (8) in the aptamer buffer solution obtained in step (9), react for 12 hours, and then dry to obtain the flexible injectable organic electrochemical transistor sensor. Figure 4 This is a top view of the flexible injectable organic electrochemical transistor sensor prepared in Example 1, wherein the width W of the organic semiconductor conductive channel is 50 μm, the length L is 25 μm, and the widest part of the flexible injectable organic electrochemical transistor sensor is 350 μm.

[0056] Example 2 The only difference between this embodiment and Embodiment 1 is that the channel width of the organic semiconductor channel layer is 150 μm, the channel width-to-length ratio is 2:1, the width and length of other parts are proportionally enlarged, the thickness remains unchanged, and the maximum size of the flexible injectable organic electrochemical transistor sensor is 1050 μm.

[0057] Example 3 The only difference between this embodiment and Embodiment 1 is that the channel width of the organic semiconductor channel layer is 700 μm, the channel width-to-length ratio is 2:1, the width and length of other parts are proportionally enlarged, the thickness remains unchanged, and the maximum size of the flexible injectable organic electrochemical transistor sensor is 4900 μm.

[0058] Example 4 The only difference between this embodiment and Embodiment 1 is that the channel width-to-length ratio of the organic semiconductor channel layer is 1:1; the other parts are the same.

[0059] Example 5 The only difference between this embodiment and Embodiment 1 is that the channel width of the organic semiconductor channel layer is 150 μm, the channel width-to-length ratio is 1:1, the width and length of other parts are proportionally enlarged, the thickness remains unchanged, and the maximum size of the flexible injectable organic electrochemical transistor sensor is 1050 μm.

[0060] Example 6 The only difference between this embodiment and Embodiment 1 is that the channel width of the organic semiconductor channel layer is 700 μm, the channel width-to-length ratio is 1:1, the width and length of other parts are proportionally enlarged, the thickness remains unchanged, and the maximum size of the flexible injectable organic electrochemical transistor sensor is 4900 μm.

[0061] Application Example 1 INF-γ protein was dissolved in PBS buffer at concentrations of 10 pg / mL, 1 ng / mL, 100 ng / mL and 10 μg / mL, respectively. The mixture was added to a hollow cylindrical cell with an inner diameter of 5 mm and a depth of 4 mm, which was made of resin material obtained by 3D printing, to obtain the detection container.

[0062] The flexible injectable organic electrochemical transistor sensors prepared in Examples 1-6 were fixed on a probe stage. Three electrodes were connected to a Keithley semiconductor parameter analyzer via a microprobe. 100 μL of INF-γ protein solutions of different concentrations (10 pg / mL, 1 ng / mL, 100 ng / mL, and 10 μg / mL) were added dropwise to the chamber in the gate region. The drain-source voltage (V...) was then... DS The gate-source voltage (V) is fixed at -0.4V. GS The drain current (I) is measured by gradually increasing the voltage from -0.2V to +1.0V. DS The reaction time was set to 8 minutes. Transport characteristic curves of the flexible injectable organic electrochemical transistor sensors prepared in Examples 1-6 to the INF-γ protein were obtained. The obtained transport curves were compared with I... DS -V GS The curve is numerically differentiated to calculate the transconductance g of the device, which reflects the device's signal amplification capability. By comparing different devices at V... GS The maximum transconductance value in the range from -0.2V to +1V, such as Figure 5 As shown, under the same device size, the larger the channel width-to-length ratio, the higher the transconductance performance of the device. Under the same channel width-to-length ratio, the smaller the overall size of the device, the better its performance and the better its signal amplification capability.

[0063] Sample Analysis Figure 6 In Figures (a) and (b), the transfer and transconductance curves of the flexible injectable organic electrochemical transistor sensor prepared in Example 1 were obtained by testing under different concentrations (10 pg / mL, 1 ng / mL, 100 ng / mL, 10 μg / mL) of INF-γ solution. The curve measured in PBS solution (without the target protein) was used as the baseline to calculate the rate of change of current Δ. I D ,like Figure 6 As shown in (a) of the figure, under constant drain voltage conditions, the transfer curve of the device shifts upward as the concentration of IFN-γ increases, indicating that the drain current increases with increasing concentration. Figure 6 As shown in (b), the transconductance curve has a maximum value under PBS conditions. With increasing IFN-γ concentration, the transconductance peak gradually decreases, and at V... GS The variation is most significant in the range of approximately 0.4–0.6 V. These results demonstrate that the device of the present invention can achieve sensitive detection of IFN-γ concentration and exhibits a stable and distinguishable electrical response within a concentration range of 10 pg / mL–10 µg / mL.

[0064] like Figure 7 As shown, the protein concentration is taken as the logarithm to base 10 (log 10 C) Plot the device's response curve using the x-axis as the abscissa and the rate of change of current as the y-axis. Perform linear fitting on the obtained data to obtain the fitting equation: y = 0.31x + 0.05, with a goodness of fit of R0. 2 =0.99, indicating that the device responds well and exhibits significant linearity within this concentration range. Measurements were repeated three times under each concentration condition, demonstrating good repeatability. Therefore, the device's limit of detection for INF-γ protein can reach 10 pg / mL.

[0065] Figure 8 The transfer curve of the flexible injectable organic electrochemical transistor sensor prepared in Example 1 was obtained by testing in PBS and 10 μg / mL INF-γ solution. Under constant drain voltage conditions, the transfer curve of the device shifted upward as the concentration of IFN-γ increased, which is consistent with... Figure 6 The pattern described in the paper is consistent with that in the paper. Then, pulse voltages of different amplitudes at 100Hz were applied to the gate, and the transfer curve of the device shifted downward. Therefore, the application of electrical pulses can assist protein elution to a certain extent and promote signal regeneration after detection, so as to enable continuous dynamic monitoring function.

[0066] like Figure 9As shown in (a), PVA was used as an adhesive to fix the device released from the silicon substrate to the outside of the syringe to evaluate the injectability of the flexible injectable organic electrochemical body tube sensor prepared in Example 1 and its performance stability before and after release from the silicon substrate and injection. The sensor performance before and after release from the silicon substrate and injection was systematically tested, and the results are as follows. Figure 9 As shown in (b), the release and injection process from the silicon substrate has no significant impact on sensor performance. Furthermore, Figure 9 (c) and (d) in the figure demonstrate the process of sensor injection and release from the needle in PDMS simulated skin tissue bonded to a glass slide using an indwelling needle device, thereby verifying the injectability of the flexible injectable organic electrochemical transistor sensor.

[0067] In the reusability test of the device, the experiment confirmed that after more than 20 incubation / test / rinse cycles (~8h), the device still retains more than 95% of its original performance.

[0068] The embodiments described above provide a detailed explanation of the technical solutions of the present invention. It should be understood that the above descriptions are merely specific embodiments of the present invention and are not intended to limit the present invention. Any modifications, additions, or similar substitutions made within the scope of the principles of the present invention should be included within the protection scope of the present invention.

Claims

1. A flexible injectable organic electrochemical transistor sensor for the detection of protein immunomarkers, characterized in that, It includes a bottom flexible substrate, a metal source, a metal drain and a metal gate disposed on the surface of the flexible substrate, and a top encapsulation layer. An organic semiconductor conductive channel is disposed between the metal source and the metal drain. The metal gate and the organic semiconductor conductive channel are arranged on the same plane. The top encapsulation layer plays an encapsulation role and the organic semiconductor conductive channel and the metal gate are exposed after encapsulation. The metal gate is modified with a monolayer formed by the self-assembly of 11-mercaptoundecanoic acid, and an amino-modified single-stranded DNA aptamer is covalently coupled to the monolayer.

2. The flexible injectable organic electrochemical transistor sensor according to claim 1, characterized in that, The flexible substrate is made of polyimide, SU-8, polyethylene terephthalate, or polydimethylsiloxane; The organic semiconductor conductive channel was obtained by curing an organic semiconductor composite solution consisting of PEDOT:PSS, ethylene glycol, 3-glycidoxypropyltrimethoxysilane and dodecylbenzenesulfonic acid. In the organic semiconductor composite solution, the content of PEDOT:PSS is 90-96 wt%, the content of ethylene glycol is 2-8 wt%, the content of 3-glycidoxypropyltrimethoxysilane is 0.1-2 wt%, and the content of dodecylbenzenesulfonic acid is 0.01-0.50 wt%.

3. The flexible injectable organic electrochemical transistor sensor according to claim 1, characterized in that, After a flexible injectable organic electrochemical transistor sensor binds to a target protein-based immunomarker, the conjugate is dissociated using an electro-pulse or buffer elution strategy, thereby regenerating the signal after sensor detection.

4. The flexible injectable organic electrochemical transistor sensor according to claim 1, characterized in that, The protein-based immunomarker is IFN-γ, and the amino-modified single-stranded DNA aptamer is 5′-NH2-TAATTCCCAATCCATGTGTTGTGGGTTGTGTTGGTTGGGG-3′ with 5′ amino-modified end.

5. The flexible injectable organic electrochemical transistor sensor according to claim 1, characterized in that, The flexible substrate has a thickness of 3-4 μm; the metal source, metal drain, and metal gate all have a thickness of 80-140 nm; the encapsulation layer has a thickness of 1-3 μm; the organic semiconductor conductive channel has a thickness of 90-110 nm, a width of 50-700 μm, and a width-to-length ratio of 1-3:

1.

6. The flexible injectable organic electrochemical transistor sensor according to claim 1, characterized in that, The flexible injectable organic electrochemical transistor sensor has a maximum size of 350-4900 μm. It is used by rolling up the flexible injectable organic electrochemical transistor sensor and then injecting it with a syringe. The flexible injectable organic electrochemical transistor sensor has outward-leading signal lines.

7. The method for fabricating a flexible injectable organic electrochemical transistor sensor according to any one of claims 1-6, characterized in that, include: (1) The metal source, metal drain and metal gate are deposited on the surface of a flexible substrate by evaporation and photolithography. (2) The organic semiconductor composite solution is spin-coated, cured and patterned to obtain an organic semiconductor channel layer, which is then placed between the metal source and the metal drain to obtain a sensor component; (3) Encapsulate the sensor components to expose the encapsulated organic semiconductor conductive channels and metal gates, clean them, and incubate them in an ethanol solution of 11-mercaptoundecanoic acid for ≥16 h. After complete incubation, activate them for 1-5 h, then immerse them in an amino-modified single-stranded DNA aptamer solution for ≥12 h, and dry them to obtain the flexible injectable organic electrochemical transistor sensor.

8. The method for fabricating a flexible injectable organic electrochemical transistor sensor according to claim 7, characterized in that, The concentration of the ethanol solution of 11-mercaptoundecanoic acid is 1-1.5 mg / mL.

9. The method for fabricating a flexible injectable organic electrochemical transistor sensor according to claim 7, characterized in that, The activation process used an EDC / NHS activation buffer, in which the concentration of EDC was 150-250 mM and the concentration of NHS was 40-60 mM; the concentration of the amino-modified single-stranded DNA aptamer solution was 5-10 μM.

10. A method for detecting IFN-γ for non-disease diagnosis or non-treatment purposes, characterized in that, Use the flexible injectable organic electrochemical transistor sensor according to any one of claims 1-6.

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