A Flexible and Wearable Chemiresistive Biosensor Fabricated by Laser Inducing for Real‐Time Glucose Analysis of Sweat

In this study, a flexible and wearable chemiresistive biosensor (FWCB) is developed for the real‐time analysis of glucose in sweat on the human skin surface based on a novel detection strategy of p‐type reduced graphene oxide (rGO) sensing film, which met the requirements of rapid, nondestructive testing. The proposed FWCB is fabricated in the form of interdigital electrodes (IEs) made of laser‐induced graphene (LIG) synthesized by the laser inducing of a polyimide (PI) film. Additionally, a semiconducting rGO sensing film modified on the surface of IEs is synthesized by thermal reduction of graphene oxide (GO), which is functionalized with glucose oxidase (GOx) by chemical cross‐linking to obtain GOx/FWCB. Moreover, the key parameters for FWCB fabrication are optimized, and the sensing strategy of the proposed GOx/FWCB is also investigated. The results show that the proposed GOx/FWCB can be used for the detection of glucose in the range of 0.01–3.0 mM with satisfactory selectivity, and the limit of detection (LOD) is calculated to be as low as 0.8 µM (S/N = 3). These dramatic advantages endow the proposed FWCB with broad application prospects in the field of portable, wearable, and real‐time detection of glucose in human sweat for health monitoring.


Introduction
Diabetes mellitus (DM) is a serious incurable disease that is extremely harmful to human life. By 2021, the number of DM patients in the world has reached 537 million, and the annual rate of increase continues to rise. [1][2][3] DM poses a serious threat to human physiological health and social development because it can cause illness with high mortality, such as kidney disease, cardiovascular disease, and neurological disease. [4] Therefore, it is necessary to monitor the blood glucose levels of humans, especially diabetic patients. Traditional glucose detection methods include continuous glucose monitoring, venous plasma glucose monitoring, and glycosylated serum protein monitoring. Although these methods have accurate detection results, the majority of blood glucose monitoring is performed by invasive diagnostic tools, which will have a certain impact on the patient's physical and mental health, while the glucose content in sweat is closely related to the blood glucose level, so the detection of glucose concentration in sweat can be expected to be used as a new way to indirectly reflect the concentration of glucose in blood for monitoring the health status of diabetic patients.
In recent years, with the development of material science and the popularization of information technology, sensor technology has played a vital role in improving people's living standards. Typically, traditional sensors fabricated with rigid substrates are hard, inflexible, and hard-to-wear devices. Consequently, flexible wearable biosensors (WBs), celebrated for their lightness, flexibility, and portability, have captured the interest of global researchers. Their practical applications, particularly in the realtime and continuous monitoring and diagnosis of physiological states, are increasingly recognized and valued. [5,6] Presently, the primary application of WBs lies in real-time analysis of human sweat constituents, specifically concentrating on sweat glucose detection. This is particularly beneficial for diabetes patients, as it significantly enhances the convenience and effectiveness of their disease management. [7][8][9] Among the existing biosensors for glucose detection, electrochemical sensors are the most widely used. [10][11][12][13] However, the integration of traditional electrochemical detection devices into the fabrication of flexible sensors is challenging, and their detection accuracy falls short of the rigorous standards required for sweat glucose monitoring. [14] In contrast, field-effect transistor (FET)-based chemiresistive biosensors have the characteristics of low cost, easy preparation, and high sensitivity and selectivity. These benefits stem from their working principle, wherein the biomolecules captured by the biorecognition element modified on the semiconducting sensing interface would change the charge density of the sensing interface and affect the arrangement of holes and electrons in the semiconducting sensing interface to regulate the resistance of the semiconducting sensing interface between the source electrode and drain electrode and finally achieve target analyte detection based on the relationship model between the concentration of biomolecules and the resistance change rate. [15][16][17] Compared with non-enzymatic glucose biosensors, enzymatic glucose biosensors have attracted great interest due to the high selectivity and rapid response. [18,19] However, enzyme biosensors still have some problems, such as a lack of stability, poor immobilization effect, and inactivation in long-term detection processes. [20] Therefore, to improve the immobilization of glucose oxidase (GOx), a chemical covalent-coupling method was applied to biosensor fabrication, which enabled GOx to be closely immobilized on the functionalized FET-based semiconducting sensing interface by a coupling reaction. [21] Graphene, a 2D carbon material, is packed tightly with sp 2 hybrid carbon atoms, forming a 2D honeycomb structure and having many dramatic properties, such as excellent electrical conductivity, high thermal conductivity, stable chemical properties, large specific surface area and great flexibility, [22,23] resulting in the application of graphene in FET-based chemiresistive biosensors, and attracting extensive attention. [24] During the fabrication of FET-based chemiresistive biosensors, the semiconducting sensing interface usually needs to be functionalized for the immobilization of various biorecognition elements. [25] Consequently, abundant carboxyl groups can be uniformly modified on the semiconducting sensing interface based on the -stacking interaction between the six-membered ring of graphene and the substance, [26][27][28] which can be expected to supply a large number of evenly distributed functional groups for the immobilization of enzymes.
To date, traditional electrodes, such as glassy carbon electrodes, [29][30][31] carbon paste electrodes, [32][33][34] and screenprinted electrodes, are made up of rigid and solid substrates, which makes them difficult to use for the wearable sensing of glucose on human skin. To resolve the above problems, a new method for the batch fabrication of flexible electrodes, which is easy to use, low-cost, and efficient, has been developed by the laser inducing of a flexible polyimide (PI) film to generate porous graphene, i.e., laser-induced graphene (LIG), with specific patterns under controllable laser power. [35,36] Interdigital electrodes made in 3D porous LIG (LIG-IEs) exhibit not only excellent conductivity, but impressive flexibility. This makes them a superior alternative to traditional rigid and solid electrodes, making them ideal for applications in flexible and wearable sensors. [37,38] In this paper, a flexible and wearable FET-based chemiresistive glucose biosensor (FWFCGB), i.e., rGO-GOx/LIG-IEs, was fabricated on a PI film by a facile one-step laser inducing technique, thereby providing a new strategy for the large-scale production of flexible and wearable glucose biosensors. The sensing area of the fabricated LIG-IEs was modified with a rGO film based on the thermal reduction of GO. Meanwhile, 1-Pyrenecarboxylic acid (PCA) with a -bond structure was used to functionalize rGO and provide abundant carboxyl groups for the subsequent immobilization of GOx on the carboxyl-functionalized rGO sensing interface based on chemical cross-linking, in which rGO was used as a p-type semiconducting layer. Additionally, the obtained interdigital LIG electrodes and sensing film were characterized by scanning electron microscopy (SEM), Raman spectroscopy and Fourier transform infrared spectroscopy (FTIR). Moreover, the proposed rGO-GOx/LIG-IEs FWFCGB demonstrated a linear detection range of 0.01 to 3.0 mM with a limit of detection (LOD) of 0.8 μM (S/N = 3). Furthermore, the feasibility and reliability of the fabricated rGO-GOx/LIG-IEs FWFCGB was verified by practical application to human sweat analysis, which exhibited excellent performance in the detection of glucose in vitro sweat.

Reagents and Instruments
The polyimide (PI) film was purchased from Tianjin Jiayin Nanotechnology Co., Ltd. The LIG interdigital electrodes were fabricated on PI film by a computer-controlled laser scribing micromachine (Nano Pro-III, Tianjin Jiayin Nanotechnology Co., Ltd., China). All electrical measurements were performed on a Keithley 2614B System Source Meter (Tektronix, Beaverton, OR, USA). The morphology of LIG, GO and rGO was characterized by a metallographic microscope (13XB-PC, Shanghai, China) and scanning electron microscopy (SEM, Zeiss Gemini 300). Fourier transform infrared spectroscopy (FTIR) was carried out by an FTIR spectrophotometer (Thermo Fisher Scientific, USA). Raman spectroscopy were obtained using a Horiba Scientific LabRAM monochromator equipped with a 532 nm (2.33 eV) laser. www.advancedsciencenews.com www.advmatinterfaces.de

Fabrication of the Flexible and Wearable rGO-GOx/LIG-IEs Biosensor
The flexible and wearable rGO-GOx/LIG-IEs biosensor was fabricated as follows (Figure 1): the PI film was first placed on glass and cleaned with ethanol. Then, LIG-IEs were patterned on a PI sheet by the laser inducing based on a computer-controlled laser system ( Figure 1a). Next, the conductive silver ink was brushed on both sides of the exposed IEs and heated at 80°C for 15 min to prepare the connecting wires of the IEs (Figure 1b). After that, 7 μL of 4.0 mg mL −1 GO solution was dropped on the sensing area of the IEs and dried at room temperature until a solid film was formed. Subsequently, the obtained LIG-IEs modified with GO (GO/LIG-IEs) were annealed at 200°C for 120 min in a thermostatic drying oven ( Figure 1c). Next, the PI solution was mixed with NMP in a volume ratio of 4:1, and then the obtained PI solution was brushed around the rGO sensing area as a solid hydrophobic insulating layer ( Figure 1d). The obtained rGO/LIG-IEs were further incubated with 100 mM PCA dissolved in DMF for 60 min (Figure 1e). The obtained rGO-PCA/LIG-IEs were washed with DMF and then blow-dried with N 2 gas. After that, the electrode was sequentially incubated with 10.0 mM EDC and 10.0 mM NHS in 100 mM MES buffer (pH 5.5) for 20 min to activate the carboxyl groups ( Figure 1f). Subsequently, the obtained rGO-PCA-EDC-NHS/LIG-IEs were incubated in 10 mL of GOx solution at 4°C for 3 h ( Figure 1g); thereafter, the electrode was incubated with 100 mM EA for 30 min to passivate residual ester groups of NHS and then rinsed thoroughly with 10 mM PBS (pH 7.2). Then, 1.0% (v/v) PBST was dropped on the sensing area and kept for 30 min to block the nonfunctionalized rGO modified on the sensing area (Figure 1h), and finally, a flexible rGO-GOx/LIG-IEs biosensor was obtained. The chips were washed with PBS buffer at a proper pH after each of the previous modification steps. The resistance of the electrode obtained above each step was measured by a source meter. To avoid direct contact between the skin and the sensing interface, two different layers of PDMS with specific patterns were designed as a double-channel sweat drainage layer and a circular sweat storage layer (Figure 1i). Finally, the flexible and wearable rGO-GOx/LIG-IEs biosensor was attached to the skin surface using noninvasive skin tape to facilitate real-time monitoring of sweat glucose levels.

Glucose Detection Using the Proposed FET-Based Chemiresistive Biosensor
The procedures of glucose detection using the proposed rGO-GOx/LIG-IEs biosensor could be briefly described as follows: 10 μL of PBS solution (pH 7.2) was dropped on the sensing interface, and the initial resistance of the obtained sensing interface was measured as R 0 . Subsequently, 10 μL of PBS solution (pH 7.2) with varying concentrations of GLC (0.01, 0.05, 0.1, 0.5, 1, 2, and 3 mM) were carefully added to the above droplet on the sensing interface, and the resistance after each addition was recorded as R n (n = 1, 2, 3, 4, 5, 6, and 7). The whole measurement process was performed using the R-t module of the digital source meter, wherein the source-drain voltage was set to 0.1 V. The resistance change rate (RCR) of each different concentrations of glucose was calculated by Equation 1.
A linear fitting equation was established using glucose concentration as a function of RCR over a concentration range of 0.01-3.0 mM, which was applied to the further prediction of glucose concentration in human sweat. Additionally, the experiments involving human subjects were performed with the full and informed consent of the volunteer.

Characterization of the Modified LIG-IEs
Considering that the electrochemical property and morphology of the rGO and porous LIG were important factors that affected the detection performance of LIG-IEs, the parameters affecting the rGO/LIG-IEs detection performance, i.e., laser power, laser scanning rate, and GO concentration, used for the fabrication of rGO/LIG-IEs were optimized in this study by the analysis of resistance characteristics, and the electrochemical property and morphology of the obtained LIG-IEs and rGO/LIG-IEs were investigated as well. As shown in Figure S1 (Supporting Information), the higher the laser power is, the better the electrical conductivity of the IEs. [39] Moreover, Figure S2a-d (Supporting Infor-mation) shows that with increasing laser power, the color of the laser-generated IEs gradually darkened under the optical microscope as the carbonization degree of the PI film was promoted. [40] When the laser power exceeds 2.2 W, the carbonization effect of PI is aggravated, resulting in incomplete IEs generated by laser printing. Consequently, a laser power of 2.2 W was selected for the following sensor fabrication. Besides, the morphologies of LIG-IEs prepared by different laser scan rates with a laser power of 2.2 W were also characterized by optical microscopy, as shown in Figure S2f-i (Supporting Information). Finally, the optimal laser scanning rate was verified to be 4 cm s −1 .
Additionally, Raman spectroscopy and FTIR were applied to investigate the essential attributes and surface functional groups of the obtained LIG, respectively. As shown in Figure 2a, the porous graphene obtained by the laser inducing of PI had three main peaks in the Raman spectroscopy, including the D band (≈1355 cm −1 ), G band (≈1585 cm −1 ), and 2D band (≈2705 cm −1 ), wherein the 2D band and G band exactly corresponded to the typical peaks for graphene materials, which demonstrated that a large number of carbon atoms were transformed from the SP3 lattice to the SP2 lattice, while some incomplete transformation of carbon atoms would induce defects in the generated LIG, which was proven by an ID/IG ratio of 1.13. As shown in Figure 2b, the FTIR spectra showed that the LIG synthesized with optimal laser power and scanning rate lacked oxygen-containing functional groups, which enabled the good conductivity of LIG due to the fewer defects. The morphology of LIG was investigated by SEM at different magnifications, as shown in Figure 2c,d. The LIG films shown in Figure 2c,d exhibit a foam shape with a porous structure due to the rapid liberation of gases produced during the formation of LIG by high-energy laser inducing of the PI film. The 3D porous structure of LIG provided abundant and various electron transport channels for improving the electrical conductivity of the fabricated LIG-IEs. [41]

Optimization and Characterization of the rGO Sensing Layer
To further improve the responses of the developed rGO-GOx/LIG-IEs biosensor, the concentration of GO was optimized depending on the resistance analysis of the IEs modified with different concentrations of GO, as shown in Figure 3. The resistance of rGO/LIG-IEs gradually decreased from 2.4 to 0.4 kΩ as the concentration of GO increased from 0.5 to 4.0 mg mL −1 , while the resistance leveled off in the concentration range of 4.0-5.0 mg mL −1 , as shown in Figure 3a. This was because the rGO film was formed more completely as the GO concentration increased, and thus, the device resistance decreased. Additionally, a large standard deviation of 2 kΩ was observed at a concentration of 0.5 mg mL −1 because GO with a low concentration made it difficult to construct enough crosslinked graphene layers and form a stable rGO film after thermal reduction, resulting in poor repeatability of the electrical conductivity. By comparing the resistance of IEs modified with different concentrations of GO after thermal reduction, a GO concentration of 4 mg mL −1 was preferred. As shown in Figure 3b, an ultrathin rGO film was modified on the surface of LIG-IEs, in which the width of each electrode and the distance between each electrode were ≈449.49 and 170.38 μm, respectively, as a semiconducting sensing layer based on the thermal reduction of GO at 200°C for 1 h. From Figure 3c, after reduction, GO changes from brownish yellow to silver gray. This is because GO had unsaturated double-bond groups and hydroxyl chromophore groups formed in the oxidation process, while as the functional group was reduced, the color changed from brownish yellow to black. [42,43] By comparing the optical images of GO/LIG-IEs formed by drop coating GO solution at different concentrations, i.e., 0.5, 1, 2, and 4 mg mL −1 , before thermal reduction, the highest light transmission of GO/LIG-IEs was found to be produced at a concentration of 0.5 mg mL −1 . As the concentration of the drop coating increases, the light transmission decreases, and the dark brown color of the GO/LIG-IEs can be clearly observed at a concentration of 4 mg mL −1 . This result demonstrates that the GO layers generated at this concentration are denser. In addition, optical images of rGO/LIG-IEs formed with different concentrations of GO solutions after thermal reduction reveal that the rGO layer produced by the reduction of 0.5 mg mL −1 GO solution has more defects. As the concentration of the GO solution increases, the reduced graphene oxide layer becomes darker, more complete and more uniform. Additionally, SEM was used to characterize the morphologies of GOand rGO-modified LIG-IEs, respectively. As can be seen from Figure 3d,e, rGO generally displays the fewer stacking layers and more densely structure compared to GO, due to the removal of oxygen-containing functional groups during the reduction process. Simultaneously, the SEM images revealed a wrinkled texture that was associated with the presence of flexible and ultrathin graphene sheets, which provided functionalized sites for the immobilization of biometric recognition elements, i.e., GOx.
Moreover, the prepared GO and rGO nanolayers were characterized by Raman spectroscopy, as shown in Figure 4a. Two fundamental vibrations can be observed in the Raman spectroscopy over the range from 1100 to 1700 cm −1 for the GO and rGO nanolayers, wherein 1300 and 1500 cm −1 correspond to the disordered (D) band and tangential (G) band of GO and rGO nanosheets, respectively. As shown in Figure 4a, the D band intensity of rGO is smaller than that of GO, which indicated that most oxygen components in GO were removed by thermal reduction at a high temperature, resulting in a more isolated graphene domain in rGO. On the other hand, the I D /I G ratio for GO was calculated to be 0.71, while rGO had an even higher I D /I G ratio of 0.88, which indicated that graphene had fewer defects due to deoxidation during thermal reduction, making it easier for electrons to be transferred. [42] As shown in the FTIR spectra of GO (Figure 4b, black line), the stretching vibrational peaks of the C-O single bond stretching vibration, carbonyl groups, C=C double bonds, and hydroxyl groups appeared at 1140, 1660, 1710, and 3340 cm −1 , respectively, which were induced by the oxidation of graphite leading to the generation of a large amount of oxygen-containing functional groups. However, with thermal reduction at 200°C, synthetic rGO removed most of the oxygen-containing functional groups, which was clear from the significant decrease in their peak strengths. The thermal reduction process is actually the deoxidation of GO, where most of the oxygen-containing functional groups are removed in the form of CO 2 and water vapor, [48] which in turn helps to remove the foreign atoms (oxygen) and convert more sp 3 carbons into sp 2 ones. Therefore, the electrochemical performance of rGO obtained by thermal reduction was significantly improved.

Characterization and Optimization of the Proposed FWFCGB
The resistance characteristics of the sensing interface after each modification step were investigated to determine whether the molecular linkers were successfully modified on the sensing interface, as shown in Figure 5a. The resistance of the sensing interface increased after each modification step of Ag/rGO, PCA, EDC/NHS-GOx, EA, and Tween-20 compared to the bare LIG-IEs, which demonstrated that the molecular linkers and GOx were successfully modified on the sensing interface due to the blocking effect caused by the binding of large amounts of macromolecular compound to the binding sites on the surface of LIG-IEs. [44] Additionally, the incubation time of GOx was also optimized for the fabrication of the proposed FWFCGB by analyzing the resistance change rate of the sensing interface. As shown in Figure 5b, the resistance of the sensing interface increased with increasing incubation time over a range of 1-3 h because the PCA modified on the rGO semiconducting layer reacted with EDC/NHS to form an amide bond, i.e., -CO-NH-, for the immobilization of GOx. When the incubation time exceeds 3 h, the resistance change rate remains almost constant because the specific binding sites on the semiconducting sensing interface for GOx immobilization tend to be saturated, resulting in the amount of modified GOx reaching the maximum. Based on the above results, an incubation time of 3 h was selected for this study.

The Sensing Performance of the Proposed FWFCGB
Under the optimized conditions, the sensing performance of the proposed FWFCGB was investigated by analysing the relationship between the resistance change rate and the corresponding glucose concentration. As shown in Figure 6a, real-time resistance measurements were performed on the proposed FWFCGB using the R-t module of the digital source meter, where the source-drain voltage was 0.1 V by dropping 10 μL of different concentrations glucose every 40 s, while by this time period, glucose was essentially completely decomposed onto the sensing area of the proposed FWFCGB with a droplet of 10 μL of PBS used as a baseline solution. To facilitate the test of resistance changes after successive drops of glucose solution, we adopted Origin software to obtain the corresponding baseline-corrected curves (Figure 6a), wherein the resistance increased with each drop of glucose solution. As shown in Figure 6b, a linear fitting equation was set up, wherein the resistance change ratio of the sensing interface has a linear relation with the glucose concentration over a concentration range from 0.01 to 3.0 mM. The obtained linear equation is y = 7.92x + 3.21 with an R 2 of 0.99, where y and x represent the resistance change rate and glucose concentration, respectively. The detection limit of the proposed FWFCGB was calculated to be 0.8 μM based on the formula LOD = 3 × SD/Slope (SD: blank standard deviation; Slope: slope of the calibration curve), which was lower than that of most other GOx sensors, as shown in Table 1. In addition, the FWFCGB developed in this study was wearable. The results show that the FWFCGB proposed in this study is capable of monitoring the blood glucose level of diabetic patients since the glucose concentration detection range contained the typical range of glucose concentrations of diabetic patients sweat exceeds 0.1 Mm. [54]

Repeatability and Specificity of the Developed FWFCGB
To further explore the performance of FWFCGB prepared by the above method for the detection of glucose, the stability and specificity were also verified, as shown in Figure 7. Under the optimal conditions, the stability verification experiment was carried out by five repeated measurements of 10 μL 0.4 mM glucose using the same rGO-GOx/LIG-IEs, wherein after each measurement, the sensing area of the sensor was washed with PBS standard solution (pH 7.2). Then, to maintain the same enzyme activity in each test, the sensor was kept at 4°C for drying before the next measurement. The results indicated that the rGO-GOx/LIG-IEs  biosensor had good repeatability with a relative standard deviation (RSD) of 4.7%, which was less than 5%. Meanwhile, in addition to the target analyte, i.e., glucose, the interference from other organic compounds present in sweat secreted by human skin, such as urea, ascorbic acid (ASA), acetaminophen (APAP), and uric acid (UA), on the responses of the proposed FWFCGB was investigated. As shown in Figure 7b, the measurements of 0.2 mM glucose were performed in the presence of 1 mM urea, 1 mM ASA, 1 mM APAP, and 1 mM UA. A maximum response could be observed for ASA, with a resistance change rate of 1.36%, which was 25.8% of glucose's response. Although ASA interferes with the response of FWFCGB, its concentration in sweat is as low as 10-50 μM, so the interference can be ignored. The results show that the proposed rGO-GOx/LIG-IEs biosensor has good anti-interference capability and can be expected to be applied to the analysis of glucose on the human skin surface.

Sensing Mechanism of the Proposed FWFCGB
The semiconducting property of the rGO synthesized by the thermal reduction of GO used for the fabrication of FWFCGBs was investigated by analyzing the charge transport characteristics of rGO using a liquid-ion gated FET system in 0.1 m KCl, wherein an Ag/AgCl electrode was used as a gate electrode. As shown in Figure 8a, the transfer curve, i.e., V g -I ds curve, of rGO was measured with a V ds of 0.1 V over a V g range from −1.0 to 1.0 V. The p-type semiconducting property of the rGO FET is evident from the decrease in I ds with increasing V g , which indicates that holes are the majority carriers in our synthesized rGO-based FETs. To understand the sensing mechanism of the proposed FWFCGB, the fundamentals of doping and charge transport phenomena related to the detection of glucose are illustrated based on the V g -I ds curves of different concentrations of glucose using a V ds of 0.1 V over a V g range from −0.3 to 0.3 V, as seen in Figure 8b. The rGO-GOx/LIG-IEs-based FET still exhibited p-type semiconductor properties when it was dropped with PBS solution without glucose, while the slope of the transfer curve, i.e., V g -I ds curve, increased significantly after the concentration of glucose increased to 0.5 mM. This is because the GOx immobilized on the sensing area reacted with glucose and produced gluconic acid and hydrogen peroxide (Equation 1); consequently, the resulting gluconic acid and hydrogen peroxide would be ionized into large amounts of e − (Equation 2). [55] Glucose Gluconicacid The e − produced by the above reactions is negatively charged, which attracts hole carriers in rGO. This leads to a decrease in the hole carrier concentration in rGO, which indicates that rGO capped with e − becomes less p-doped (i.e., more n-doped) due to the decrease in the hole carrier concentration, finally resulting in an increase in the slope of the transfer curve. Therefore, as the concentration of glucose increased, more corresponding e − was produced, which consequently increased the resistance of the sensing interface.

Real-Time Analysis of Glucose in Sweat Secreted from the Skin
Usually, persons who have blood sugar disorders secrete a certain amount of glucose into their sweat. Therefore, it is of great instructive significance to obtain timely information on glucose content in human sweat. In this paper, the feasibility and reliability of the proposed flexible and wearable chemiresistive biosensor was further evaluated by the real-time analysis of glucose in sweat secreted by the skin of a healthy subject with a portable sensing device, in which the glucose content data in sweat can be sent and received in real time based on a Bluetooth module, as shown in Figure 9a,b. It must be noted that the optimum temperature and pH range of GOx are between 25 and 37°C and between pH 5.0 and pH 6.0, respectively, while the surface temperature of human skin is usually 33°C and the sweat ranges from pH 4.0 to pH 6.0, so it does not affect the detection performance of the sensor. [56] Two PDMS flexible films with specific patterns were used to encapsulate the naked sensor and achieve the backflow storage of sweat, which effectively avoided interference from the external environment on the sensing area of LIG-IEs, as shown in Figure 9c. In addition, to reduce the damage of the long-term attachment of the sensor on human skin to normal sweating and skin permeability, medical body-friendly skin tape was applied to fix the sensor on skin. In order to validate the applicability of the sensor, we opted for subjects diagnosed with type T 2 DM diabete and prepared for five hours to measure changes in glucose concentration in sweat before, during and after meals by using FWFCGB, which was used to analyze GLC concentrations. As shown in Figure 9d, the glucose concentration in the sweat of the patient continued to rise within 1 h after a meal. Through the calculation of linear fitting model, it was found that the glucose concentration in the sweat finally stabilized at 1.03 mM 1 h after a meal.The actual test results show that the proposed FWFCGB is useful for detecting glucose in sweat to indirectly monitor blood glucose levels, which provides a new theoretical and methodological basis for engineering enzymes to produce tailor-designed biorecognition molecules for the development of wearable human health monitoring equipment.

Conclusion
In this work, a flexible and wearable FET-based rGO-GOx/LIG-IEs chemiresistive biosensor was developed for the real-time analysis of glucose in sweat on the surface of human skin based on a novel detection strategy of p-type rGO semiconducting sensing film. An efficient method for the batch fabrication of biosensors was developed by laser inducing, wherein the IEs can be easily patterned on a PI film by controlling the laser power and scanning rate. The rGO sensing film was modified on the surface of IEs based on the thermal reduction of GO, which was further functionalized with GOx by chemical cross-linking to obtain the rGO-GOx/LIG-IEs biosensor. Additionally, the developed biosensor was further encapsulated with double-layer patterned PDMS films and medical body-friendly skin tape to meet the requirements of wearable and noninvasive analysis of glucose. The developed flexible and wearable chemiresistive rGO-GOx/LIG-IEs biosensoer could be used to detect glucose over a concentration range of 0.01 to 3.0 mM with a LOD of 0.8 μM, which mostly covered the reference range of medical examination and screen tests for diabetes diagnosis and held great development prospects and broad application space in the field of wearable, continuous and real-time monitoring of glucose for the identification of diabetes with good resolution.

Supporting Information
Supporting Information is available from the Wiley Online Library or from the author.