Integrated sweat sensing system for health condition monitoring and safety alerts

Abstract

The present invention provides an integrated sweat sensing system comprising a multi-layer structure containing a flexible circuit sensing layer with multiple electronic and mechanical components for sensing specific physiological changes of a user and giving feedback signals to the user when the detectable biological, chemical or physiological signals exceed a certain threshold. The system also comprises a microfluidic based biosensing platform for sensing changes in sweat composition components including changes in ions, glucose and pH, which subsequently triggers feedback signal output from the flexible circuit sensing layer, thereby providing corresponding status change signals to the user. The system is bendable and can be detachably mounted on multiple skin areas of a user.

Description

TECHNICAL FIELD

[0001] The present invention relates to an integrated sweat sensing system for health condition monitoring and safety alerting, in particular, a wearable integrated sweat sensing system for detecting compositional changes in different sweat constituents of an individual as well as other physiological changes. BACKGROUND

[0002] In modern society, personal health is an important aspect that many people pay attention to in our daily life, which creates great market potential. Therefore, many consumer products have been implanted with functions for tracking and even analyzing health-related data received from individuals. Since biological fluids such as sweat contain a large number of biomarkers reflecting the health status of individuals, for example, Na + responsible for regulating acid-base balance and maintaining muscle and heart function; glucose in sweat indicates the blood glucose content in the human body, and recent studies have focused on different flexible and sensitive biosensors for detecting a variety of sweat constituents such as Na + , Cl - , K + , NH4 + , glucose, pH, lactate, etc. The integration of these biosensors in wearable devices can provide an alternative, non-invasive, real-time platform to monitor the health status of individuals. In order to integrate multiple biosensors in a single wearable device for sweat sensing, at least the following features should be considered: (a) a flexible substrate for conformal mounting on the curved human skin and allowing the construction of sensing components thereon; (b) a multiple sensing mechanism that allows sensing of the main chemical substances in sweat, as well as a processing mechanism capable of providing a comprehensive health assessment in real time; (c) a self-alerting system that provides a synchronous reminder to the user when abnormal physiological signals are detected during sweat monitoring; (d) a carefully designed microfluidic as an interface between the human skin and the biosensor for spontaneously directing fresh sweat to the target sensor area quickly.

[0003] One of the challenges faced by most conventional skin-integrated biosensors is how to avoid false sensing or interpretation of different biomarkers from sweat. Providing accurate health assessment becomes a major obstacle in developing the next generation of wearable sweat sensors. Another challenge is how to provide an effective sampling and sensing platform while not compromising the flexibility and durability of the platform.

[0004] Therefore, there is a need for an improved integrated system having a variety of sensing mechanisms and mapping functions suitable for sensing complex biological fluids such as sweat, thereby at least reducing or eliminating the above-mentioned drawbacks and problems. SUMMARY

[0005] Accordingly, the first aspect of the present application provides a flexible, detachably mountable integrated sensing system on the skin of a human body for simultaneous differentiation of multiple sweat components in order to provide real-time monitoring and assessment of the individual's health condition. The system is a multi-layered assembly comprising, from the nearest layer to the contact surface with the human body skin:

[0006] an adhesive layer for enhancing the adhesion of the system to the human body skin;

[0007] a sensing layer comprising:

[0008] a flexible printed circuit board (FPCB) having disposed thereon a plurality of electronic and mechanical components including at least a mechanical actuator for outputting a vibration signal to the user, a microcontroller (MCU), and a plurality of feedback signal outputs each producing at least two opposite visual, audio, or audio-visual signals representing at least two different states of each respective biological, chemical, or physiological signal that the corresponding biosensor is capable of detecting; and a biomarker sensing platform having a plurality of biosensors associated with the electrode layer and a microfluidic, wherein the microfluidic comprises a plurality of microchambers, a microfilter that is an interface between the microfluidic and the contact surface with the human body skin, and a plurality of microchannels connecting the corresponding microchambers together and with the microfilter, the microchambers being disposed in matching positions with the biosensors on the biomarker sensing platform to provide a shortest path for the sweat to flow from the contact surface with the skin to the corresponding biosensors and to avoid any contamination of the sweat sample; and

[0009] an encapsulation layer.

[0010] In particular embodiments, the biomarker sensing platform is configured to be flexible, bendable, and detachably mountable on the contact surface of the human body skin.

[0011] In particular embodiments, the biomarker sensing platform is integrated to the FPCB forming a continuous layer with the circuitry area (or control panel) of the FPCB having a plurality of electronic and / or mechanical components soldered thereon.

[0012] In particular embodiments, the biomarker sensing platform is detachably connected to the circuitry area of the FPCB for forming an extension of the sensing layer.

[0013] In particular embodiments, the electrode layer of the biomarker sensing platform is made of two metals and is patterned.

[0014] Preferably, the electrode layer of the biomarker sensing platform is made of gold and chromium (Au / Cr).

[0015] In particular embodiments, the electrode layer of the biomarker sensing platform is supported by a flexible, bendable, and durable substrate.

[0016] Preferably, the support substrate of the biomarker sensing platform is made of the same material as the material of the substrate forming the FPCB.

[0017] Preferably, both the support substrate of the biomarker sensing platform and the FPCB substrate are made of polyimide (PI).

[0018] In particular embodiments, each microchamber has a hydrophilic inner surface to increase the flow rate of the sweat obtained from the human skin contact surface through the plurality of microchannels to the corresponding biosensor.

[0019] Preferably, the inner surface of the microchamber is deposited with polyvinyl alcohol and then subjected to plasma treatment.

[0020] In particular embodiments, the biosensors are formed on the patterned electrode layer by initially electrodepositing platinum black, polyaniline (PANI) and poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) (PEDOT:PSS) respectively, thereby forming different films or coatings on the patterned electrodes, and then subjecting the different patterned electrodes to respective mixtures to form ion-selective membranes for sensing respective ions or sensing electrodes for sensing glucose.

[0021] Preferably, the mixture for forming the sensing electrodes for sensing glucose comprises glucose oxidase (GOx), chitosan and carbon nanotubes, and the mixture is deposited on one of the electrodes having a platinum black layer.

[0022] Preferably, the ion-selective membrane for sensing NH 4+ ions is formed on the respective electrode having a PEDOT:PSS coating by an ion-selective solution comprising bis(2-ethylhexyl) sebacate (DOS), non-actin, sodium tetra[3,5-bis(trifluoromethyl)phenyl]borate (Na-TFPB), polyvinyl chloride (PVC) dissolved in tetrahydrofuran.

[0023] Preferably, the ion-selective membrane for sensing Na + ions is formed on the respective electrode having a PEDOT:PSS coating by an ion-selective solution comprising bis(2-ethylhexyl) sebacate (DOS), Na ionophore X, sodium tetra[3,5-bis(trifluoromethyl)phenyl]borate (Na-TFPB), polyvinyl chloride (PVC) dissolved in tetrahydrofuran.

[0024] Preferably, the electrode having a PANI film is used as a pH sensor.

[0025] In particular embodiments, the microfluidics of the biomarker sensing platform are made of a thermoset polymer and are fabricated based on photolithography.

[0026] Preferably, the thermoset polymer of the microfluidics for forming the biomarker sensing platform is polydimethylsiloxane (PDMS).

[0027] In particular embodiments, the replication mold is provided by a process of patterning photoresist on a silicon wafer, then exposing to UV light under a photomask and removing unmasked photoresist to form a patterned mold base, depositing an anti-adhesion layer thereon, followed by injecting a mixture of PDMS monomer and curing agent into the cavities of the patterned mold base to form the microchambers, microfilters and microchannels of the microfluidics of the biomarker sensing platform.

[0028] In particular embodiments, the plurality of microchannels has an average channel width of about 15 pm.

[0029] In particular embodiments, the plurality of microchambers is five cylindrical chambers having an average diameter of about 2 mm and connected to each other by microchannels.

[0030] In particular embodiments, the microfilter is a microcolumn array, each microcolumn has a height of about 20 pm, a side length of about 40 pm, a center-to-center spacing of about 90 pm between two adjacent microcolumns, and a gap distance of about 18 pm between each microcolumn and the sidewall of its adjacent microchannel.

[0031] Preferably, the microcolumn array contains at least ten microcolumns and is disposed at the inlet of the microfluidics.

[0032] The number and / or their dimensions of the microcolumns, microchannels and microchambers can vary as needed.

[0033] In particular embodiments, the PDMS-based microfluidics and the PI-based support base of the biomarker sensing platform are bonded using oxygen plasma treatment.

[0034] In particular embodiments, the mechanical actuator includes two copper coils, each attached with a polyethylene terephthalate (PET) film, and the two PET films seal the magnets surrounded by an epoxy ring, thereby forming a double-sided electromagnetic coil structure.

[0035] In particular embodiments, the plurality of feedback signal outputs is a plurality of light-emitting diodes (LEDs), each LED producing two opposite light signals representing two different states of each respective biological, chemical or physiological signal that the corresponding biosensor is capable of detecting.

[0036] In particular embodiments, the electronic and / or mechanical components further include a temperature sensor, an impedance sensor, an amplifier, a crystal oscillator, a resistor, a capacitor and an internal battery.

[0037] In particular embodiments, the integrated sweat sensing system further comprises at least two elastic strips on two opposite sides of the sensing layer integrated or detachably connected adjacent to the biomarker sensing platform, each of the elastic strips being fixed on each of the two opposite sides of the sensing layer for fastening the system to the contact surface of the human skin.

[0038] In particular embodiments, the encapsulation layer is configured to enclose at least the top surface of the sensing layer and the electronic and / or mechanical components disposed thereon.

[0039] Preferably, the encapsulation layer is made of a thermoset polymer including PDMS.

[0040] In particular embodiments, the adhesive layer is a replaceable adhesive layer consisting of one or more conductive double-sided adhesive strips, wherein one side of the double-sided adhesive strips is attached to the contact surface of the human skin and the other side is adhered to the bottom surface of the sensing layer.

[0041] The second aspect of the present application provides a real-time non-invasive method for monitoring physiological changes of an individual in a plurality of states, comprising:

[0042] fastening the integrated sweat sensing system described in the first aspect or according to various embodiments of the present application on one or more skin areas;

[0043] activating the power supply of the integrated sweat sensing system;

[0044] calibrating each biosensor of the biomarker sensing platform with reference readings of biological, chemical or physiological signals so as to preset a threshold or range of the biological, chemical or physiological signals;

[0045] determining a first state of the individual according to the preset threshold or range through one of the two opposite signals as a first feedback signal given by one of the feedback signal outputs;

[0046] if the corresponding biological, chemical or physiological signal detectable by the corresponding biosensor exceeds the preset threshold or range, generating a second feedback signal according to the other of the two opposite signals given by the feedback signal output and received by the individual, thereby confirming a second state of the individual; and

[0047] when the corresponding detectable biological, chemical or physiological signal exceeds the preset threshold or range, generating a synchronous vibration signal by at least one mechanical actuator together with the generation of the second feedback signal.

[0048] In particular embodiments, the integrated sweat sensing system is fastened to one or more skin areas of the individual by at least two strips attached to the sensing layer and the adhesive layer of the present system.

[0049] In particular embodiments, the biological, chemical, or physiological signals that the respective biosensors are capable of detecting include ions, glucose, pH, temperature, and humidity.

[0050] Preferably, the ions that the respective biosensors are capable of detecting include ammonium cations (NH 4+ ) and sodium (Na + ) ions.

[0051] In particular embodiments, the integrated sweat sensing system is configured to have a microfluidic-based sweat collection module that is an extension from the circuit area of a flexible printed circuit board (FPCB) or a control panel, which is coupled with the array of biosensors for rapid delivery of the sweat collected on the skin surface of the individual through the microfluidic structure to the individual microchambers of the sweat collection module, wherein the microchambers are configured to match the placement of the biosensors on the patterned electrodes of the biomarker sensing platform to sense the changes in the concentration of pH, glucose, different ions in the collected sweat over a period of time.

[0052] In particular embodiments, the FPCB is further incorporated with some sensors for sensing other physiological changes of the individual, including temperature and skin impedance over a period of time.

[0053] Exemplarily, the changes in the concentration of pH, glucose, ions of the collected sweat can be detected in real-time through the electrical response in terms of the voltage changes of the respective sensing electrodes of the biosensors. If the voltage changes of the respective sensing electrodes exceed the threshold or range of the biological, chemical, or physiological signals, the feedback signal output can generate a feedback signal.

[0054] Preferably, the humidity of the one or more regions of the skin can be detected by the impedance sensors incorporated on the FPCB according to the inverse relationship between impedance and skin humidity.

[0055] In particular embodiments, the feedback signal can be output as a visual, audio, or audiovisual signal that represents the state change of the individual in terms of the compositional changes of any of the sweat components.

[0056] Preferably, each pair of opposite signals is generated by a pair of light emitting diodes with opposite color emissions that represent two different states of the biological, chemical, or physiological signals that the respective biosensors are capable of detecting.

[0057] In particular embodiments, the one or more regions of the skin include the skin of the upper arm, and the threshold ranges of the detectable NH4 + , Na + , pH, glucose, skin impedance, and temperature from sweat are 0.1-1 mM, 10-100 mM, 3-8, 10-200 pM, 0.07-0.91 MΩ, and 14-30 °C, respectively.

[0058] In particular embodiments, the threshold range of skin impedance for the skin area of the leg and chest of the human body is 0.34-1.68 MΩ and 0.08-2.63 MΩ, respectively.

[0059] Other aspects of the present application include methods of manufacturing the present integrated sweat sensing system. Details of each of the manufacturing methods with respect to the different parts of the present integrated sweat sensing system and variations thereof can be found in the various embodiments and examples described herein, or can be found in any method known to one of ordinary skill in the art.

[0060] This summary is intended to introduce some concepts in a simplified form, which are further described below in the detailed description. This summary is not intended to identify key or essential features of the claimed subject matter, nor is it intended to be used to determine the scope of the claimed subject matter. Other aspects of the present application are disclosed in the examples below. BRIEF DESCRIPTION OF DRAWINGS

[0061] The drawings contain figures of particular embodiments to further illustrate and describe the above and other aspects, advantages, and features of the present application, in which like numbers denote similar elements throughout the several views. It is to be understood that the drawings depict embodiments of the application and are not, therefore, to be considered limitations of its scope. The application will be described and explained with additional specificity and detail through the use of the accompanying drawings in which:

[0062] Figures 1A-1E The structure and appearance of the integrated sweat sensing system according to particular embodiments of the present application are shown, wherein:

[0063] Figure 1A The structure of the integrated sweat sensing system is schematically depicted through an exploded perspective view;

[0064] Figure 1B A schematic of the sensing layer of the integrated sweat sensing system is shown through one overhead perspective view and an optical image of the sensing layer is shown through another overhead perspective view, wherein the scale bar represents 1 cm and the dashed box indicates the biomarker sensing platform of the sensing layer;

[0065] Figure 1C The structure of the biomarker sensing platform of the sensing layer as indicated by the dashed box in Figure 1B is schematically depicted and an image of a prototype of the detachable and flexible biomarker sensing platform is shown;

[0066] Figure 1D A series of images are shown for a prototype of the integrated sweat sensing system (top left image) and its appearance during the implementation of a study on the skin of multiple sites of a subject, including the arm (top right image), the leg (bottom left image), and the abdominal region (or abdomen) (bottom right image);

[0067] Figure 1E An exploded view showing an embodiment of a mechanical actuator of the integrated sweat sensing system;

[0068] Figures 2A-2L Characteristics of different biosensors and mechanical actuators in the integrated sweat sensing system according to a particular embodiment of the present application, wherein:

[0069] Figure 2A NH4 + Electrical response of the biosensor;

[0070] Figure 2B Na + Electrical response of the biosensor;

[0071] Figure 2C Electrical response of the glucose biosensor;

[0072] Figure 2D Electrical response of the pH sensor;

[0073] Figure 2E Electrical response of the humidity sensor as a function of skin humidity for three different skin regions (arm, leg, and chest);

[0074] Figure 2F Comparison of the skin temperature of a human subject measured by both the temperature sensor in the present system and a commercial temperature measuring device (temperature gun);

[0075] Figure 2G Operating performance of the mechanical actuator (2.8 mm in thickness) as a function of increasing vibration frequency (left) and normalized pressure generated by the actuator under different types of current input (including sine wave, square wave, and pulse) at a constant frequency and current amplitude of 10 Hz and 63.75 mA (right);

[0076] Figure 2H Cycling performance of the mechanical actuator over 134400 test cycles at a power input of 63.75 mA;

[0077] Figure 2I is Figure 2H Enlarged view of the results in over a certain time interval;

[0078] Figure 2J Temperature peak of the mechanical actuator as a function of operating time at a constant square wave current of 63.75 mA and a duty cycle of 1:1;

[0079] Figure 2KA thermal profile image showing the continuous running of the actuator embedded in the control panel for 80 s, with the maximum temperature of 43 °C;

[0080] Figure 2L A theoretical analysis of the actuator is shown when a distributed pressure is applied on the top surface. The results show that the actuator can function normally at a pressure of 7.52 MPa without any damage.

[0081] Figures 3A-3L Microfluidic performance and anti-interference ability of different biosensors according to specific embodiments of the present system are shown, wherein:

[0082] Figure 3A Images of the microfluidic of the biomarker sensing platform are shown, wherein the inset is a magnified view, and the scale bar of the lower magnification image is 3 mm;

[0083] Figure 3B Images of three different sites with three different bending radii (1 cm, 1.7 cm, and 2.5 cm) are shown, on which the microfluidic of the biomarker sensing platform as shown in Figure 3A is mounted, wherein the scale bar is 1 cm;

[0084] Figure 3C Images of the microfluidic of the biomarker sensing platform mounted on three different sites as shown in Figure 3B are shown, wherein the skin debris indicated by the arrows is filtered under three different bending conditions (scale bar = 0.1 mm);

[0085] Figure 3D The filtration rate of the microfluidic filtering out standard microspheres under three different bending conditions as shown in Figure 3B is shown;

[0086] Figure 3E The occupancy of sweat filling in the microfluidic as a function of running time is shown;

[0087] Figure 3F A theoretical model showing the sweat flow trajectory generated in the chamber of the simulated microfluidic is shown;

[0088] Figure 3G The anti-interference ability of the NH4 + sensor under the interference of NaCl, NH4Cl, and KCl (1 mM each) at different times during the test is shown;

[0089] Figure 3H The anti-interference ability of the Na + sensor under the interference of NaCl, NH4Cl, and KCl (20 mM each) at different times during the test is shown;

[0090] Figure 3I The anti-interference capability of the glucose sensor under the interference of NaCI, glucose and KCI (20 μΜ each) at different times during the test is shown;

[0091] Figure 3J The anti-interference capability of the pH sensor under the interference of NaCI, KCI and formic acid (30 μΜ each) at different times during the test is shown;

[0092] Figure 3K The comparison of the electrical response of the humidity (skin impedance) sensor before and after 1000 bending cycles at a constant bending angle of 60° and a frequency of 1 Hz is shown;

[0093] Figure 3L The comparison of the electrical response of the temperature sensor before and after 1000 bending cycles at a constant bending angle of 60° and a frequency of 1 Hz is shown.

[0094] Figures 4A-4H An implementation example of the system and the performance of different sensors on human skin at the same site installed on three different test subjects are shown, wherein:

[0095] Figure 4A An image of one of the test subjects wearing the system on his right upper arm to monitor the sweat composition concentration during running is shown;

[0096] Figure 4B The electrical response of the six biosensors (NH4 + , Na + , glucose, pH, humidity and temperature) of the test subject shown in Figure 4A during a 20-minute running exercise is shown;

[0097] Figure 4C The change in the readings of the NH4 + sensor installed on three different sites (arm, chest, leg) of each of the three test subjects is shown;

[0098] Figure 4D The change in the readings of the Na + sensor installed on three different sites (arm, chest, leg) of each of the three test subjects is shown;

[0099] Figure 4E The change in the readings of the glucose sensor installed on three different sites (arm, chest, leg) of each of the three test subjects is shown;

[0100] Figure 4FThe changes in pH sensor readings at three different sites (arm, chest, and leg) installed on each of the three test subjects are shown.

[0101] Figure 4G The changes in readings of humidity (impedance) sensors installed at three different locations (arm, chest, and leg) on ​​each of the three test subjects are shown.

[0102] Figure 4H The changes in temperature sensor readings at three different locations (arm, chest, and leg) on ​​each of the three test subjects are shown.

[0103] Those skilled in the art will understand that the elements in the accompanying drawings are shown for simplicity and clarity and are not necessarily depicted to scale. Detailed Implementation

[0104] It will be apparent to those skilled in the art that modifications, including additions and / or substitutions, can be made without departing from the scope and spirit of the invention. Specific details have been omitted to avoid obscuring the invention; however, this disclosure is written to enable those skilled in the art to practice the teachings herein without undue experimentation.

[0105] Turning Figures 1A-1E It provides an integrated sweat sensing system 10. Figure 1A The basic structure of this integrated sweat sensing system 10 according to a specific embodiment is schematically depicted, comprising: a thin encapsulation layer 15, which may be made of an elastic biocompatible polymer such as polydimethylsiloxane (PDMS) to resist potential interference from various external stimuli; a sensing layer 12, which includes a flexible printed circuit board (FPCB) on which a series of electronic and mechanical components and mechanical actuators are soldered, serving as a control panel for data acquisition, processing, and feedback; and a detachable and flexible chemical sensing platform 13 with matched microfluidics for simultaneously detecting NH4 in fresh sweat via a multi-sensor array. + Na + pH value and glucose concentration ( Figure 1C The three images show the same prototype; and a thin, replaceable adhesive layer 11, such as a double-sided conductive strip, to enhance the bond between human skin 20 and the system 10 for subsequent continuous monitoring of skin impedance. To reduce the impact of human movement on the multi-biosensor signals acquired by the system, four fabric strips 14 are used to secure the device to the user's target skin area. The number of strips can be varied as required. To provide simultaneous vibration-based feedback, a low-power (<0.2W) miniaturized vibration tactile actuator based on a bilateral electromagnetic coil design is used. Figure 1E), the vibration strength of which can be widely enhanced by a coupled electromagnetic field induced by two copper (Cu) coils 1211. The magnet 1213 is inserted into the hollow space of the epoxy resin ring 1214, and then sealed by two PET films 1212. The two Cu coils 1211 are combined on the two PET films 1212 so as to sandwich the sealed magnet 1213 in the ring 1214, thereby forming an electromagnetic coil structure.

[0106] Figure 1B The device layout is shown with all electronic components soldered on the FPCB, including the mechanical actuator 121. In a particular embodiment, the mechanical actuator has a diameter of 11 mm and a thickness of 2.8 mm. The electronic components include a microcontroller (MCU), six LEDs with two alternative light colors (e.g., green and red), an amplifier, a crystal oscillator, resistors, capacitors, and an 80 mAh lithium-ion battery. In a particular embodiment, the system is configured to remain in a silent state when all detected physiological signals are within the normal range, in which all LEDs emit green light and the mechanical actuator is at rest, in which the NH4 + , Na + , pH, glucose, skin impedance, and temperature of the human arm are in the normal ranges of 0.1-1 mM, 10-100 mM, 3-8, 10-200 mM, 0.07-0.91 MΩ, and 14-30 °C, respectively. These physiological signal ranges can vary from one measurement site to another. When the signal of any one biomarker is outside the normal range at a particular measurement site, the corresponding LED will turn red within 10 seconds, while the mechanical actuator will sequentially cause vibrations, while alerting the user to take necessary measures, such as slowing down the exercise intensity, drinking water rich in electrolytes, or resting, etc.

[0107] Figure 1C An embodiment of a flexible and detachable biomarker platform is shown, which has four biosensors (for sensing NH4 + , Na + , pH, and glucose, respectively) integrated into one patch. The biomarker platform 13 includes a base layer 131 as a support platform, e.g., a 150-μm polyimide (PI) layer; an electrode layer 132, e.g., a 240-nm patterned electrode layer made of gold / chromium (Au / Cr, 200 nm / 40 nm); four catalyst patches 133, which serve as the sensing sites for NH4 + , Na +, pH and glucose sensitive biosensors; and microfluidics 134 made of PDMS. By directly interfacing with human skin, the four sensors can accurately analyze the concentration of corresponding sweat components in a short time via high-performance microfluidics with a response time less than 15 s. This advanced device layout and material selection provides great flexibility for the system, which can be bent more than 90° without any structural damage, enabling it to be applied to various parts of the human body, including the arm, leg and abdomen Figure 1D

[0108] Figures 2A-2L The electrical characteristics of the biosensors and mechanical actuators are shown. For NH4 + and Na + ion-selective electrodes, ion-selective membranes with corresponding ion carriers are used as sensing layers, while a layer of poly(3,4-ethylenedioxythiophene) (PEDOT): poly(4-styrenesulfonic acid sodium salt) (PSS) is electrodeposited on the electrodes as an ion-to-electron transducer to minimize the potential drift of the biosensors. As shown in Figure 2A , the response of the NH4 + sensor positively correlates with the increase of NH4 + concentration, and the corresponding sensor exhibits excellent linearity in a wide detection range from 0.1 mM to 100 mM. Similarly, the Na + sensor also shows excellent sensing behavior for Na + in a physiological relevant concentration range of 10-200 mM Figure 2B . Glucose and pH sensors are prepared using a similar potential sensing method with glucose oxidase (GOx) and polyaniline (PANI) as sensing components, respectively. Figure 2C and 2D show the voltage responses of the glucose sensor and the pH sensor to various glucose concentrations (10 to 200 mM) and pH levels (3 to 8), respectively. Both sensors exhibit good linearity in the physiological relevant concentration range in sweat. Figure 2E The electrical response of the skin impedance sensor measuring the change in skin impedance with skin humidity is shown, which is also verified by a commercial skin humidity sensor at three human body locations, including the arm, leg and chest. It is found that the skin impedance detected by the sensor decreases linearly with the increase of skin humidity, 0.91-0.07 MΩ@25%-83% for the arm, 1.68-0.34 MΩ@30%-90% for the leg, and 2.63-0.08 MΩ@27%-67% for the chest. In addition to the skin impedance sensor, the skin temperature is measured by a thermistor embedded in the electrical control panel. The thermistor-based temperature sensor can be calibrated by a commercial temperature sensor, where the temperature range is 24-32 °C, covering most of the application prospects of users at room temperature​Figure 2F ). The results show that the electrical performance of the sensor is comparable when the slope coefficient of the fitted curve is about 0.9.

[0109] Figures 2G-2I The electrical characteristics of the mechanical actuator according to a particular embodiment of the wearable form of the present system are shown. To achieve low energy consumption characteristics, it is necessary to operate the mechanical actuator at the resonance frequency. Figure 2G The normalized amplitude of the actuator is shown as a function of the operating frequency from 10 Hz to 150 Hz at intervals of 30 Hz, where it can be clearly seen that at a constant square wave current of 63.75 mA, at a maximum yield pressure of 0.43 kPa, the resonance peak is at 70 Hz (left graph); the right graph shows the normalized pressure of the actuator as a function of the current input type (including sine wave, square wave and pulse) at a constant frequency and current amplitude of 63.75 mA, where the square wave current can generate much higher pressure than the other two types, since at the same peak amplitude value, the power input of the square wave is higher. To verify the stability of the actuator, a long duration test of 32 minutes of continuous operation at the resonance frequency (70 Hz) was carried out >10 5 cycles of operation, where the actuator remained very robust without any fatigue or damage Figure 2H ). Figure 2I The detailed electrical signal of the actuator during long-term operation is shown, further demonstrating its high stability.

[0110] Since the actuator generates a large amount of heat during continuous operation, the temperature peak of the actuator embedded in the control panel was monitored for more than 80 s Figure 2J ) in full load operation mode, the results show that the maximum temperature is 43 °C, which is further verified by the thermal distribution image of the actuator in Figure 2K . The maximum temperature of the actuator (43 °C) does not cause any discomfort to the user, which further demonstrates its practicality. When a distributed force with a pressure of 7.52 MPa is applied on the top surface of the actuator, the PET layer begins to suffer plastic deformation, with a maximum pressure of 40 MPa Figure 2L ).

[0111] Figures 3A-3F The characteristics of the microfluidic are shown. As Figure 3AAs shown, the microfluidic according to certain embodiments comprises two parts: a micro-filter for sweat filtration (small inset), and a micro-chamber for biomarker signal monitoring. In certain embodiments, the micro-filter is integrated with a plurality of micro-pillars (e.g., 10 micro-pillars), where each micro-pillar can have the following dimensions: height: 20 pm; side length: 40 pm; center-to-center spacing: 90 pm; gap distance between micro-pillar and sidewall of micro-channel: 18 pm, disposed at the inlet to constitute a sweat filtration module. The micro-chamber for biomarker signal monitoring comprises 5 cylindrical chambers (each with a diameter of about 2 mm) and are connected to each other by micro-channels (channel width: 15 pm). To investigate the effect of the micro-filter in the micro-chamber on reducing biomarker sample contamination during sweat sampling, sweat samples from volunteers were processed through the micro-filter. The microfluidic-based sweat filtration module can be applied to filter skin debris and microspheres of ~20 pm in diameter when applied to cylinders with different curvatures Figures 3B-3C ). Microspheres of 20 pm in diameter mixed in distilled water were also used to verify the filtration function of the micro-filter. As shown in Figure 3D , even in the case of bending, nearly 95% of the microspheres can be filtered, which verifies that the microfluidic of the present application can provide stable skin debris filtration performance when applied to different types of human interfaces (e.g., forehead, elbow joint, etc.). Since the sweat collection rate plays a key role in biomarker signal monitoring, the micro-chamber was provided with a hydrophilic surface by depositing polyvinyl alcohol (PVA) followed by plasma treatment to improve the collection rate. The occupancy time of the sweat achieved in the five micro-chambers was calculated, where the sweat can cover up to 80% of the volume of each micro-chamber within 15 s and about 100% of the volume within 1 min Figure 3E . Streamline simulation of the flow of sweat into the micro-chamber shows that the sweat collected from the human body can be evenly distributed to the micro-chamber, which indicates that the microfluidic-based sweat collection module can provide efficient sweat collection rate and reduce the interference of skin debris Figure 3F .

[0112] Figures 3G-3J The anti-interference ability of the biosensors according to certain embodiments of the present application on NH 4+ , Na + , pH value and glucose, respectively, is shown. It can be observed that when interference is generated on the electrolyte by adding other electrolytes, the fluctuation of the voltage output of the sensor is negligible, while there is a clear response by adding the corresponding electrolyte matching the sensor at the same concentration as the other electrolytes, which indicates that these biosensors have excellent anti-interference ability. The stable performance is mainly determined by the specificity of the corresponding sensing components of the enzyme or ion carrier.

[0113] In terms of humidity and temperature sensing, Figure 3K and3L The electrical responses of the skin impedance sensor and the temperature sensor integrated in the control panel (circuit area) of the FPCB are shown at a constant bending angle and frequency of 60° and 1 Hz, respectively, after 1000 bending cycles. The results show that these two biosensors for skin impedance and temperature can still accurately detect the corresponding physiological signals, with a signal range of 0.819-0.848 MΩ @ the current value of the skin impedance sensor is 0.835 MΩ, and 28.1-28.3 °C @ the current value of the temperature sensor is 28.2 °C, which proves their high resistance to external stimuli.

[0114] Figures 4A-4H An embodiment of the present system on a human subject running is shown, in which the system is mounted on the arm of the corresponding volunteer for recording the physiological information during his / her jogging, including NH4 + , Na + , glucose, pH, skin impedance, and skin surface temperature. As shown in Figure 4A , in order to better present the running condition of the electronic device, a thin transparent PDMS is used to encapsulate the other parts of the present system, especially the FPCB with electronic components soldered. In order to read the real-time biomarker signals, the wired connection between the electronic device and the personal computer has Figure 4B six physiological signals as shown. In some other embodiments, the present system can incorporate a wireless communication module and utilize the corresponding protocol to transmit the recorded signal data to the paired device or system outside the present system and receive any feedback or instructions from it. The wireless communication module and the corresponding protocol can be any wireless communication module and protocol known to those skilled in the art for achieving the same purpose. During the 20-minute continuous operation, the NH4 + concentration, Na + concentration, and arm skin temperature increase with the increase of the amount of sweat produced (0.09-0.97 mM @ NH4 + , 32.6-55.9 mM @ Na + , 21.7-22.2 °C @ temperature sensor), which is due to the high NH4 + and Na +The concentration is caused by the large amount of heat generated internally during exercise. In the first 200 seconds, the glucose concentration in the target skin area stabilizes at around 66.7 μM, then begins to decrease to 13.4 μM at 1200 seconds, due to the significant increase in sweat at 200 seconds, which dilutes the concentration. Unlike the previous four biomarkers, pH and skin impedance decrease with increasing sweat volume throughout the process, ranging from 6.65 to 5.9 @ pH and 2.25 to 0.89 MΩ @ impedance sensor. It is particularly clear that in the first 200 seconds, skin impedance rapidly decreases from 2.25 MΩ to 1.38 MΩ, due to the significant reduction in skin impedance caused by the small amount of water in fresh sweat.

[0115] Figure 4C -H represents the NH4 content of sweat measured after 20 minutes of running in place, compared to initial values ​​at the arms of the three volunteers. + Na + The results showed that NH4+ levels at 20 minutes were related to glucose, pH, skin impedance, and skin surface temperature. + Na + And skin temperature levels higher than at 0 minutes, which is related to Figure 4B The results shown are consistent with previous studies. Figure 4F As shown, the pH levels of the three volunteers exhibited slight changes at the start and 20 minutes. In addition to the four biosensors, glucose concentration and skin impedance (…) Figure 4D and 4G All showed a negative correlation with exercise time, which is consistent with Figure 4B The results are consistent with those obtained in previous studies. These results demonstrate the potential of this system to monitor user physiological signals by analyzing sweat component concentrations using a synchronous feedback signal mechanism, without any instructions or feedback mechanisms transmitted from external devices such as smartphones and smartwatches. Thanks to its high reliability, stability, long uptime, and detection of multiple biochemical signals, this system holds great promise as an integrated smart skin electronics device for healthcare monitoring.

[0116] The following non-limiting examples are intended to help understand specific embodiments of the invention. The scope of the invention should be determined with reference to the appended claims.

[0117] Example 1 - Manufacturing of a Biomarker Sensing Platform

[0118] Fabrication of the biomarker sensing platform as a sensor patch started with PI (150 pm) adhered on glass. First, the thin film was cleaned with deionized water, ethanol and acetone accordingly. Then, a layer of Cr / Au (10 nm / 100 nm) was coated on the cleaned PI thin film by e-beam. To obtain the specific pattern of the sensor patch, a layer of photoresist (AZ 4620, AZ Electronic Materials) was spin-coated at 3000 rpm for 30 s, baked on a hot plate at 110 °C for 5 min, exposed to UV light for 45 s, and developed in the solution (AZ 400K) for 1 min accordingly. Finally, the excess Au and Cr were etched off, followed by cleaning with acetone to remove the residual photoresist, thus obtaining the patterned electrode layer. To prepare the glucose, pH and various ion biosensors, platinum black, PANI and PEDOT:PSS were electrodeposited onto the corresponding locations of the patterned electrode layer, respectively. Specifically, platinum black was electrodeposited in an electrolyte consisting of 24 mM chloroplatinic acid and 2.1 mM lead acetate at a constant voltage of -0.8 V for 150 s; PANI was electrodeposited in a mixed solution of 0.1 M aniline and 1 M H2SO4 by cyclic voltammetry from -0.2 V to 1 V for 20 cycles; PEDOT:PSS was obtained by constant current electrodeposition (1 mA / cm2) for 15 min in a solution containing 0.01 M 3,4-ethylenedioxythiophene (EDOT) and 0.1 M polystyrene sulfonate (NaPSS). More specifically, the electrodeposition process was carried out in a three-electrode system: the Au electrode as the working electrode; platinum wire and Ag / AgCl as the counter and reference electrodes, respectively. A 2 pL mixed solution of glucose oxidase (GOx) (2.5 U / pL), chitosan (1 mg / mL) and carbon nanotubes (2 mg / mL) was dropped on the platinum black layer as the sensing electrode for the glucose biosensor; 2 pL of ion-selective solutions of NH4+and Na+were dropped on the corresponding electrodes to form the ion-selective membranes, respectively. Specifically, the solution of NH4+was obtained by dissolving a 200 mg mixture of non-actin (1% w / w), sodium tetra[3,5-bis(trifluoromethyl)phenyl]borate (Na-TFPB, 0.55% w / w), polyvinyl chloride (PVC, K value 72 - 1, 33% w / w) and bis(2-ethylhexyl)sebacate (DOS, 65.45% w / w) in 1.2 mL tetrahydrofuran; the Na+ion-selective solution was formed by replacing non-actin with Na+ionophore X. Finally, the common reference electrode was realized by printed Ag / AgCl ink. 2 4+ + 4+ +

[0119] Example 2 - Fabrication of microfluidics for sweat collection ​​​​​

[0120] Fabrication of the microfluidic-based sweat collection module was based on photolithography and PDMS replica molding. Initially, the mold was fabricated by patterning SU-8 photoresist (SU-8 2015, Microchem) with a thickness of 20 pm on a 4-inch silicon wafer, followed by isopropanol, acetone, and deionized water cleaning, and finally isopropanol rinsing. After soft baking at 95 °C for 3 min on a hot plate, the coated photoresist was exposed to UV light by mounting to a photomask, followed by another 3 min baking at 95 °C. The exposed substrate was immersed in a developer solution (SU-8 Developer, Microchem) for 5 min to remove any unexposed photoresist. Next, the mold was salted by depositing a molecular layer of trichloro(lH, lH, 2H, 2H-perfluorooctyl)silane (Sigma-Aldrich, St. Louis, MO, USA) to facilitate the release of PDMS from the mold master. After the completion of mold fabrication, the PDMS monomer was mixed with the curing agent at a weight ratio of 15: 1; and the mixture was degassed for 5 min. The degassed PDMS mixture was poured into the control layer mold with a thickness of 250 pm. After baking at 70 °C for 30 min, the PDMS substrate was cut and peeled off from the wafer. The inlet was obtained by punching a hole with a diameter of 1.5 mm, and the outlet was formed by cutting an extra PDMS at the end of the microchannel. Finally, the PDMS substrate was bonded to the PI support substrate of the sensing layer using the following steps: 2 min of oxygen plasma treatment (energy: 5 kJ; Harrick Plasma Cleaner PDC002) and 15 min of baking at 70 °C to enhance the adhesion between the PDMS and the PI layer.

[0121] Example 3 - Fabrication of the mechanical actuator

[0122] Reference Figure 1E The structure depicted in FIG. 6 was fabricated. First, the Cu coil was fabricated by 0.5 mm thick copper wire to form an electromagnetic coil with a thickness of 0.25 mm (Yisuo Electronics Co., Ltd, Dongguan, China). The inner diameter was 2 mm and the outer diameter was 11 mm. The support epoxy ring was 3D printed with a constant thickness of 1.9 mm. The diameter of the magnet used was 8 mm. The 0.2 mm thick PET film was cut into a circle with a diameter of 11 mm by a laser cutting machine. First, the Cu coil was fixed on a clean glass plate, and then the PET film was glued to the top of the coil. After the glue was cured at 80 °C for 1 min, the epoxy ring was bonded to the PET film with quick-drying glue. The magnet was placed inside the hollow space of the ring, and then the chamber was sealed with the PET layer. After the top Cu coil was bonded to the PET, the actuator was obtained.

[0123] In terms of mechanical analysis, the FEA commercial software ABAQUS (Analysis User’s Manual 2020) was used to obtain the impact resistance of the actuator. Cu, PET, epoxy, and magnets were modeled by 860,000 hexahedral elements (C3D8R). The minimum element size was 0.002 mm to ensure the convergence and accuracy of the simulation results. A uniform distributed force with a pressure of 5.53 MPa was applied on the top surface of the actuator, and the magnets were set as rigid bodies. The elastic modulus (E) and Poisson’s ratio (v) used in the analysis were E = 131 GPa, v = 0.33 for Cu, E = 3.5 GPa, v = 0.35 for PET, E = 3.35 GPa, v = 0.35 for epoxy, and E = 1.8 GPa, v = 0.3 for the magnets. Cu Cu PET PET Epoxy

[0124] Example 4 - Fabrication of the integrated sweat sensing system

[0125] The fabrication of the integrated sweat sensing system started with the circuit, which was fabricated on a PI substrate with gold-plated (thickness ~ 50 nm) copper (thickness ~ 10 pm) by flexible printed circuit board technology. An insulating layer was covered on the exposed circuit to prevent short circuits. All the electrical components, including the microcontroller (ATmega328p-mu), capacitors (14-22 pF), resistors, crystal oscillator (16 MHz), connectors, LEDs, and mechanical actuators, were soldered to the flexible printed circuit board (FPCB) and to the corresponding contact pads on the Cu / PI substrate. Two strips were integrated in the control panel of the FPCB to ensure close contact between the human skin and the sensing layer of the system. To enhance the resistance to the tearing strength caused by the strips, a customized hard printed board was fixed on both opposite sides of the FPCB by quick-drying glue, and then encapsulated with PDMS (PDMS: crosslinking = 10: 1) and cured at 80 °C for 10 minutes. Finally, two conductive double-sided tapes (3J ordinary double-sided conductive tape, Dongguan Xinxin Packaging Material Co., Ltd.) were pasted on the exposed Cu patches on the bottom surface of the control panel, each with a constant size of 1 cm x 1 cm.

[0126] Example 5 - Operation of the integrated sweat sensing system

[0127] Since the current integrated sweat sensing system mainly includes two sensing regions, i.e., the sweat sensor array (for sensing ions, such as NH4 + , Na + ​​​​​; glucose; and pH) and a control panel (circuit area) on the FPCB for processing data and measuring skin impedance and temperature, in this example, an application programmable MCU (ATMEGA328P-MU, Microchip Technology Inc.) is used, while the entire system is powered by an external 5V battery. During operation, the analog signals obtained by the sensor array are converted into digital signals by using the 10-bit ADC function of the MCU. Furthermore, since the amplitude of the analog signals generated by the sweat sensors is small, in the millivolt range, the signals are passed through an instrumentation amplifier (INA321EA / 250, Texas Instruments) before the ADC step to enable accurate measurements. Moreover, to measure positive and negative signals from the sensors, 3.3V, which can be directly reduced from 5V by an LDO (TPS76933DBVR, Texas Instruments), is applied to the raw signals, and the return values are calculated internally in the MCU. To sense skin impedance and temperature, two independent copper pads that directly contact the individual's exposed skin and a negative temperature coefficient (NTC) thermistor mounted on the control panel are employed. Similar to the sweat sensor array, the signals from the skin sensors are converted into digital signals by the ADC pins of the MCU and processed in the MCU. Finally, to produce mechanical and visual alert functions, the actuators are activated by the forward and reverse current direction functions of an H-bridge (BD6211F-E2, Rohm Semiconductor) and six bi-color LEDs (green and red) are employed to indicate the status of each parameter in real time, while they are controlled in real time by the digital pins of the MCU.

[0128] Although the present application has been described in terms of certain embodiments, other embodiments apparent to those of ordinary skill in the art are within the scope of the present application. Therefore, the scope of the present application is intended to be limited only by the claims appended hereto.

[0129] Industrial applicability

[0130] The present application provides a flexible, bendable, and reliable health monitoring system with at least six integrated electrochemical sensors for real-time tracking and analyzing human sweat composition, skin impedance, and surface temperature, whose sensing signals can be associated with an embedded actuator based on mechanical vibration and six LEDs of variable light color as a feedback module. The biosensor exhibits excellent sensing performance and outstanding mechanical performance in terms of sensitivity, anti-interference ability, and reliability. The microfluidic module of the present application, which is configured with a hydrophilic surface for sampling, can greatly improve sweat collection efficiency and avoid sweat contamination. In addition to the sensor, the wearable mechanical actuator also has a high pressure (0.43 kPa) generated at a square wave current of 63.75 mA and also exhibits high stability, thereby being able to operate under a distributed stress of more than 7.52 MPa. The demonstration of real-time recording of six human physiological signals during exercise also demonstrates the potential of the present application in various practical applications such as entertainment, health care, and medicine.

Claims

1. An integrated sweat sensing system comprising a multi-layer structure, the multi-layer structure comprising: an adhesive layer configured to enhance adhesion of a contact surface of the integrated sweat sensing system to a human skin of a user; a sensing layer; and an encapsulation layer configured to encapsulate the sensing layer, wherein the sensing layer comprises: a flexible printed circuit board (FPCB) having a plurality of electronic and mechanical components disposed thereon, the plurality of electronic and mechanical components comprising: a mechanical actuator configured to output a vibration signal to the user; a microcontroller (MCU); and a plurality of light emitting diodes (LEDs) configured to emit a visual signal to indicate whether a detected physiological signal is within a normal range, a biomarker sensing platform comprising: an electrode layer; a microfluidic comprising a micro-filter configured to filter sweat and a plurality of micro-chambers configured to monitor biomarker signals; and wherein the micro-filter is configured to serve as an interface between the microfluidic and the contact surface of the human skin, and a plurality of micro-channels connecting the respective micro-chambers together and the micro-chambers to the micro-filter, a plurality of biosensors configured to sense ammonium cations (NH4 + ), sodium (Na + ) ions, pH, and glucose, wherein the mechanical actuator is a double-sided electromagnetic coil structure comprising two copper coils, each copper coil having a polyethylene terephthalate film attached to seal a magnet surrounded by an epoxy ring. the biomarker sensing platform is configured to be flexible and detachably mounted on the contact surface of the human skin of the user.

2. The system of claim 1, wherein, the biomarker sensing platform is integrated to the FPCB forming a continuous layer with a circuit area of the FPCB on which the plurality of electronic and mechanical components are soldered; or the biomarker sensing platform is detachably connected to a circuit area of the FPCB on which the electronic and mechanical components are disposed.

3. The system of claim 1, wherein, the electrode layer of the biomarker sensing platform is made of two metals including gold and chromium and is patterned.

4. The system of claim 1, wherein, the electrode layer of the biomarker sensing platform is supported by a substrate made of polyimide (PI).

5. The system of claim 1, wherein, each micro-chamber of the plurality of micro-chambers has a hydrophilic inner surface configured to increase a flow rate of the sweat obtained from the contact surface of the human skin to flow through the plurality of micro-channels to the respective biosensors.

6. The system of claim 1, wherein, the plurality of biosensors are formed on the patterned electrode layer by initially electrodepositing platinum black, polyaniline (PANI) and poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) (PEDOT:PSS) respectively, thereby forming different films on the patterned electrodes, and then subjecting the different patterned electrodes to respective mixtures to form ion-selective membranes for sensing respective ions or a sensing electrode for sensing glucose.

7. The system of claim 4, wherein, the microfluidic of the biomarker sensing platform is made of a thermoset polymer and is manufactured based on photolithography and replication molding.

8. The system of claim 1, wherein, the plurality of micro-channels have an average channel width of about 15 pm.

9. The system of claim 1, wherein, the plurality of micro-chambers are five cylindrical chambers having an average diameter of about 2 mm and are connected to each other by the plurality of micro-channels.

10. The system of claim 1, wherein, the plurality of micro-channels have an average channel width of about 15 pm. the plurality of micro-chambers are five cylindrical chambers having an average diameter of about 2 mm and are connected to each other by the plurality of micro-channels.

11. The system of claim 1, wherein, The micro-filter is disposed at the inlet of the microfluidic and consists of a micro-pillar array, each micro-pillar having a height of about 20 pm, a side length of about 40 pm, a center-to-center spacing of about 90 pm between two adjacent micro-pillars, and a gap distance of about 18 pm between each micro-pillar and the sidewall of its adjacent micro-channel.

12. The system of claim 1, wherein, The electronic and mechanical components further include a temperature sensor, an impedance sensor, an amplifier, a crystal oscillator, a resistor, a capacitor, and an internal battery.

13. The system of claim 1, further comprising at least two elastic strips fixed on two opposite sides of the sensing layer, adjacent to the positions where the biomarker sensing platform is integrated or detachably connected thereto, the at least two elastic strips configured to fasten the integrated sweat sensing system to a contact surface of a user’s human skin.

14. The system of claim 1, wherein, The encapsulation layer is configured to at least enclose the top surface of the sensing layer and the electronic and mechanical components disposed thereon.

15. The system of claim 1, wherein, The adhesive layer is a replaceable adhesive layer consisting of one or more conductive double-sided adhesive strips, wherein one side of the double-sided adhesive strips is configured to be attached to the contact surface of the human skin, and the other side is adhered to the bottom surface of the sensing layer.

16. A real-time non-invasive method for monitoring physiological changes of an individual in multiple states, comprising: fastening the integrated sweat sensing system of claim 1 on one or more skin regions of an individual; activating the power supply of the integrated sweat sensing system; calibrating each biosensor of the biomarker sensing platform with reference readings of biological, chemical, or physiological signals, so as to preset the threshold values of the biological, chemical, or physiological signals; outputting a first feedback signal indicating that the individual is in a normal state when the detected signal is within the preset threshold value; and outputting a second feedback signal when the detected signal exceeds the preset threshold value, and simultaneously generating a vibration signal through a mechanical actuator, thereby indicating an abnormal state.

17. The method of claim 16, wherein, Corresponding biosensors are capable of detecting biological, chemical or physiological signals including ions, glucose, pH, temperature and humidity, and wherein the ions include ammonium cations (NH4 + ) and sodium (Na + ) ions.

18. The method of claim 17, wherein, The changes in ion concentration, glucose, and pH value of the collected sweat can be detected in real time through the electrical response of the voltage change of the corresponding sensing electrode of the biosensor; wherein the humidity of one or more regions of the skin can be detected by the impedance sensor combined on the FPCB according to the inverse relationship between impedance and humidity.

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