A wearable sweat real-time detection device and detection method thereof

Abstract

The present invention provides a wearable real-time sweat detection device and a detection method thereof. Using PDMS as a flexible substrate material, a microfluidic sweat collection channel is designed, and a wedge-shaped structure and a layered capillary pump are added to accelerate the collection efficiency of sweat. At the same time, PANI is used as a substrate, and a Cu-MOF glucose sensor is constructed on it. The detection performance of the sensor is greatly improved, and it has good selectivity and long-term stability. At the same time, combined with sodium and potassium ion selective electrodes, instant detection of sodium, potassium and glucose in sweat can be achieved. Data transmission is carried out through Bluetooth, which can be used for real-time monitoring of physical health information of pilots during long-term missions.

Description

Technical Field

[0001] The present invention belongs to the technical field of sweat analysis and detection, and in particular relates to a wearable sweat real-time detection device and a detection method thereof. Background Art

[0002] When pilots are on missions for a long time, even the slightest change in their physical condition will have a huge impact on flight safety. Therefore, real-time health monitoring of pilots is particularly important. Traditional health monitoring usually involves regular visits to hospitals for checkups. Although this method can provide certain health information, it cannot achieve real-time dynamic health monitoring during flight. The ease of sweat production provides new possibilities for wearable real-time health monitoring. Figure 2 As shown. Sweat is a physiological fluid secreted by epidermal glands, containing electrolytes, glucose, lactic acid and a small amount of protein, etc. These biomarkers are closely related to the physiological state of the human body. Since sweat evaporates easily and the concentration of biomarkers in it is low, inaccurate measurement will bring challenges to establishing the relationship between biomarker concentration and health status, thus placing higher requirements on the efficiency and sensitivity of sweat collection.

[0003] There are many existing sweat detection technologies, mainly colorimetric and electrochemical methods to detect the concentration of biomarkers in sweat. Among them, the preparation technology of sodium and potassium ion sensors is basically mature, and for glucose concentration detection, the most commonly used detection method is to use glucose oxidase as a probe for detection, but enzyme probes have high requirements for the detection environment and are prone to inactivation. Currently, emerging non-enzyme sensors, such as metals and their derivatives, have lower detection limits, but complex synthesis methods and high-cost sensing materials will limit their application in portable and wearable devices.

[0004] Efficient sweat collection efficiency is a prerequisite for accurately detecting the concentrations of sodium, potassium and glucose in sweat. Most existing sweat collection devices are flexible patches made of materials such as hydrogel, PDMS or paper, which contain a sweat collection channel and a detection function. Improving sweat collection efficiency mostly involves expanding the sweat collection entrance or enhancing the transmission efficiency of the microfluidic channel. Although the above-mentioned researchers have made some improvements to the sweat collection channel, the collection time from the entrance to the detection function is long, which may cause sweat evaporation or reflux, resulting in inaccurate test results. Therefore, there is an urgent need for a wearable sweat detection application that can improve sweat collection speed and detection sensitivity. Summary of the invention

[0005] In view of the above problems in the prior art, the present invention provides a wearable real-time sweat detection device and a detection method thereof, so as to improve the sweat collection speed and the sensitivity of sweat detection.

[0006] The present invention is implemented by the following technical solutions:

[0007] The present invention provides a wearable real-time sweat detection device, a microfluidic chip, the microfluidic chip comprises at least two layers, the two layers of chips are vertically laminated up and down, the lower layer is a microfluidic chip covering layer, the microfluidic chip covering layer is provided with sweat collection holes, the sweat collection holes are laminated to the upper layer of the microfluidic chip, the lower surface of the upper chip is provided with a sweat collection channel, the front end of the sweat collection channel is a wedge-shaped structure, the rear end of the sweat collection channel is provided with a capillary pump, the capillary pump is an upper and lower two-layer channel, the upper channel is formed by a short convex portion and inner walls on both sides of the channel, the lower channel is formed by a plurality of short convex portions arranged at intervals, the plurality of short convex portions are arranged at the same intervals in the transverse direction, the upper channel of the capillary pump is a detection function part, the detection function part is a cavity, and the rear end of the capillary pump is connected to a sweat outlet channel;

[0008] A screen-printed electrode is arranged between two adjacent layers of the microfluidic chip, the screen-printed electrode comprises a working electrode, a reference electrode and a counter electrode, the working electrode is modified by Cu-MOF / PANI and a sodium and potassium ion selective membrane, the reference electrode is a silver paste, the counter electrode is a carbon paste, each electrode is connected to an external circuit through a corresponding lead, the screen-printed electrode is connected to a multi-parameter detector circuit through a detector electrode interface, the multi-parameter detector circuit is connected to a Bluetooth transmission module, and the multi-parameter detector circuit is powered by a rechargeable lithium battery;

[0009] The microfluidic chip and the screen-printed electrode constitute a sweat collection flexible patch; the sweat collection flexible patch, the multi-parameter detection circuit system, the battery, and the Bluetooth transmission module are integrated to form a sweat detection functional system.

[0010] Furthermore, the microfluidic chip uses PDMS as a flexible substrate material.

[0011] Furthermore, the sweat collection channel has a dimension D of 0.2 mm in width with a floating value of ±0.02 mm, a dimension H of 0.25 mm in depth with a floating value of ±0.01 mm, and an extension line angle α of the wedge-shaped channel is 1.38° with a floating value of ±0.02°.

[0012] Furthermore, the height of the upper channel of the capillary pump is L2, the height of the lower channel of the capillary pump is L3, and the ratio of L3 / L2 is 0.25.

[0013] Furthermore, the detection limit of the glucose sensor of the screen-printed electrode is 1.94 μM.

[0014] The present invention also provides a wearable sweat real-time detection method, which uses the above-mentioned detection device and includes the following steps:

[0015] S1 Sweat Collection

[0016] The wearable real-time sweat detection device is worn on the surface of the human body. The sweat on the skin surface flows into the upper layer of the microfluidic chip through the sweat collection holes on the microfluidic chip cover layer. After passing through the front end of the sweat collection channel, the Laplace force spontaneously formed by the structure causes the sweat to continue to flow backward along the collection channel until it reaches the capillary pump. After the lower channel of the capillary pump is filled with sweat, the sweat continues to flow along the upper channel and contacts the screen-printed electrode, and the excess sweat is discharged from the outlet channel.

[0017] S2 system power-on and initialization and reference electrode voltage setting

[0018] The device is powered by a lithium battery, and the output of the lithium battery is converted into voltage through a voltage regulator module; at the same time, the voltage regulator module converts the voltage again to power the microcontroller and other components that need power. After the microcontroller is powered on, it performs initialization operations, including configuring the I / O port, ADC module, LMP91000 parameter settings, UART interface for communication with the Bluetooth module, and hardware resources; after the voltage regulator chip generates voltage, it generates a stable voltage through a voltage divider circuit as the input voltage of the reference electrode in ion detection;

[0019] S3 ion and glucose detection

[0020] The potential output of the sodium and potassium ion working electrode is buffered by a buffer composed of an operational amplifier. The buffered analog signal is sent to the on-chip ADC of the STM32F103RCT6 for analog-to-digital conversion, converting the analog potential into a digital signal. The voltage regulator chip controls the bias voltage between the reference electrode and the working electrode through its internal circuit and keeps it constant.

[0021] Glucose detection: The current signal on the working electrode is converted into a voltage signal by the voltage regulator chip and amplified. The amplified voltage signal is converted into a digital signal again by the integrated circuit ADC. The three ADC channels of the microcontroller respectively collect the potential of the working electrode and reference electrode of the ion electrode and the output voltage of the electrochemical detection chip. Based on the voltage value and the preset calibration curve or algorithm, the microcontroller calculates the corresponding ion concentration and glucose concentration. The calculated concentration data is wirelessly transmitted to the mobile phone APP through the Bluetooth module to realize real-time display and recording of data.

[0022] S4 Interaction and Feedback

[0023] After the mobile phone APP receives the data, it displays the changes in ion concentration and glucose concentration in real time. The APP performs statistical analysis on the data over a long period of time and provides health advice or abnormal reminders. It realizes wireless transmission and real-time display of data through Bluetooth technology. Through the above steps, this sweat detection device can complete the detection of ion concentration and glucose concentration in sweat, and realize wireless transmission and real-time display of data through Bluetooth technology, providing health monitoring services.

[0024] Beneficial effects: In terms of glucose sensors, economical metal material Cu is used, Cu-MOF is synthesized using a homemade micromixer, and economical conductive material PANI is added as a substrate, thereby improving the sensing performance of the glucose sensor; in the design of the sweat collection patch, a wedge-shaped structure is used in the microfluidic channel part to make the sweat spontaneously flow to the detection functional part, thereby accelerating the transmission efficiency of the sweat, and at the same time, two layers of capillary pumps are designed in the detection functional part to enhance the capillary force of the microfluidic channel, thereby further enhancing the sweat collection efficiency; the sweat collection patch and the three-electrode detection circuit are integrated, and Bluetooth is used for signal transmission for use in a wearable sweat detection device. Therefore, the device provided by the present invention can realize real-time dynamic health monitoring during flight, thereby providing a strong guarantee for the health monitoring and flight safety of pilots. BRIEF DESCRIPTION OF THE DRAWINGS

[0025] Figure 1 This is a system principle diagram of a wearable real-time sweat detection device according to an embodiment of the present invention;

[0026] Figure 2 This is a diagram of the common sweat detection mechanism;

[0027] Figure 3 A detailed diagram of a sweat collection flexible patch according to an embodiment of the present invention;

[0028] Figure 4 This is a schematic diagram of the structure of a sweat collection flexible patch according to an embodiment of the present invention;

[0029] Figure 5 A 3D schematic diagram of a flat channel and a schematic diagram of PMMA after bonding according to an embodiment of the present invention;

[0030] Figure 6 A 3D schematic diagram of a wedge-shaped channel and a microfluidic chip entity according to an embodiment of the present invention;

[0031] Figure 7 Optimizing and analyzing data of a sweat collection channel according to an embodiment of the present invention;

[0032] Figure 8 A top view of a sweat collection channel according to an embodiment of the present invention;

[0033] Fig. 9is a 3D schematic diagram of a capillary pump according to an embodiment of the present invention;

[0034] Fig.10 A finite element analysis cloud diagram of a capillary pump according to an embodiment of the present invention;

[0035] Fig.11 Velocity analysis data of four cross sections of a capillary pump according to an embodiment of the present invention;

[0036] Fig.12 A schematic diagram of the volume fraction of a capillary pump according to an embodiment of the present invention;

[0037] Fig.13 This is an electron microscope image of Cu-MOF according to one embodiment of the present invention;

[0038] Fig.14 This is a graph of electrochemical test analysis data of a glassy carbon electrode according to an embodiment of the present invention;

[0039] Fig.15 This is an infrared analysis diagram of Cu-MOF and Cu-MOF / PANI according to an embodiment of the present invention;

[0040] Fig.16 This is a graph of electrochemical test analysis data of a screen-printed electrode according to an embodiment of the present invention;

[0041] Fig.17 is a cross-sectional view of a capillary pump according to an embodiment of the present invention;

[0042] Fig.18 This is a diagram of the test results of the test system according to an embodiment of the present invention.

[0043] In the figure: 1-capillary pump, 10-short convex part, 2-upper channel, 3-lower channel. DETAILED DESCRIPTION

[0044] In order to make the purpose, technical solution and advantages of this patent more clear, the following is a further detailed description of this patent in conjunction with specific embodiments. It should be understood that the specific embodiments described here are only used to explain this patent, and are not used to limit the present invention.

[0045] The present invention is implemented through the following technical scheme: a wearable real-time sweat detection device, a microfluidic chip includes at least two layers, the two layers of chips are vertically attached to each other, the lower layer is a microfluidic chip covering layer, the microfluidic chip covering layer is provided with sweat collection holes, the sweat collection holes are attached to the upper layer of the microfluidic chip, the lower surface of the upper chip is provided with a sweat collection channel, the front end of the sweat collection channel is a wedge-shaped structure, the collection channel of the wedge-shaped structure helps to guide sweat to flow quickly and evenly into subsequent detection functional parts, reducing the risk of evaporation and contamination, the rear end of the sweat collection channel is provided with a capillary pump 1, the The capillary pump 1 utilizes the principle of capillary action to achieve directional flow of sweat without external power. The capillary pump 1 is a two-layer channel, the upper channel 2 is formed by a short protrusion 10 and the inner walls on both sides of the channel, and the lower channel 3 is formed by a plurality of short protrusions 10 arranged at intervals, and the plurality of short protrusions 10 are arranged at the same intervals in the horizontal direction. The upper channel 2 of the capillary pump 1 is a detection functional part, and the detection functional part is a hollow cavity. The rear end of the capillary pump 1 is connected to the sweat outlet channel, and the outlet channel ensures that the sweat after detection can be discharged smoothly to avoid channel blockage. The top view of the sweat collection channel is shown in FIG. Figure 8 As shown;

[0046] A screen-printed electrode is arranged between two adjacent layers of the microfluidic chip, and the screen-printed electrode contains a working electrode, a reference electrode, and a counter electrode. The working electrode is modified by Cu-MOF / PANI and a sodium and potassium ion selective membrane, the reference electrode is a silver paste, and the counter electrode is a carbon paste. Each of the electrodes is connected to an external circuit through a corresponding lead, and the screen-printed electrode is connected to a multi-parameter detector circuit through a detector electrode interface. The multi-parameter detector circuit is connected to a Bluetooth transmission module. The Bluetooth transmission does not require a wired connection, which reduces the sense of restraint when worn and improves the comfort and convenience of the device. At the same time, the multi-parameter detector circuit is powered by a rechargeable lithium battery;

[0047] The microfluidic chip and the screen-printed electrode constitute a sweat collection flexible patch, such as Figure 3 and Figure 4 As shown; the sweat collection flexible patch, the multi-parameter detection circuit system, the battery, and the Bluetooth transmission module are integrated to form a sweat detection function system. The system principle of the wearable sweat real-time detection device is as follows Figure 1 As shown, multiple detection functions are integrated into one circuit, which realizes the synchronous detection of multiple parameters in sweat and improves the detection efficiency and accuracy.

[0048] A method for preparing a wearable real-time sweat detection device comprises the following steps:

[0049] (1) Optimization and preparation of microfluidic chips

[0050] The specific steps are as follows:

[0051] 1. Conduct comparative experiments on flat channels and obtain samples such as Figure 5 After the display, determine the best width D of the channel, first fix the depth H to 0.3mm, and the incremental width D to 0.1-0.7mm, that is, Figure 7 The depth H of channels 1-7 in d is fixed at 0.3 mm, and the width D is 0.1-0.7 mm respectively. Figure 7 g in the figure shows the solution progress image at T = 60s, which is different from Figure 7 The width D in d is 0.2mm by analyzing the arrival time of the end point. Then the fixed channel width D is 0.2mm, and the depth H value of the channel is changed. The incremental depth H is 0.1-0.3mm. Figure 7 h in Figure 2 shows the solution progress image at T = 30 s, which is different from Figure 7 The depth H in e is 0.25 mm by analyzing the time of reaching the end point;

[0052] After determining the optimal depth and width of the collection channel, the overall sweat collection device is designed, and a wedge-shaped channel is used at the front end of the sweat collection channel. Figure 6 The 3D schematic diagram of the wedge-shaped channel is shown. D is 0.2 mm, H is 0.25 mm and remains unchanged. The channel length L is 30 mm. The extension line angle α of the wedge-shaped channel is optimized to determine the optimal α value. Figure 6 The left part is the 3D image of the wedge-shaped channel, D is 0.2mm wide, H is 0.25mm deep, and the incremental angle α is 0°-1.41°. Figure 6 The right part of the figure shows the PDMS and PMMA chips. Microchannels with different extension angles α were prepared on PMMA. The microchannels were engraved on the PMMA chip using a precision CNC machine tool, and the PDMS layer was used as a covering layer to seal the microchannels. PDMS and PMMA need to be repeatedly cleaned with deionized water and ethanol, and no additional bonding process is required. PMMA is hydrophilicized using a plasma cleaner, and then PDMS is attached. Finally, a hole puncher is used to manually punch holes on the PDMS. The test is performed with a pipette, and a mobile phone is used to record the process of the solution from the inlet to filling the entire detection function part. Figure 7 The i in the figure shows the solution progress image at T = 7s, which is different from Figure 7 The angle α in f is 1.38° by analyzing the arrival time of the endpoint;

[0053] Design a two-layer capillary pump. The sweat will first fill the bottom capillary channel. After it is filled, a layer of water film will form at the bottom of the upper capillary. The design diagram is shown in the figure. Fig.17 As shown, Fig. 9This is a 3D schematic diagram of one of the capillary pumps, where a=0.45mm, b=3.55mm, L1=0.25mm, and the values ​​of L3 and L2 are changed to change the ratio of L3 / L2, and L3 / L2=0, L3 / L2=0.2, L3 / L2=0.25, L3 / L2=0.3, L3 / L2=0.38888, L3 / L2=0.5, L3 / L2=1 are taken respectively. Comsol6.0 is used for two-phase flow simulation, and capillary force is used to drive the water to flow to the air area. The time is 0.003s. Fig.10 Four sections were selected for velocity analysis, namely x=-0.073m, x=-0.075m, x=-0.074m, and x=-0.0745m. Fig.11 Figures e, f, g, and h are velocity analysis diagrams of the four cross sections. It can be seen from these diagrams that when the ratio of L3 / L2 is 0.25, that is, L3=0.05mm, L2=0.20mm, the velocity is ahead compared with the other L3 / L2. Fig.12 The volume fraction diagrams of L3 / L2=0, L3 / L2=0.25, and L3 / L2=1 are selected. It can be seen that when L3 / L2=0.25, the area where the water flows is the longest. Therefore, we can determine that L3 / L2=0.25;

[0054] 2. Carve microfluidic channels of different sizes onto PMMA, then ultrasonically clean two sheets of PMMA with ethanol solution and water for five minutes, and then put them into a drying oven at 60°C for 10 minutes to remove residual moisture on the surface, take out two sheets of dried PMMA, and bond them with a bonding agent (prepared by 1,2-dichloroethane and 98% ethanol in a weight ratio of 1:4), align the two PMMA plates, and slowly inject the bonding agent into the gap between the two sheets of PMMA through a syringe. When the capillary force phenomenon is observed, clamp the bonded chip with multiple clips to make the overall force uniform, and then put it into a drying oven and keep it at 60°C for 5 minutes; After a single PMMA chip is made, use the PDMS layer as a cover to seal the microchannel. PDMS and PMMA need to be repeatedly cleaned with deionized water and ethanol, and then hydrophilized. After the hydrophilic treatment is completed, they can be directly bonded;

[0055] 3. Pour the mixture of PDMS (liquid) and crosslinker (cured PDMS, commonly used sylgard184) into a test tube at a ratio of 10:1, stir thoroughly with a glass rod, and use a drying oven for vacuum treatment to remove bubbles in the PDMS mixture after it is uniform; pour the PDMS mixture into the mold after cleaning it, and heat it in a drying oven at 60°C for 3 hours. Wait for the PDMS to harden and then take it out. After the mold and PDMS cool down, cut the substrate, peel off the PDMS to get the device, and obtain a replica of the microchannel on the PDMS block; use a PDMS puncher to process the fluid inlet and fluid outlet on the microfluidic; use a plasma cleaner to treat the surface of the PDMS block with microchannels and slides to remove surface impurities; finally, perform PDMS bonding, and use a 2kg weight to press on the two PDMS blocks for 24 hours to make them firmly bonded.

[0056] (2) Preparation of screen-printed electrodes

[0057] The specific steps are as follows:

[0058] 1. Optimization of Cu-MOF. To determine the structure of Cu-MOF, two low flow rates were used for the synthesis of Cu-MOF, namely 1519 μL / min (Re=100) and 4557 μL / min (Re=200). To observe the microstructure of Cu-MOF, the synthesized Cu-MOF powder was tested by scanning electron microscopy. Fig.13 a and b in the figure. As can be seen from the figure, the Cu-MOF particles synthesized at low flow rates are uneven and contain too many impurities, which are not suitable for the preparation of sensors. In response to this, the flow rate was further increased, and two high flow rates of 7595μL / min (Re=500) and 12152μL / min (Re=800) were used to synthesize Cu-MOF. After the synthesized Cu-MOF was subjected to the preparation process, Cu-MOF powders at different flow rates were obtained, and then subjected to scanning electron microscopy tests, and the results were obtained. Fig.13 c and d in;

[0059] For the preparation of Cu-MOF, a microfluidic mixer was used for preparation. First, 10 mL of 40 mM 2-methylimidazole solution was prepared as reagent one, and 10 mL of 10 mM copper nitrate solution was prepared as reagent two. Using the designed microfluidic channel, a syringe pump was used to simultaneously inject the two reagents from two inlets at a speed of 12152 μL / min (Re=800). 20 mL of a blue solution containing Cu-MOF was obtained from the outlet. The obtained solution was placed in a high-speed centrifuge, the parameters were set to 10000 r / min, and centrifuged for 5 hours to obtain Cu-MOF precipitate and supernatant. The supernatant was removed, and the Cu-MOF was placed in a 60°C constant temperature box and vacuum dried for half an hour to obtain Cu-MOF powder;

[0060] 2. For the first PANI conductive layer of the glucose sensor, a thin PANI film was formed by continuous dripping test on the glassy carbon electrode. Four concentrations of Cu-MOF were configured, namely 10, 20, 30, and 40 mg / mL. Time-current and EIS tests were performed using an electrochemical workstation. When performing the time-current test, the test voltage was 0.5 V, the test time was 1600 s, and three-electrode detection was used, with the glassy carbon electrode modified with Cu-MOF as the working electrode, the Pt electrode as the counter electrode, and the Ag / AgCl electrode as the reference electrode. The three electrodes were placed in a 0.1 M NaOH solution, and the test was first performed for 200 s. After the current was stable, a certain amount of glucose solution was added to the solution every 100 s to form a step-like concentration gradient curve. The test glucose concentration range was 0-1 mM. The test curve is as follows: Fig.14 As shown in a in the figure, the glucose sensors with the three modification methods were then electrochemically tested. The three electrodes were tested for time-current according to the above method. The test results are shown in Fig.14 As shown in b;

[0061] The PANI / Cu-MOF glucose sensor was subjected to cyclic voltammetry test with a test voltage of 0-0.8V, a scan rate of 50mV / s, and a scan number of 20 cycles. It was placed in a 0.1M NaOH solution for testing and the test curve was recorded. Glucose solution was then added to the solution to prepare a 0.1M NaOH solution containing 1mM glucose. The test was then conducted and the test curve was recorded. Finally, the last cycle of the two test curves was plotted to obtain the CV test curve. Fig.14 c, the PANI / Cu-MOF electrode was tested for time-current curve in 5mL 0.1M NaOH solution, the test voltage was set to 0.5V, the time was 1600s, and the test curve results were as follows Fig.14 As shown in d, in order to obtain linear correlation and sensitivity, a linear fitting was performed on the test results, and the fitting curve is shown in Fig.14 e in the figure, it can be seen from this figure that the fitting degree reaches 0.99314, and the detection limit (LOD) is 1.94μM (S / N=3), which can be simplified to LOD=3σ / S, where σ is the standard deviation of the response and S is the slope of the calibration curve; at the same time, for the long-term stability of the PANI / Cu-MOF electrode, the prepared PANI / Cu-MOF electrode was subjected to DPV test in a NaOH solution containing 1mM glucose, with a test voltage of 0-0.8V. The electrode was subjected to a DPV test every 1 hour for five tests, and the voltage value data at 0.5V was obtained. Fig.14 g in the figure, and infrared spectroscopy analysis was performed on Cu-MOF and Cu-MOF / PANI samples. Fig.15It can be seen that the absorption peak of Cu-MOF sample is found at 3417.24. 3100-3600cm -1 The bands near are attributed to -OH and HOH stretching vibrations, which are produced by water molecules adsorbed on the Cu-MOF surface. In addition, the Cu-MOF sample also has absorption peaks near 2920.02, 1617.12, 1382.05, 745.49, 606.56, 509.24, and 430.68. In addition, 1617.12 and 1382.05 cm -1 The bands nearby are the HOH bending vibration in the H2O molecule or the OH deformation vibration of the surface hydroxyl group, indicating the presence of carboxylic acid groups. 745.49 and 606.56 cm -1 The nearby bands come from the stretching vibration of Cu-O when oxygen atoms coordinate with copper atoms.

[0062] Some of the vibration frequencies of the Cu-MOF / PANI sample have higher absorbance than bare Cu-MOF. It can be clearly seen that the absorbance of the peak of the nanocomposite is much higher than that of bare Cu-MOF. The broadening of the peak can also be observed in different bands, especially in the OH band around 3450cm −1 The increase in the intensity of the absorbance band can be attributed to the effect of hydrogen bonding. In addition, the nanocomposites have −1 The presence of a broad band in the region is attributed to the vibrational modes associated with the nitrogen bonds (i.e., due to the presence of the conductive polymer). In addition, the nanocomposite spectrum also shows typical absorption bands for both parent materials. −1 The band at 1308 cm corresponds to the bending and stretching vibration modes of Cu−Cu and Cu−O. Similarly, in the Cu-MOF / PANI nanocomposite, the -1 The peak at is related to the CN stretching vibration of the quinone benzene ring;

[0063] 3. Preparation of PANI / Cu-MOF glucose sensor: 10 μL of 0.5 mg / mL PANI was added to the glassy carbon electrode and dried at room temperature to form a PANI film. 10 μL of 20 mg / mL Cu-MOF solution was added on top and dried at room temperature to form a Cu-MOF film. 10 μL of Nafion membrane solution was added to the surface of the modified electrode and dried at room temperature.

[0064] 4. Preparation of sodium and potassium ion sensors. Pre-treat the glassy carbon electrode. First, polish the glassy carbon electrode (GCE) on 1μm, 0.3μm, and 0.05μm alumina polishing powder, respectively. Then, use ethanol (v:v=1:1), HNO3 (v:v=1:1), and plasma aqueous solution for ultrasonic cleaning for 30s each. Prepare sodium and potassium ion selective membranes on the glassy carbon electrode. Before modifying the ion selective membrane, it is necessary to cover the surface of the glassy carbon electrode with a layer of PEDOT:PSS membrane. Prepare a mixed solution containing 0.01MEDOT and 0.1MNaPSS, place the glassy carbon electrode in the mixed solution, and perform electrodeposition by constant current electrodeposition. Set the current to high. The minimum current was 14 μA, and the electrodeposition was performed for 700 s to form a layer of PEDOT:PSS film on the surface of the glassy carbon electrode. Then, a sodium ion selective membrane was modified on the surface of the glassy carbon electrode. 1 mg of sodium ion carrier, 0.55 mg of Na-TFPB, 65.45 mg of DOS, and 33 mg of PVC were dissolved in 660 μL of cyclohexanone solution. The solution was placed in an ultrasonic cleaner and vibrated for 30 min until it was completely dissolved. For the potassium ion selective membrane, 2 mg of valinomycin, 0.5 mg of NaTPB, 32.7 mg of PVC, and 64.7 mg of DOS were dissolved in 660 μL of cyclohexanone solution. The solution was placed in an ultrasonic cleaner and vibrated for 30 min until it was completely dissolved.

[0065] (3) Real-time detection of the device

[0066] The specific steps are as follows:

[0067] The sweat collection channel and the multi-parameter electrode detection circuit system are integrated to form a wearable sweat real-time detection device. First, the screen-printed electrodes of the sodium, potassium and glucose sensors are tested. Fig.16 a and b in the figure are the test curves of screen-printed electrodes for sodium and potassium ions, respectively. Fig.16 c and d in the figure are the test and fitting curves of the screen-printed electrode of the glucose sensor, respectively. Fig.16 From c and d in the figure, we can find that the two test curves are consistent with the above Fig.14 The curves of d and e are basically consistent, and the linearity is above 0.99. The multi-parameter detection circuit is combined with the PDMS sweat collection microfluidic chip prepared above. The whole system is powered by a battery and the test data is transmitted using Bluetooth. Finally, a simulation test was performed using 80μM glucose, 60μM sodium and 10μM potassium. The test results are shown in Figure 2. Fig.18 As shown, it can be seen from the App interface that the concentrations of the three markers tested are basically maintained at the corresponding concentrations, with small fluctuations, indicating that the system has good detection stability and can be used for real-time detection of biomarkers in actual sweat.

[0068] A wearable sweat real-time detection method, using the wearable sweat real-time detection device, comprises the following steps:

[0069] S1

[0070] When the pilot wears the device for activities, the sweat on the skin surface flows into the upper layer of the microfluidic chip through the sweat collection holes on the microfluidic chip cover layer. After passing through the front end of the sweat collection channel, the Laplace force spontaneously formed by the structure causes the sweat to continue to flow backward along the collection channel until it reaches the capillary pump. After the lower channel of the capillary pump is filled with sweat, the sweat continues to flow along the upper channel and contacts the screen-printed electrode, and the excess sweat is discharged from the outlet channel.

[0071] S2 system power-on and initialization and reference electrode voltage setting

[0072] Powered by a 3.7V lithium battery, the voltage regulator module generates 2.5V and 3.3V for reference electrode and microcontroller, etc. After the microcontroller is powered on, the system configuration is initialized; LM4120-2.5 generates 2.5V, which is divided into 0.625V as the reference electrode voltage.

[0073] S3

[0074] The working electrode potential is buffered by AD8606 and converted into a digital signal by the MCU ADC. For glucose detection, LMP91000 controls the bias voltage to 0.5V, converts and amplifies the working electrode current signal, and then converts it by the MCU ADC. The MCU calculates the open circuit potential and the LMP91000 output voltage and converts them into concentration data.

[0075] S4

[0076] The concentration data is sent to the mobile phone APP via Bluetooth and displayed and recorded in real time.

[0077] In this embodiment, the microfluidic chip uses PDMS as the flexible base material.

[0078] In this embodiment, the sweat collection channel dimension D is 0.2 mm wide, with a floating value of ±0.02 mm, H is 0.25 mm deep, with a floating value of ±0.01 mm, and the extension line angle α of the wedge-shaped channel is 1.38°, with a floating value of ±0.02°.

[0079] In this embodiment, the height of the upper channel of the capillary pump is L2, the height of the lower channel of the capillary pump is L3, and the ratio of L3 / L2 is 0.25.

[0080] In this embodiment, the detection limit of the glucose sensor of the screen-printed electrode is 1.94 μM.

[0081] Compared with the prior art, in this embodiment, a wedge-shaped structure is adopted in the front end part of the microfluidic channel to make the sweat spontaneously flow toward the detection functional part, thereby accelerating the transmission rate of the sweat. At the same time, two layers of capillary pumps are designed in the detection functional part to enhance the capillary force of the microfluidic channel, thereby further enhancing the sweat collection efficiency. For the glucose sensor, the non-enzymatic economical metal material Cu is used, and Cu-MOF is synthesized using a homemade micromixer. At the same time, the conductive material PANI is used as the substrate to enhance the sensing performance of the glucose sensor.

[0082] In summary, a wearable real-time sweat detection device, through its sophisticated structural design and advanced technology application, can achieve rapid and accurate detection of multiple parameters in pilots' sweat, and has the advantages of wireless transmission and portable use, providing strong protection for pilots' health monitoring and flight safety.

Claims

1. A wearable real-time sweat detection device, characterized in that: A microfluidic chip is provided, wherein the microfluidic chip comprises at least two layers, wherein the two layers of the chip are vertically laminated up and down, wherein the lower layer is a microfluidic chip covering layer, wherein the microfluidic chip covering layer is provided with sweat collecting holes, wherein the sweat collecting holes are laminated with the upper layer of the microfluidic chip, wherein the lower surface of the upper layer chip is provided with a sweat collecting channel, wherein the front end of the sweat collecting channel is a wedge-shaped structure, wherein the rear end of the sweat collecting channel is provided with a capillary pump, wherein the capillary pump is an upper and lower two-layer channel, wherein the upper layer channel is formed by the upper edge of a short convex portion and the inner walls on both sides of the channel, wherein the lower layer channel is formed by a plurality of short convex portions arranged at intervals, wherein the plurality of short convex portions are arranged at the same intervals in the transverse direction, wherein the upper layer channel of the capillary pump is a detection function component, wherein the detection function component is a hollow cavity, and wherein the rear end of the capillary pump is connected to a sweat outlet channel; A screen-printed electrode is arranged between two adjacent layers of the microfluidic chip, the screen-printed electrode comprises a working electrode, a reference electrode and a counter electrode, the working electrode is modified by Cu-MOF / PANI and a sodium and potassium ion selective membrane, the reference electrode is a silver paste, the counter electrode is a carbon paste, each electrode is connected to an external circuit through a corresponding lead, the screen-printed electrode is connected to a multi-parameter detector circuit through a detector electrode interface, the multi-parameter detector circuit is connected to a Bluetooth transmission module, and the multi-parameter detector circuit is powered by a rechargeable lithium battery; The microfluidic chip and the screen-printed electrode constitute a sweat collection flexible patch; the sweat collection flexible patch, the multi-parameter detector circuit, the battery, and the Bluetooth transmission module are integrated to form a sweat detection functional system.

2. A wearable real-time sweat detection device according to claim 1, characterized in that: The microfluidic chip uses PDMS as a flexible base material.

3. A wearable real-time sweat detection device according to claim 1, characterized in that: The width D of the sweat collection channel is 0.2 mm, with a floating value of ±0.02 mm, the depth H is 0.25 mm, with a floating value of ±0.01 mm, and the extension line angle α of the wedge-shaped channel is 1.38°, with a floating value of ±0.02°.

4. A wearable real-time sweat detection device according to claim 1, characterized in that: The height of the upper channel of the capillary pump is L2, the height of the lower channel of the capillary pump is L3, and the ratio of L3 / L2 is 0.

25.

5. A wearable real-time sweat detection device according to claim 1, characterized in that: The detection limit of the glucose sensor of the screen-printed electrode is 1.94 μM.

6. A wearable real-time sweat detection method, using the wearable real-time sweat detection device according to claim 1, characterized in that: The following steps are involved: S1 Sweat Collection The wearable real-time sweat detection device is worn on the surface of the human body, and the sweat on the skin surface flows into the upper layer of the microfluidic chip through the sweat collection holes on the cover layer of the microfluidic chip. After passing through the front end of the sweat collection channel, the Laplace force spontaneously formed by the structure causes the sweat to continue to flow backward along the collection channel until it reaches the capillary pump. After the lower channel of the capillary pump is filled with sweat, the sweat continues to flow along the upper channel and contacts the screen-printed electrode, and the excess sweat is discharged from the outlet channel; S2 system power-on and initialization and reference electrode voltage setting The device is powered by a lithium battery, and the output of the lithium battery is converted into voltage through a voltage regulator module; the voltage regulator module converts the voltage again to power the microcontroller and other components that need power. After the microcontroller is powered on, it performs initialization operations, including configuring the I / O port, ADC module, LMP91000 parameter settings, UART interface for communicating with the Bluetooth module, and hardware resources; after the voltage regulator chip generates voltage, it generates a stable voltage through a voltage divider circuit as the input voltage of the reference electrode in ion detection; S3 ion and glucose detection The potential output of the sodium and potassium ion working electrode is buffered by a buffer composed of an operational amplifier. The buffered analog signal is sent to the on-chip ADC of the integrated circuit for analog-to-digital conversion, and the analog potential is converted into a digital signal. The voltage regulator chip controls the bias voltage between the reference electrode and the working electrode through its internal circuit and keeps it constant. Glucose detection: The current signal on the working electrode is converted into a voltage signal by the voltage regulator chip and amplified. The amplified voltage signal is converted into a digital signal again by the integrated circuit ADC. The three ADC channels of the microcontroller respectively collect the potential of the working electrode and reference electrode of the ion electrode and the output voltage of the electrochemical detection chip. Based on the voltage value and the preset calibration curve or algorithm, the microcontroller calculates the corresponding ion concentration and glucose concentration. The calculated concentration data is wirelessly transmitted to the mobile phone APP through the Bluetooth module to realize real-time display and recording of data. S4 Interaction and Feedback After the mobile phone APP receives the data, it displays the changes in ion concentration and glucose concentration in real time. The APP performs statistical analysis on the data over a long period of time and provides health advice or abnormal reminders. It realizes wireless transmission and real-time display of data through Bluetooth technology. Through the above steps, the sweat detection device can complete the detection of ion concentration and glucose concentration in sweat, and realize wireless transmission and real-time display of data through Bluetooth technology, providing health monitoring services.

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