Fully integrated and minimally invasive continuous blood glucose monitoring system (CGM) with coin size based on organic electrochemical transistors

Abstract

The invention provides a coin-sized, fully integrated and wearable continuous blood glucose monitoring (CGM) system in combination with frontier technologies in crossing fields of biosensors, minimally invasive tools, hydrogels and the like. The invention comprises the following three main parts: 1) an emerging biochemical amplifier, namely an organic electrochemical transistor (OECT), for improving the sensitivity and exceeding the traditional electrochemical sensor; 2) a microneedle array for sampling interstitial fluid (ISF), which reduces pain in the skin penetration process; and 3) a tough and viscous enzyme hydrogel film which can improve the reliability of glucose sensing on the skin. Different from a traditional CGM, the CGM (OECT-CGM) provided by the invention has high anti-noise capability, adjustable sensitivity and current regeneration capability due to the adoption of the OECT amplifier, so that long-term stable glucose sensing can be realized within a specific clinical range (1-20mM). The CGM provided by the invention can provide high and adjustable sensitivity, minimally invasive property and better wearable property at the same time, and also paves a way for developing the next generation of CGM.

Description

Cross - Reference to Related Applications

[0001] This application claims the priority of U.S. Provisional Application No. 63 / 617,436, filed on January 4, 2024, the entire disclosure of each of which is incorporated herein by reference in its entirety. Technical Field

[0002] The present invention generally relates to the technical field of wearable continuous glucose monitoring systems. More specifically, the present invention relates to a continuous glucose monitoring system based on organic electrochemical transistors, which has a lower detection limit, high resolution, high anti - noise ability, automatic calibration, and current regeneration ability. Background Art

[0003] Diabetes is one of the malignant chronic diseases threatening human health. Improper intervention in diabetic patients may cause sudden hypoglycemia and increase the risk of complications. Therefore, accurately detecting blood glucose levels is crucial. In the past two decades, continuous glucose monitoring systems (CGMS) have made great progress. CGMS can inform, notify, and warn diabetic patients of persistent hyperglycemia and sudden hypoglycemia, providing necessary information for closed - loop blood glucose control systems. Wearable CGM systems are considered promising candidate systems for providing personalized lifestyle advice through all - weather continuous blood glucose monitoring. However, current wearable CGM devices still have problems such as pain during sensor implantation and maintaining high accuracy during continuous blood glucose level tracking. Therefore, developing the next - generation CGM with higher sensitivity, minimally invasive nature, and comfortable wearability has become the main research focus of all sectors of society to promote diabetes management in different cases.

[0004] Organic electrochemical transistors (OECTs) have recently emerged as one of the most promising next - generation biosensor technologies because they combine the advantages of electrochemical transducers and transistor amplifiers. OECTs based on ion - electron hybrid semiconductors, such as poly(3,4 - ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS), can operate at low power in a liquid environment at low voltage (<1V) while maintaining stable performance for several months. The operation of OECTs relies on the electrochemical doping / dedoping process of the ion - electron hybrid semiconductor channel in contact with the electrolyte, where mobile ions can penetrate between the channels and then adjust the entire channel volume. These characteristics enable OECT - based biosensors to in - situ amplify weak electrochemical signals with a high signal - to - noise ratio (SNR) at low voltage and low power. Therefore, OECT technology can effectively solve the inherent low - power / high - SNR trade - off problem of biomedical sensors.

[0005] So far, the research on OECT-based glucose biosensors has mainly focused on enhancing the performance metrics of the sensors, such as amplification ability (i.e., transconductance), mechanical stretchability, and operational stability. The following have been reported in various reports and academic articles: OECT-based glucose sensors that work in a neutral pH environment; high-selectivity glucose sensors can be obtained by modifying the gate electrode of OECT with carbon nanotubes (CNT); OECT-based glucose sensors with micromolar sensitivity, where the gate electrode is coated with a gel containing glucose oxidase (GOx); stretchable OECTs for glucose detection, and the device can still work properly under a strain of up to 30%. A recent paper reported a high-transconductance OECT that achieved highly sensitive glucose detection with a transconductance of 180 mS using an interdigitated electrode structure.

[0006] Although OECT glucose biosensors have been widely studied, due to the lack of a system-level development strategy, their practical applications in the wearable CGM scenario have not been explored. To maximize their competitiveness as the next-generation CGM, OECT glucose sensors need to be integrated with a miniaturized readout system to achieve comfortable wearability, as well as minimally invasive sampling techniques to reduce pain during skin penetration and minimize the risks induced during wearing. In addition, strategies to ensure skin interface stability and reliability during long-term use are also an essential part of filling the missing part of the system challenges.

[0007] In the present invention, an intrinsically stretchable glucose-sensing enzyme membrane was developed by combining: i) a hybrid conductive polymer hydrogel and ii) a stretchable redox hydrogel loaded with enzymes by a covalent bonding method. Specifically, a 3D interpenetrating structure was provided, where the strategic use of the PEDOT:PSS conductive network reduced the distance between the enzyme and the electron-conductive network (i.e., the electrode). This method promoted direct electron transfer (DET) and mediated electron transfer (MET), representing a significant advancement in stretchable molecular biosensing enzyme technology.

[0008] The present invention has the potential to influence a new research direction - the development of stretchable enzyme technology. This technology is customized for emerging wearable and tissue-compatible biosensing applications and will significantly improve the strain robustness of membranes, electrodes, and devices, thus promoting their wider deployment at soft biointerfaces. For example, it can be a potential supplement to the current three main glucose oxidase technologies. Summary of the Invention

[0009] In response to the above challenges and problems, the present invention hereby provides a prototype of a wearable OECT-CGM system for diabetes medical management.

[0010] In one aspect of the present invention, a continuous blood glucose monitoring system based on an organic electrochemical transistor has minimal invasiveness and no blood contact, and includes a hollow microneedle patch with surface-deposited metal, a sticky stretchable enzyme hydrogel sensing film, a glucose sensor based on an organic electrochemical transistor, and a 3D printed resin encapsulation housing. Specifically, the organic electrochemical transistor can perform current regeneration, sensitivity adjustment, and self-calibration.

[0011] In an embodiment, the sticky stretchable enzyme hydrogel sensing film includes an interpenetrating polymer network hydrogel, and the interpenetrating polymer network hydrogel includes a first network of sodium alginate and a second network of polyacrylamide, a semipermeable bioadhesive elastomer, and glucose oxidase.

[0012] In other embodiments, the vapor-deposited metal coating on the 3D printed plastic encapsulation housing is selected from gold (Au) or platinum (Pt).

[0013] In another embodiment, the continuous blood glucose monitoring system based on an organic electrochemical transistor can track cellular glucose metabolism and clinical blood glucose concentration in the blood glucose concentration range of 10 -6 M to 10 -1 M.

[0014] In another embodiment, the size of the continuous blood glucose monitoring system based on an organic electrochemical transistor of the present invention is less than 2 cm × 2 cm × 0.5 cm.

[0015] In another embodiment, the continuous blood glucose monitoring system based on an organic electrochemical transistor of the present invention achieves a high signal-to-noise ratio of at least 50 dB.

[0016] In another embodiment, the current regeneration of the organic electrochemical transistor of the system includes adjusting the gate voltage of the organic electrochemical transistor.

[0017] In another embodiment, the sensitivity adjustment of the organic electrochemical transistor of the system includes controlling the transconductance to adjust the anti-noise ability; adjusting the gate voltage to adjust the linear range of detection.

[0018] In another embodiment, the self-calibration of the organic electrochemical transistor of the system includes standardizing the transconductance curve. BRIEF DESCRIPTION OF THE DRAWINGS

[0019] In order to more clearly illustrate the technical solutions of the embodiments of the present invention, the drawings required for the description of the embodiments are briefly introduced below. Obviously, the drawings in the following description are only some embodiments of the present invention, and those skilled in the art can clearly obtain other drawings according to these drawings without creative labor, where:

[0020] Figure 1Shows the concept and design principle of the OECT-CGM system. The system includes: (i) a microneedle patch for connecting the ISF and the OECT sensor, (ii) a tough and sticky enzyme hydrogel sensing membrane, (iii) an OECT glucose sensor, (iv) a micro readout system for real-time communication with peripheral devices (such as mobile phones), and (v) a 3D printed plastic encapsulation shell. Using a stacked integration strategy, a size of 6.5 mm in height and 14 mm in radius is achieved.

[0021] Figures 2A - 2I Outlines the fabrication and characteristics of the OECT-based glucose sensor. Figure 2A Is a schematic diagram of the OECT glucose sensor and the sensing mechanism. Figure 2B Shows the optical image of the OECT glucose sensor fabricated on the PI substrate. Figure 2C Displays the output curve of the flexible OECT, where V ds Is scanned from 0 V to -0.8 V with a step size of 20 mV. Figure 2D Is the transfer curve of the OECT, where Vg is scanned from -0.1 V to 0.8 V with a step size of 20 mV and repeated for 40 cycles, indicating the stability of the OECT for biosensing. Figure 2E Is the real-time current response (ΔI -6 mM changes to 10 -1 mM) of the OECT-based glucose sensor when the glucose concentration changes from 10 ds . Figure 2F Shows the response of the OECT glucose sensor to potential interfering substances (including ascorbic acid and urea), indicating the selectivity of the sensor. Figure 2G Shows the real-time current response (ΔI -4 mM changes to 10 -2 mM) of the OECT-based glucose sensor when the glucose concentration changes from 10 ds at different environmental noise levels. Figure 2H Shows that at different environmental noise levels, the electrochemical glucose sensor responds to glucose concentration changes from 10 -6 mM to 10 -1The real-time current response (ΔI, working electrode current) in mM. The function generator generates white noise with three different amplitudes (10.0 μA, 1.0 μA, and 0.2 μA). During the data processing stage, these three different noises are respectively superimposed on the original current response curves of the OECT-type glucose sensor and the electrochemical-type glucose sensor. Even when the applied noise amplitude reaches 10.0 μA, the current response curves of the OECT-type sensor are still clearly distinguishable at different glucose concentrations. However, when the applied noise amplitude reaches 1.0 μA, the current response curves of the electrochemical sensor are already overwhelmed by the noise. This demonstrates that the OECT-type sensor has good anti-transmission noise ability. 2I shows the SNR of the OECT glucose sensor and the electrochemical-based glucose sensor under the application of artificial noise (n = 10). The data are presented as mean ± SD.

[0022] Figures 3A - 3H Summarizes the integrated OECT glucose sensing platform with tunable sensitivity, current regeneration ability, and self-calibration ability. Figure 3A Shows the optical image of the PERfECT readout system customized for wearable OECT characteristics. Figure 3B Is the circuit diagram and main components of the PERfECT system. Figure 3C Shows the comparison of the glucose responses (from 2 mM to 14 mM, step size 2 mM) of the OECT sensor measured by PERfECT and the SMU used in the laboratory. Figure 3D Shows the OECT glucose sensor at different V g Response to different glucose levels (from 2 mM to 18 mM). Figure 3E Shows the linear fitting of the results in the figure. Figure 3D Shows the tunable sensitivity achieved by controlling the Vg of the OECT to 0.2 V and 0.3 V respectively. Figure 3F Shows the current response (before calibration) of the OECT glucose sensor when the glucose concentration increases to 10 mM. Figure 3G Shows the calibrated current response (after calibration) of the OECT glucose sensor when the glucose concentration increases to 10 mM. Figure 3H Shows the calibration process of the OECT glucose sensor using a specific transfer curve and related g m

[0023] Figures 4A - 4H Outlines the prototype design and in vivo testing of a fully integrated, minimally invasive wearable OECT-CGM system. Figure 4A ​Optical images of the wearable OECT-CGM on the skin and 3D-printed microneedle array (left); Glucose diffusion mechanism between the microneedles and ISF of the IPN hydrogel (middle); and Scanning electron microscope (SEM) images of the hollow microneedle array (side view and top view) (right). Figure 4B Optical images showing a sticky and tough double-network IPN hydrogel composed of PAAm hydrogel and sodium alginate hydrogel (left) and the peeling process of the sticky and tough hydrogel (right). Figure 4C Shows the time lag comparison of OECT sensors with / without hydrogel fillers. Figure 4D Shows the size comparison between the OECT-CGM and the Dexcom glucose sensor (G6). Figure 4E Illustrates the comparison of real-time sensing results between the wearable OECT-CGM system and the Dexcom G6. Figure 4F Illustrates the MARD (mean absolute relative error) of the OECT-CGM (calculated from 12 sampling points within 1000 seconds). Figure 4G Illustrates the in vivo test setup image of the OECT-CGM for tracking blood glucose levels using a rat model. Figure 4H Illustrates the in vivo glucose measurement results obtained using the OECT-CGM and a blood glucose meter (black, 16 data points of glucose concentration collected by the blood glucose meter; red, relative channel current change of the OECT-based glucose sensor measured by the OECT-CGM). (a.u. arbitrary unit.)

[0024] Figure 5 Illustrates the exploded view of the OECT-CGM system. The OECT-CGM system is assembled by integrating four main components: i) a miniature personalized electronic reader for the electrochemical transistor (PERfECT), ii) an OECT-based glucose sensor for converting changes in glucose concentration into changes in the current Ids flowing in the semiconductor channel, ii) a hollow microneedle array for interstitial fluid (ISF) glucose collection, with a length of 0.5 mm to achieve minimally invasive, iv) an enzymatic, tough, and adhesive hydrogel layer with a thickness of 0.15 mm to improve adhesion and robustness under deformation, effectively connecting the ISF and the sensor to achieve accurate and stable glucose monitoring. v) A replaceable 2036 independent power supply for the entire system. vi) A 3D-printed resin housing encapsulates the entire system.

[0025] Figure 6 Illustrates the current response of the OECT glucose sensor. The sensor was first immersed in PBS solution. After adding 2 mL of 10 mM and 20 mM glucose solutions to the test container at time = 46 s and time = 88 s respectively, the current in the OECT glucose sensor channel decreased to 697 μA and 680 μA, Vg = 0.2 V.

[0026] Figure 7 Shows the signal amplification of the OECT glucose sensor. The power amplification of the biological signal in the OECT mainly comes from the integration of the channel to the gate-to-channel current (I g ). The product of the gate voltage (V g ) and the gate current (I g ) represents the input power of the signal. The product of the channel voltage (V ds ) and the change in the channel current (ΔI ds ) represents the output power. The power amplification of the OECT glucose sensor is 19.66 dB. n = 10. Data are mean ± SD.

[0027] Figure 8 Shows the optical image of the microneedle patch prototype with the wearable PERfECT readout system. The overall size of the device is moderate, with a length and width of 15 mm each and a thickness of 2.5 mm. This compact size facilitates portability and discreet use. We incorporated microneedles into the design to provide a minimally invasive experience for users while ensuring accurate blood glucose level measurement.

[0028] Figure 9 Shows the optical image of the microneedle array prototype with a metallized surface. The microneedle array has a 10×10 configuration, and its sharp conical microneedles are set at a 15-degree angle to achieve the best skin penetration effect. Its size is 10 mm × 10 mm × 2.2 mm, and it can be seamlessly aligned with the CGM readout system. The manufacturing process starts with precision 3D printing to achieve the required geometry. The next step involves magnetron sputtering to metallize the surface with a layer of gold (30 nm). This metallization enhances the structural strength, ensures durability, and helps with a smoother application.

[0029] Figure 10A Is a double-network hydrogel layer diffusion colorimetric test. Before gelation, the double-network hydrogel was applied to the 3D printed hollow microneedle structure. The microneedle array consists of 12 hollow needles with a pore diameter of 150 μm and a length of 1 mm, and the needles are filled with the double-network hydrogel. The thickness of the hydrogel layer on the substrate is approximately 300 μm. At Time = 0 s, 50 μL of 0.1% methyl orange solution was dropped on the tip of the microneedle patch, and stained hydrogel photos were taken at Time = 50 s and Time = 100 s, indicating the diffusion efficiency of the double-network hydrogel as a diffusion layer. Figure 10BpH diffusion tests in solution and hydrogel-filled microneedles. The pH values of different solutions with a pH range of 2 - 7 were measured using a pH meter. The purple dots represent the pH values measured directly in the solution, and the blue dots represent the pH values measured through microneedles (with a thickness of approximately 300 μm) filled with a double-network hydrogel after adding solutions of other pH values. The pH values measured after 200 s of stabilization indicate the diffusion efficiency of the double-network hydrogel within the microneedles.

[0030] Figure 11A The Bode plot (left) and Warburg resistance - angular frequency plot (right) in the hydrogel are shown respectively. Figure 11B The Bode plot (left) and Warburg resistance - angular frequency plot (right) in water are shown respectively. Substituting this value into the above equation, the diffusion coefficients of glucose in the hydrogel and in water can be estimated as: Calculations show that the diffusion efficiency of glucose molecules in water and in the double-network hydrogel is similar.

[0031] Figure 12 Scanning electron microscope images of the IPN hydrogel filling layer are shown. The IPN hydrogel layer is used as the filling material inside the hollow microneedles, aiming to establish a glucose diffusion pathway for the ISF to OECT glucose sensing system. The scanning electron microscope images illustrate the porous nature of the IPN hydrogel after the freeze-drying process. These pores are crucial for promoting the diffusion of glucose molecules, which is a key mechanism necessary for the operation of the sensor.

[0032] Figure 13 An in vitro experimental setup for glucose monitoring in artificial ISF is shown. The device used for in vitro testing of changes in the concentration of artificial interstitial fluid includes: i) an artificial ISF reservoir containing a glucose solution on which the entire OECT-CGM can be stably placed, ii) a programmable injection pump for pushing glucose solutions of different concentrations into the reservoir at a constant speed. iii) An artificial skin based on an agarose film. By setting the injection pump to different forward speeds and filling the syringe with glucose solutions of gradient concentrations, this experimental setup simulates different changes in the glucose concentration in the ISF.

[0033] Figure 14 The stability test of the OECT-based glucose sensor over a two-week monitoring period is shown, depicting the long-term stability test (14 days) of the OECT glucose sensor. This figure shows the response of the glucose sensor to a glucose solution of constant concentration (6 mM) at different time points, measured by the channel current variable (ΔI ds ) Measurement. ΔI dsValues were normalized for comparison. It is evident that the current gradually decreased over time and reached approximately 90% of the initial value by day 14. The decrease in current indicates a reduced response of the sensor to the same glucose concentration during the test, suggesting a decrease in long-term stability. n = 10. Data are mean ± SD.

[0034] Figure 15 This is a photograph of in vivo continuous blood glucose measurement using the OECT-CGM system. This image outlines the in vivo test setup used to validate the performance of the OECT-CGM system in real-time blood glucose concentration monitoring. The back area of the rat was carefully prepared by shaving to enable precise insertion of the microneedles into the skin and to ensure optimal adhesion of the IPN hydrogel, thus enabling effective application of the OECT-CGM. The highlighted area within the red dashed box represents the OECT-CGM device. In this configuration, glucose in the rat interstitial fluid diffuses passively and naturally through the hydrogel located within the hollow microneedles towards the OECT glucose sensor. Subsequently, the sensor data are acquired by the miniature PERfECT readout system and wirelessly transmitted to a mobile phone using low-power Bluetooth (BLE) technology, where the data are recorded and analyzed. To verify the blood glucose level of the rat, blood samples were regularly collected from the rat tail vein every 15 minutes for 2 hours and tested using a blood glucose meter. Detailed Description

[0035] Diabetes, especially type 2 diabetes, can have its progression significantly slowed or controlled in the early stages. Therefore, regular blood glucose monitoring is crucial for helping diabetic patients mitigate the impact of diabetes on their health.

[0036] Traditional blood glucose monitoring is performed through invasive blood sampling techniques, such as pricking a finger to draw blood for blood glucose monitoring, which can be inconvenient and may cause minor pain to patients. Therefore, minimally invasive blood glucose monitoring methods are an increasingly popular alternative.

[0037] With technological advancements, it is now possible to detect blood glucose levels without a blood sample; instead, interstitial fluid is sufficient to detect blood glucose levels. Therefore, microneedle patches have been developed, which greatly reduce the invasiveness of the monitoring procedure while also enabling continuous glucose monitoring (CGM), thus providing real-time data and trends for patients and healthcare providers to make timely adjustments based on blood glucose level fluctuations.

[0038] However, the existing technologies of sensors designed for interstitial fluid glucose readings in the art still have accuracy and calibration problems. Due to the presence of other substances in the interstitial fluid that may interfere with glucose measurement, sensors with higher precision and calibration capabilities are needed. The continuous blood glucose monitoring system based on organic electrochemical transistors of the present invention solves these problems by adjusting the gate voltage and transconductance of the transistor, integrating current regeneration, sensitivity adjustment and self-calibration capabilities, and ensuring the accuracy and reliability of blood glucose level readings.

[0039] In addition, skin conditions such as skin allergies, scars and peripheral edema that may be caused by some more invasive detection techniques may affect the accuracy of the readings. Therefore, the hydrogel-based thin film and microneedle patch continuous blood glucose monitoring system of the present invention is designed to minimize the irritation and trauma to the skin, thereby preventing the occurrence of these skin adverse reactions, not only ensuring good accuracy of blood glucose level readings, but also reducing the discomfort of patients.

[0040] As will be described in more detail below, the continuous blood glucose monitoring system of the present invention can not only minimize the trauma to the skin, thereby optimally reducing the discomfort of patients, but also, by integrating OECT into the system, the synergistic effect of the electrochemistry and the transistor amplifier enhances the sensing quality of blood glucose levels and the tunability of sensitivity. At the same time, as Figure 2I shown, the signal-to-noise ratio is greatly increased to as high as 60 dB, which is nearly 5 times higher than that of traditional electrochemistry-based blood glucose sensors. Examples

[0041] Example 1 - Development and prototyping of the OECT-CGM system

[0042] 1.1 Design principle of the OECT-CGM system

[0043] The coin-sized OECT-CGM system ( Figure 1 and 5 ) consists of five units: (i) a hollow microneedle patch, (ii) a tough and sticky enzyme hydrogel sensing membrane, (iii) an OECT-based glucose sensor, (iv) a micro readout system, and (v) a 3D printed plastic encapsulation box. The microneedle patch serves as a minimally invasive channel for connecting the ISF and the OECT sensor. The stretchable and sticky enzyme hydrogel sensing membrane is sandwiched between the microneedle patch and the OECT glucose sensor to improve interface adhesion and buffer motion artifacts. Due to the concentration gradient, glucose molecules in the ISF passively diffuse to the surface of the OECT biosensor (through the microneedles and the hydrogel). By constructing polyacrylamide (PAAm) and Na +- An interpenetrating double network structure of alginate was used to synthesize a viscous enzyme hydrogel film. The synthesized hydrogel was further cross-linked with glucose oxidase (GOx) for glucose detection. The current of the OECT biosensor was recorded by a miniature wireless readout system ( Figure 1 ) for data communication with a mobile phone. A flexible printed circuit board (fPCB) connector was used as the power supply line to facilitate the easy connection between the OECT sensor and the readout system. A stacking strategy was adopted to reduce the physical size (width and length) of the OECT-CGM system, thereby improving its wearability ( Figure 5 ).

[0044] 1.2 Fabrication and Characterization of OECT Glucose Sensors

[0045] The OECT glucose sensor was configured as shown in Figure 2A and 2B . The OECT glucose sensor was fabricated on a flexible polyimide (PI) substrate (14 mm x 7 mm) with a circular gate having a diameter of 3 mm and a channel with an aspect ratio of 1000 μm / 100 μm.

[0046] The gate electrode was strategically placed on the back side of the PI substrate ( Figure 2B ). This design is beneficial for enzyme modification of the gate electrode and avoids adverse effects on the channel. The OECT showed a typical transistor output curve operating in the depletion mode ( Figure 2C ). The operation stability of the device was verified by cyclic transfer curves. Negligible offset was observed after cyclic measurements ( Figure 2D ), which highlights the remarkable stability of the sensor in biosensing.

[0047] To enhance the sensitivity of the OECT to glucose levels, GO x was mixed in a stretchable interpenetrating double network (IPN) hydrogel film. The sensing mechanism is shown in Figure 2A and includes the following processes: i) Glucose oxidation: When glucose is present, glucose is oxidized by GO x embedded in the enzyme hydrogel, producing gluconolactone and reducing GO x ; ii) Enzyme regeneration: The reduced form of GO x is oxidized by the mediator ferrocene and regenerated to the oxidized form. In turn, the ferrocene mediator is reduced; iii) Mediator oxidation: The reduced form of the ferrocene mediator is then oxidized back to ferrocene on the surface of the gate electrode. The above series of reactions generate a Faraday current (I g ), resulting in a change in the effective gate voltage (V g ). Although this change is small, it is subsequently amplified by the OECT, leading to a significant change in the drain-source current (I ds ) (Figure 6 )。This in-situ signal amplification process enhances the tolerance to circuit noise, resulting in a high signal-to-noise ratio (SNR).

[0048] The fabricated OECT biosensor exhibits a linear response to a wide range of glucose concentrations from 10 -6 mM to 10 -1 mM ( Figure 2E ). Additionally, the results from the interference experiments indicate that due to the high selectivity of GO x , the responses of the sensor to common interfering substances (such as ascorbic acid and urea) can be neglected, and it is confirmed that the influence of these intrinsic substances on the response of OECT (I ds ) is very small ( Figure 2F ). The above results suggest that the OECT glucose biosensor can be used for body fluid analysis in harsh environments. Compared with traditional electrochemistry-based glucose sensors, the OECT-based glucose sensor exhibits enhanced noise immunity due to its inherent high I ds ( Figure 2G and 2H ). The high I ds of the OECT sensor originates from the integration and amplification of the gate signal (I g ), which makes it very practical in situations where signal transmission is vulnerable to environmental noise ( Figure 7 ). Figure 2I A comparative analysis of the SNR of traditional electrochemistry-based glucose sensors and OECT-based glucose sensors is provided, confirming the superiority of the latter.

[0049] 1.3 Tunable sensitivity, current regeneration, and self-calibration ability of the OECT glucose sensor

[0050] Current electrochemistry-based glucose biosensors have long pursued tunable sensitivity, current regeneration, and self-calibration ability to improve their diabetes management functions, although it remains difficult to achieve these functions simultaneously. First, tunable sensitivity allows for focused monitoring of patient-specific glucose ranges (between 1 mM and 20 mM) for precise diabetes management and personalized medication. Second, current regeneration ability is essential for extending the lifespan of the sensor as it can mitigate the decay of the signal over time. Third, self-calibration ability is a key function that simplifies the user experience by eliminating the need for manual calibration before each use. Despite significant progress in this field, the simultaneous achievement of these functions in a single biosensor system has not been realized.

[0051] The OECT glucose biosensor allows these new functions to be achieved simultaneously in one device. To demonstrate these functions, a coin-sized precise PERfECT readout system (with dimensions of 1.5 cm × 1.5 cm × 0.2 cm and a weight of 0.4 g) was used to measure the sensor so that this module can be integrated with smart wearable devices ( Figure 3A and 3B ). The logic diagram of the PERfECT system is as shown in Figure 3B . The potential control module receives commands from the microcontroller unit (MCU) to control V g and V ds , stepping down to 2 mV to precisely characterize the dynamic behavior of the OECT sensor. The current monitoring module can read I ds , with a detection limit as low as 1 nA. Compared with the source measurement unit (SMU) used in the laboratory, PERfECT has a benchmark resolution in applying Vg and V ds and reading I ds ( Figure 3C ).

[0052] Tunable sensitivity is achieved by changing the gate voltage of the OECT. As shown in Figure 3D and E, alternating Vds from 0.3 V to 0.2 V can increase the linear sensitivity from 0.0035 μA·mM -1 to 0.0042 μA·mM -1 . This characteristic can be achieved because the sensitivity (transconductance) of the three-terminal OECT has a linear relationship with V g , which is determined by the following formula (1): where g m is the transconductance, w is the channel width, d is the channel thickness, l is the channel length, μ is the charge carrier mobility, C* is the volume capacitance, V th is the threshold voltage, and V g is the gate voltage. OECTs have the highest μC* among biotransistors, so they have the highest sensitivity and SNR, which stems from the power amplification ability of OECTs as biotransistors ( Figure 7 ).

[0053] Current regeneration is achieved by reducing V g of the OECT, which in turn increases I ds , because the PEDOT:PSS OECT used is a p-type transistor operating in depletion mode. However, as mentioned above, changing V g also affects the sensitivity. Therefore, the self-calibration ability of the system is required. This function is achieved by using the transconductance value for I ds ( Figure 3F-H) is achieved by standardization. As Figure 3G shown, after calibration, the same current response as the glucose level can be obtained (because ΔI ds = Δg m ·ΔV) to compensate for differences between devices. The current regeneration and self-calibration capabilities can avoid current loss caused by the increase in implantation time, which may improve the anti-noise ability and long-term stability during actual use.

[0054] 1.4 Prototype design of the OECT-CGM system

[0055] Microneedles for minimally invasive ISF sampling : One of the main obstacles widely adopted in current CGM systems is the pain during the implantation process. Recent studies have shown that reducing the needle length of microneedles helps to alleviate the pain of human subjects. Microneedles, especially microneedle arrays, can collect biofluid samples with less pain, provided that the needles do not penetrate the skin to the dermal depth where most capillaries and nerve endings are located. Since it does not penetrate to this depth, the biofluid sample collected is not blood but ISF ( Figure 4A ) present in the living epidermis. For the OECT-CGM system, the microneedle length is set to 1.0 mm ( Figure 4A ). To better penetrate the skin, a 10×10 array of sharp conical microneedles with a cone angle of 15 degrees is used. The efficiency of this structure in penetrating the skin has been reported previously. The size of the microneedle array is designed to be 10 mm×10 mm×2.2 mm to be compatible with the CGM readout system ( Figure 8 ). In addition, the surface properties of the resin-based 3D printed microneedle array can be modified by metallizing the needle tips ( Figure 9 ):

[0056] The hydrogel filler promotes the diffusion of glucose from ISF to the OECT sensor: Due to the concentration gradient, glucose molecules in ISF can passively diffuse through the hollow microneedles to the sensor. Since glucose is continuously consumed on the gate electrode of the OECT glucose sensor, there is a consistent concentration gradient between the IFS and the sensor in the vertical direction, which drives glucose to diffuse through the microneedle pores to the electrode ( Figure 4A)。However, due to wound healing at the tip of the microneedles, clogging often occurs. Therefore, a hydrogel is selected as the filler in the hollow microneedles to prevent wound healing while maintaining efficient diffusion of the target molecule to the sensor (Figs. 11 and 12). To evaluate the diffusion efficiency of the hydrogel, the glucose diffusion coefficient is measured by EIS analysis based on the Warburg element for quality control (Figs. 11 and paragraph

[0042] below). The Warburg impedance depends on several factors, including the concentration of the active material, the diffusion coefficient, and the number of mobile electrons. The Warburg impedance, also known as the Warburg coefficient (σ), can be expressed as follows, yielding the diffusion coefficient term: where R is the gas constant (8.314 J·mol-1·K -1 ), T is the temperature (in K), F is the Faraday constant (96485 C·mol -1 ), A is the electrode surface area (cm 2 ), C O is the concentration of the oxidized species (in mol), C R is the concentration of the reduced oxidized species (in mol), D O is the diffusion coefficient of the oxidized species (cm·s -1 ), D R is the diffusion coefficient of the oxidized species (cm·s -1 )

[0057] To evaluate the diffusion efficiency of the proposed double-network hydrogel as a passive diffusion layer, electrochemical impedance spectroscopy (EIS) is used to measure the impedance of the proposed system as a function of frequency. The Warburg impedance (characteristic of the low-frequency EIS spectrum) can be used to estimate the glucose diffusion coefficient from the hydrogel layer to the vicinity of the electrode. The diffusion of glucose molecules near the electrode can be described by Fick's second law of diffusion: where C is the glucose concentration, t is the time, x is the distance, and D is the diffusion coefficient. To estimate D, the Warburg impedance needs to be observed in EIS. The Warburg impedance is typically observed in the low-frequency range. The relationship between the diffusion coefficient of glucose and the slope of the Warburg impedance line can be expressed as: where D is the diffusion coefficient, R is the ideal gas constant, T is the absolute temperature, A is the electrode area, n is the number of electrons transferred per mole involved in the electrode reaction, F is the Faraday constant, C is the glucose concentration, and σ is the Warburg coefficient, i.e., the slope of the Warburg resistance line in the Z'-ω -1 / 2 plane.

[0058] By controlling the composition and thickness of the hydrogel film, due to the high porosity and hydrophilicity of the hydrogel, the glucose coefficient (3.90×10 -12 cm·s -1 ) is close to 6.54×10 -12 cm·s -1 of water. Although a similar sensitivity can be obtained using hydrogel fillers, the response delay is also 120 seconds compared to the reference sensor in direct contact with the ISF. This delay is observed regardless of the glucose concentration (1 mM to 10 mM), which is consistent with microneedle-based sensors and is attributed to the glucose diffusion process from the ISF to the OECT sensor ( Figure 4F ).

[0059] Tough, sticky hydrogels can improve the adhesion between the skin and the OECT-CGM: A strong adhesion between the sensor and the skin is a key factor in improving the durability and wearability of the CGM. To achieve this goal, the introduced hydrogel was further designed to be tough and adhesive by mixing an IPN hydrogel with a semipermeable bioadhesive elastomer. The IPN hydrogel consists of a tough sodium alginate first network and a soft PAAm second network and is functionalized with a GOx mixture ( Figure 4B ). The toughness and adhesiveness of the IPN hydrogel elastomer mixture ensure that the system remains on the skin even under strain and distortion conditions. In addition, mixing GOx in the IPN hydrogel provides a simple method for glucose detection without complex enzymatic modification procedures on the electrodes. Therefore, the IPN hydrogel-based sampling and sensing system is expected to be a promising solution for improving the reliability and wearability of wearable CGM systems.

[0060] To verify the efficacy of the fully integrated OECT-CGM system, it was compared with a reference CGM system (Dexcom G6). The reference device consists of four parts: a needle-based sensor with a length of 2 cm, a sticky patch, a support frame, and a wireless data transmitter. The OECT-CGM system proposed here consists of five parts: a microneedle array with a needle length of 1.0 mm, an IPN-hydrogel-elastomer mixture for glucose detection and adhesion, an OECT transducer, a PERfECT readout system, and an fPCB-based connector with a magnetic aligner and metal contacts for easy integration with a smartwatch.

[0061] In vitro experiments were conducted to benchmark the results of the OECT-CGM and the reference device ( Figure 13)。An artificial skin made of an elastomeric film was used to verify the microneedle penetration effect. During the experiment, the glucose level in the cavity was dynamically controlled by a pump, and the glucose values were monitored using the OECT-CGM system and a reference device. During the 180-minute test, the concentration cycled between clinically relevant values (3 mM to 20 mM). The results shown by the OECT-CGM system were comparable to those of the reference device( Figure 4H ), but it had a smaller size (1.5 cm × 1.5 cm) and weight( Figure 4D ). Before reaching the equilibrium state, a response delay of approximately 8 minutes was observed in the OECT-CGM system, which was attributed to the diffusion delay of glucose molecules in the IPN hydrogel. The OECT-CGM exhibited an acceptable mean absolute relative difference (MARD) value of 15%, indicating its potential for future practical applications( Figure 4F ).

[0062] The current regeneration ability of the OECT-CGM system was further verified. As shown in Fig. 14, after two weeks of continuous monitoring, the current change of the device gradually decreased from 355 μA to 322 μA. By reducing the gate voltage from 0.3 V to 0.2 V, at a glucose concentration of 6 mM, the current recovered from 322 μA to 345 μA. In addition, after the self-calibration process using the transconductance value measured in situ, the same sensitivity could be regained, but with a higher current and without being affected by environmental noise.

[0063] 1.5 In vivo continuous blood glucose measurement using the OECT-CGM system

[0064] To verify the continuous blood glucose monitoring (CGM) ability of the proposed OECT-CGM, in vivo experiments involving rats were conducted, as Figure 4G shown. To facilitate effective monitoring, the dorsal region was selected as the application site because it allowed the microneedles to be properly inserted into the dorsal region of the rats, thus ensuring the stable operation of the OECT-CGM in this region, as Figure 15 shown. To induce a hyperglycemic state in the rat model, a glucose solution was administered to control the blood glucose level of the rats. Once the blood glucose level of the rats reached the peak, the animals relied on their natural glucose regulation mechanism to lower their blood glucose level to the normal state. Approximately one hour after the administration of the glucose solution, the blood glucose level of the rats reached the peak, approximately 15.4 mM / L.

[0065] The OECT-CGM continuously measures blood glucose levels at 90-second intervals. Additionally, blood samples were collected from the rat tail vein every 15 minutes, and their blood glucose levels were measured using a commercial glucometer for comparison. This data collection process continued for two hours. The results showed a strong correlation between the OECT-CGM glucose sensor and the commercial glucometer. Notably, the peaks of these two measurements occurred in a relatively synchronous manner (t = 110 minutes). This observation highlighted the feasibility and effectiveness of the in vivo application of this OECT-CGM. It is worth mentioning that a slight delay of about 15 minutes was observed, which can be attributed to the diffusion of glucose from the bloodstream into the interstitial fluid and then from the interstitial fluid through the hydrogel layer to the sensor, as Figure 4H shown.

[0066] Example 2 - Discussion

[0067] In conclusion, the OECT-CGM system described in this paper represents a significant advancement in the field of wearable CGM. By integrating OECT as a glucose biosensor and a microneedle array as a minimally invasive sampling technique, this system provides three key functions that are expected to improve glucose monitoring. First, it offers an adjustable linear region, providing enhanced sensitivity and customizable measurements. This function has the potential to enable accurate monitoring in real-world scenarios. Second, the OECT-CGM system demonstrates the ability of current regeneration achieved by adjusting the gate voltage. This ability allows the device to restore its performance by releasing captured ions or molecules, addressing issues related to performance degradation over time and in challenging environments. Finally, the system incorporates an auto-calibration ability, enabling accurate glucose measurement without external calibration.

[0068] In addition, the OECT-CGM system exhibits stability through various mechanisms. The ability of current regeneration, combined with the stability of GO x mixed in the tough hydrogel and its role as a mechanical adhesion buffer layer, contributes to the stability of the entire system. Additionally, the microneedle array acts as a bridge between the internal environment and the externally deployed biosensor, minimizing the foreign body response and potential biological contamination commonly seen in CGM.

[0069] Example 3 - Materials and Methods

[0070] 3.1 Materials

[0071] The PEDOT:PSS aqueous suspension (Clevios PH1000) was purchased from Heraeus Electronic Material (USA). Glycerol, dodecylbenzenesulfonic acid (DBSA), sodium chloride, (3-glycidoxypropyl)trimethoxysilane (GOPS), acrylamide (AAm), calcium chloride (CaCl2), and 3-(trimethoxysilyl)propyl methacrylate (TMSPMA) were purchased from Sigma-Aldrich (USA). Hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959), glutaraldehyde, ferrocene, GOx, and chitosan were provided by Aladdin Co. (Shanghai, China). The PI film was provided by DuPont Company (USA). The biocompatible adhesive polyurethane was provided by 3M Company (USA). Unless otherwise stated, the chemicals used in this work were not further purified.

[0072] 3.2 Animals

[0073] Animal experiments were approved by the Animal Ethics Committee of Guangzhou Medical University and were conducted in accordance with relevant guidelines. Female Sprague-Dawley rats (300 to 350 g) were purchased from the Guangdong Medical Experimental Animal Center and acclimated in an approved animal facility.

[0074] 3.3 Microneedle Fabrication

[0075] To fabricate the hollow microneedle patch, a high-resolution 3D printer (S240 from Boston MicroFabrication) was used. This printer employs ultra-precision stereolithography (uPSL) technology, which can achieve a printing accuracy of up to 10 μm in both the transverse and longitudinal directions. The microneedle structure was designed using computer-aided design (CAD) software and then printed using a photopolymer resin. The printed structure was then post-processed by curing them in an ultraviolet chamber and rinsing them with isopropyl alcohol to remove any uncured resin. The mechanical strength and skin penetration ability of the microneedle patch were tested using an in vitro skin model.

[0076] 3.4 OECT Sensor Fabrication

[0077] To fabricate an OECT on a flexible PI substrate, the process starts with constructing source / drain / gate electrodes by depositing a thin film of gold on it, the pattern of which is defined by a metal shadow mask. Then, the active channel layer PEDOT:PSS is deposited between the source and drain using inkjet printing, with a minimum droplet size of 1 pL. The printable PEDOT:PSS ink is prepared by first stirring for 3 minutes and then mixing it with GOPS (1 w / w.%), glycerol (5 v / v.%), and DBSA (0.1 v / v.%) using a Vortex (MX-S). Glycerol is added to increase the film conductivity. DBSA is added to promote the wettability of the film on the substrate. Before printing, the mixed suspension is filtered through a PTFE membrane (pore size: 0.45 μm) to remove aggregates and avoid nozzle clogging. Next, an ultraviolet-curable plastic insulating layer is also deposited using an inkjet printing system (Yiwu Yangtian Electronic Technology Co., Ltd., Yiwu, China), and in-situ curing is carried out using an ultraviolet lamp (wavelength: 345 nm) attached to the printer. To complete the OECT structure, a tough IPN hydrogel composed of PAAm / sodium alginate double network is prepared to bridge the gate electrode and the channel.

[0078] To synthesize the tough IPN hydrogel, 2% (w / w) sodium alginate, 13% (w / w) AAm, 0.015% (w / w) MBAA, and 0.24% (w / w) Irgacure 2959 are dissolved in deionized water. Then the mixture is centrifuged at 9000 rpm to remove air bubbles. The formation of the PAAm network can be achieved by curing the solution in a UV chamber (364 nm, 10 W power) for 60 minutes. The second network is formed by crosslinking sodium alginate by soaking the previous gel in 1 M CaCl2 solution for 24 hours to reach an equilibrium state.

[0079] To make the system specific for glucose sensing, an enzymatic hydrogel is prepared by embedding GOx and ferrocene into the tough IPN hydrogel, where GOx can selectively react with glucose and ferrocene serves as a redox mediator. For this purpose, the IPN hydrogel is first freeze-dried for 48 hours to obtain a double-network scaffold. After the scaffold is swollen overnight in an aqueous solution containing GOx (500 units / mL), ferrocene (10 mM), and glutaraldehyde (0.3% (w / w)), the enzymatic tough hydrogel can be obtained.

[0080] To make the IPN hydrogel adhesive, the hydrogel was mixed with a bioadhesive elastomer film (provided by 3M). The elastomer film was punched with a puncher (200 microns in diameter) to facilitate the ISF exchange between the body, the hydrogel, and the OECT sensor. The adhesive elastomer was first treated with a TMSPMA solution (100 mL deionized water, 10 μL acetic acid at pH 3.5, and 2 wt.% TMSPMA), then washed with ethanol and dried. The gel precursor was carefully dropped onto the TMSPMA-grafted elastomer to avoid air bubbles and cured under UV irradiation for 60 minutes. The formation of the IPN hydrogel followed the same procedure as described above. The grafted TMSPMA tended to interconnect the IPN hydrogel and the elastomer, forming a mixture that could be firmly adhered to human skin.

[0081] 3.6 Miniaturized Readout System Fabrication and Evaluation

[0082] To reduce the size of the readout system, the recently developed miniaturized OECT readout system PERfECT was modified to meet the specific requirements of CGM. The PERfECT system adopted firmware customization and was designed to include a customized MCU, a low-power Bluetooth (BLE) unit, and other necessary components, all of which were soldered onto a printed circuit board (PCB) using standard manufacturing processes. Then, the readout system was connected to the OECT glucose sensor using an fPCB connector. The typical power consumption of the PERfECT system was measured to be 1 mW when sampling at a rate of 1 sample per second (SaPS) using a Keithley B2902B. The performance of the improved PERfECT system was evaluated through in vitro experiments using an analog glucose solution, and the accuracy and precision of the system were analyzed using statistical methods.

[0083] 3.7 Integration Strategy

[0084] To achieve the integration of individual components, we adopted a meticulous approach. The microneedle patch produced by high-resolution 3D printing was seamlessly connected to the OECT sensor. This was achieved using an fPCB connector, which ensured a reliable electrical connection between the sensor and the miniaturized readout system. The readout system itself was customized to include an MCU, a BLE unit, and other key components. This customized design facilitated a smooth and optimized interface with the OECT sensor. By adopting efficient connection technologies and carefully assembling these components, a compact and fully integrated CGM system was successfully realized, enabling continuous blood glucose monitoring with maximum user-friendliness and operational efficiency.

[0085] 3.8 Anti-Interference Test

[0086] To rigorously evaluate the selectivity of the OECT-based glucose sensor, a comprehensive anti-interference test was conducted. The OECT-based glucose sensor is an integral part of the compact OECT-CGM system. This evaluation aimed to determine the sensor's ability to discriminate and accurately measure glucose levels in the presence of various interfering substances that may coexist in the human body.

[0087] To mimic real-world conditions and cover a wide range of potential interferents, four representative substances were selected, namely uric acid, ascorbic acid, paracetamol, and lactic acid. These substances were introduced into the test container at a concentration of 1 M to simulate concentrations relevant to the in vivo scenario.

[0088] To ensure the precision and control throughout the experiment, an advanced programmable injection pump (XMSP-2CI, China Ximai Nanotechnology Co., Ltd.) was employed. This enabled us to systematically adjust the concentration of the interfering substances in the test container, controlling the movement of the syringe through a stepper motor with a precision as low as 10 μm. By incorporating this precision in this method, the aim was to simulate the dynamic changes and fluctuations in the levels of these interfering substances commonly found in the human body.

[0089] 3.9 In vitro experiments

[0090] To evaluate the actual performance of the OECT-CGM system, in vitro experiments were conducted to simulate dynamic glucose changes in the human body. The test setup included a glucose solution reservoir with an inlet and an outlet, a programmable syringe that allowed precise control of the injection rate, and a programmable injection pump ( Figure 13 ) that enabled us to adjust the glucose concentration in the reservoir. In these experiments, the concentration changes were carried out in two cycles in the range of 0 to 20 mM and vice versa, with the rate increasing at 0.3 mM per minute. Each concentration change cycle took approximately 2 hours to simulate the potential fluctuations in blood glucose levels in vivo.

[0091] 3.10 In vivo trials

[0092] Female Sprague-Dawley rats (300 to 350 g) were used for in vivo OECT-CGM glucose tracking trials. The rats were fasted for 20 hours but provided with water to obtain a stable initial blood glucose level. First, to facilitate blood sampling, the rats were anesthetized with 2.5% (v / v) isoflurane. To control the blood glucose level of the rats, 400 mM glucose solution was injected into the rats at a dose of 10 mL / kg through a syringe 40 minutes after the start of the experiment. The OECT-CGM was applied to the back of the rats, with the sampling rate set at 90 seconds / sample and monitored for 2 hours. To verify the blood glucose level of the rats, blood was continuously drawn from the rat tail vein at a frequency of 15 minutes for 2 hours using a blood glucose meter (Yuwell Accusure 590, China) for testing.

[0093] All patents, patent applications, provisional applications, and publications mentioned or cited herein are hereby incorporated by reference in their entirety, including all figures, to the extent not inconsistent with the explicit teachings of this specification.

[0094] It should be understood that numerous specific details, relationships, and methods are set forth herein to provide a thorough understanding of the present invention. However, one of ordinary skill in the relevant art will readily recognize that the present invention may be practiced without one or more of the specific details, or with other methods, protocols, reagents, cell lines, and animals. The present invention is not limited by the order of acts or events shown, as some acts may occur in a different order and / or concurrently with other acts or events. In addition, not all acts, steps, or events shown are necessary to practice the methods according to the present invention. Many of the techniques and procedures described or cited herein are well known to those skilled in the art and are generally employed using conventional methods.

[0095] Unless otherwise defined, all professional terms, symbols, and other scientific terms or terminology used herein are intended to have the meaning commonly understood by one of ordinary skill in the art to which this invention pertains. In some cases, terms with commonly understood meanings are defined herein for clarity and / or for ease of reference, and the inclusion of such definitions herein should not necessarily be construed as representing a substantial difference from the meaning commonly understood in the art. It should also be understood that terms (such as those defined in common dictionaries) should be interpreted as having a meaning consistent with the meaning in the relevant technical context and / or as otherwise defined herein.

[0096] As used herein, the terms “about,” “substantially,” “essentially,” and “approximately” are used to describe and account for small variations. When used in conjunction with an event or circumstance, the term can refer to the instance in which the event or circumstance occurs precisely as well as instances in which the event or circumstance occurs approximately. As used herein, with respect to a given value or range, the term “about” generally means within ±10%, ±5%, ±1%, or ±0.5% of the given value or range. This range may be indicated herein as from one endpoint to the other endpoint or between two endpoints. Unless otherwise stated, all ranges disclosed in this disclosure include the endpoints. When referring to “substantially” the same numerical value or property, the term can refer to a value within ±10%, ±5%, ±1%, or ±0.5% of the average value.

Claims

1. A minimally invasive and blood-contact-free continuous blood glucose monitoring system based on organic electrochemical transistors, characterized in that Comprising: A hollow microneedle patch with a surface-deposited metal; A viscous and stretchable enzyme hydrogel sensing film; A glucose sensor based on an organic electrochemical transistor; A wearable micro readout system; and A 3D printed plastic encapsulation housing; wherein the organic electrochemical transistor has functions of current regeneration, sensitivity adjustment, and automatic calibration.

2. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, characterized in that The hydrogel sensing film comprises: An interpenetrating polymer network hydrogel including a first network of sodium alginate and a second network of polyacrylamide; A semipermeable bioadhesive elastomer; and Glucose oxidase.

3. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, wherein The surface-deposited metal is selected from gold or platinum.

4. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, characterized in that The system is capable of tracking 10 -6 M to 10 -1 Cellular glucose metabolism and clinical blood glucose concentrations within the concentration range of M.

5. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, wherein The size of the system is less than 2 cm × 2 cm × 0.5 cm.

6. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, wherein The current regeneration function is achieved by adjusting the gate voltage of the organic electrochemical transistor.

7. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, characterized in that Its signal-to-noise ratio is at least 50 dB.

8. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, wherein The sensitivity adjustment function includes controlling the transconductance to adjust the noise resistance; and adjusting the gate voltage to adjust the linear range of detection.

9. The continuous blood glucose monitoring system based on an organic electrochemical transistor according to claim 1, characterized in that, The automatic calibration function includes standardizing the transconductance curve.

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