+86 400-003-5559 CN


Wireless battery-free wearable sweat sensor powered by human motion | Science Advances

tagscapacitor resistor and inductor

These authors contributed equally to this work.

Wireless wearable sweat biosensors have received great attention due to their potential for non-invasive health monitoring. Since high energy consumption is a key challenge in this field, the effective collection of energy from human movement represents an attractive way to sustainably power future wearable devices. Despite a lot of research activities, most wearable energy harvesters still have the disadvantages of complex manufacturing process, poor robustness and low power density, so they are not suitable for continuous biosensing. Here, we propose a highly durable, mass-produced wearable platform that does not require batteries. This platform can effectively remove the body from human movement through a free-standing frictional electric nanogenerator (FTENG) based on flexible printed circuit boards (FPCB) Get energy. The carefully designed FTENG shows a high power output of approximately 416 mW m

. Through seamless system integration and effective power management, we demonstrated a battery-free friction electric drive system that can provide multiple powers for human sweat biosensors, and wirelessly transmit data to the user interface through Bluetooth during human human testing .

A large number of studies on the development of wearable bioelectronics technology have greatly expanded the vision of personalized health monitoring (


). Wireless wearable devices provide a non-invasive means to extract real-time physiological parameters that indicate health conditions and transmit continuous data to user equipment. Wearable devices capable of detecting various vital signs (such as pulse, respiratory rate, and temperature) have been widely commercialized and integrated into daily life (

). Sweat is another attractive medium, which contains a variety of molecular biomarkers, including electrolytes, metabolites, amino acids, hormones, and drugs that can be analyzed by wearable sensors (

). Continuous monitoring of these biomarkers may supplement laboratory-based blood tests, thereby realizing real-time monitoring of daily health conditions and early disease detection and management (


In the past few years, extensive interest and efforts have focused on developing novel sensors and improving the wearability of these platforms (


). So far, most wearable sensor prototypes have relied on bulky rigid battery packs to power electronic circuits for data collection, processing, and transmission. Some people suggest using flexible batteries to make skin conformal contact (

), combined with low-power electronic devices, which greatly reduces the power requirements of wearable devices and allows the use of small button batteries. Despite these efforts, batteries still face limitations because they need to be charged and replaced frequently. In addition, although unlikely, lithium-ion batteries are prone to explosion, causing safety hazards. Reported on battery-less systems powered by Near Field Communication (NFC) (

), but the operating distance is short. As an alternative, energy can be harvested from renewable, portable and sustainable energy sources (such as solar energy, biological fluids and human movement) to power future wireless wearable electronic devices (

The triboelectric nanogenerator (TENG) converts the mechanical energy generated by human motion into electrical energy through the coupling of induction and triboelectric effect (

), provides an attractive energy harvesting strategy for powering wearable sweat sensors in intensive physical exercise, because their operation is independent of uncontrolled external sources, such as sunlight or wireless power transmitters. Despite the advantages, most existing TENG-based devices still have the problems of low power intensity, low power management efficiency, and insufficient power continuity and life. Therefore, there are no reports of using TENG to continue to power fully integrated wireless wearable molecular sensor systems (

Here, we propose a battery-free, fully self-powered wearable system, which consists of a high-efficiency wearable stand-alone TENG (FTENG), low-power wireless sensor circuit and a microfluidic sweat sensor patch, located in a single A flexible printed circuit board (FPCB) platform that can dynamically monitor key sweat biomarkers (for example, pH and Na

) (

). This wearable sweat sensor system (FWS) driven by FTENG

) The design and manufacture are compatible with the traditional FPCB manufacturing process, which can realize mass production and high reliability. Our FPCB-based independent design combined with effective power management can efficiently collect energy from human skin, and is particularly suitable for powering wearable devices in contact with the skin. With waterproof medical tape, FWS

Can be stacked conformally on the side torso to maximize energy collection (

). The integrated Bluetooth Low Energy (BLE) module can easily transmit sensor data to the mobile interface to track the health status during exercise. This is the first demonstration of a fully integrated battery-free friction electric drive wearable system for hyperhidrosis sensing.


) Schematic diagram illustrating FWS

The product integrates human movement energy collection, signal processing, microfluidic sweat biosensing and Bluetooth-based wireless data transmission into the mobile user interface for real-time health tracking. (


) Optical image of FWS based on FPCB

Can be worn on the side of the human body. Scale bar, 4 cm. (

) Schematic diagram of FPCB-based FTENG with grating slider and fork designator. (

) Schematic diagram of FWS

Shows a microfluidic-based sweat sensor patch connected to a flexible circuit. (

) System-level block diagram showing the power management, signal conversion, processing and wireless transmission of FWS

From FTENG to biosensor to user interface. Image courtesy: Yu Song, California Institute of Technology.

FTENG consists of an interdigitated stator and a sliding block with grating pattern (

). In order to obtain a strong charging effect, polytetrafluoroethylene (PTFE) and copper are used as friction pairs in the flexible FTENG. FTENG is manufactured through commercial FPCB technology (as shown in Figure S1), and the detailed size parameters are shown in Figure 5. S2. The distance between electrodes was optimized by FTENG's transfer charge density study (Figure S3). The stator and the slider are respectively patterned into periodic complementary interdigital structures and grating structures by photolithography. After electroless nickel/immersion gold (ENIG) surface treatment is performed on the electrode area, the stator is further laminated with PTFE. The reusable flexible circuit and the disposable micro-sweat sensor patch can continuously perform electrochemical measurements of key biomarkers in sweat (

). During the movement, the power generated by FTENG is stored and released from the capacitor under the control of the power management integrated circuit (PMIC).

). After being fully charged, the storage capacitor releases its stored energy, which is adjusted to a stable voltage to power the BLE system-on-chip (PSoC) module and instrumentation amplifier to collect and transmit potential measurements through BLE.

The working mechanism of FTENG can be explained as the coupling effect of contact charging and in-plane sliding induced charge transfer, as shown in Figure 1.

. Since the triboelectricity of copper is higher than that of PTFE, electrons will accumulate on PTFE during sliding. In the initial state, the grating slider completely overlaps a stator electrode, and due to the electrostatic balance, no charge flow occurs between the designated sub-electrodes of the fork. The one-way sliding process causes a charge flow between the stator electrodes until the grating slider completely overlaps the second stator electrode with the opposite polarity. The numerical simulation using COMSOL Multiphysics further verified the working process (Figure S4). The detailed model of FTENG under open circuit and short circuit conditions is illustrated in Figure 5. S5 and notes S1 and S2. Our optical microscopic image of FTENG based on FPCB and typical short-circuit current (

) Cross-sectional view of FTENG at different operating frequencies

. FTENG operates continuously at varying frequencies of 0.5, 1.25, and 3.3 Hz to obtain maximum

They are 8.39, 19.11 and 42.25μA respectively. Open circuit voltage (

The frequency obtained at a frequency of 0.5 Hz is shown in FIG. 5. S6A, the signal polarity of the envelope waveform oscillates rapidly along the sliding process. To evaluate the use of our FPCB-based FTENG as a power source, voltage and power were measured under a series of different load resistances (

), its operating frequency is 1.5 Hz. An increase in resistance exceeding 1 megohm will cause a rapid increase in voltage. The load resistance of FTENG is 4.7 megohms and the maximum output power is 0.94 mW (equivalent to 416 mW m

) Schematic diagram of FTENG's working mechanism and charge distribution. (

) Microscopic images and optical images of FPCB-based interdigital stators, which have ENIG surface finish on the patterned electrode area. The scale bars are 200 μm and 5 mm. (

) Current output of FTENG at different operating frequencies. (

) The peak voltage and corresponding average power of FTENG under different external load resistances (

= 5). The working frequency is 1.5 Hz. (

) After 20,000 test cycles, the durability of M-PDMS–, W-PDMS– and PTFE stators.

Indicates the maximum open circuit voltage after and before the endurance test. Illustration after 20,000 test cycles, SEM images of M-PDMS, W-PDMS and PTFE. The scale bars are 5, 50 and 50 μm. (

) A schematic diagram of a flexible FTENG with different stator layouts (one parallel, three or six panels) and corresponding slider layouts. (

) Compare the voltage of capacitors from 10 to 1000 μF, and charge 30 cycles with one, three and six panel FTENG. (

) The long-term stability of the three-panel FTENG charging a 47μF capacitor for 2 hours at a working frequency of 1.5 Hz.

After 20,000 working cycles, FPCB-based FTENG's PTFE exhibited superior durability to traditional miniature pyramidal polydimethylsiloxane (M-PDMS) and wrinkled PDMS (W-PDMS) (


As shown in Figure 2, the attenuation is minimal. S6B. Scanning electron microscope (SEM) image

And figure. S7 revealed the shape of different friction materials before and after the durability test: PTFE showed excellent mechanical strength without scratches, while M-PDMS and W-PDMS both suffered obvious surface damage. During normal use, the influence of normal force and shear force (reflected by sliding frequency) on the performance of FTENG is shown in Figure 1. S8: The peak output voltage increases with the increase of the normal force, and then reaches saturation. For a given normal force, the output voltage remains stable under the changing sliding frequency. FTENG is mechanically robust and shows similar electrical output even at a high normal force of 100 N. The response of FTENG is stable after 1000 bending cycles (with a radius of curvature of 5 cm) (Figure S9) and at varying physiological temperatures (Figure 9). S10). In addition, FTENG can maintain high performance after 100 washing cycles, which shows that it has superior wearable performance compared with traditional TENG (Figure S11 and Table S1). When designing future TENG power supply equipment, it is important to consider factors such as cost, materials, mechanical properties and power density. The cost of TENG prepared by fabric weaving and polymer coating process is very low, but it is subject to lower manufacturing resolution and repeatability. In contrast, FPCB-based FTENG provides a high-resolution, cost-effective and mechanically robust energy harvesting solution.

In order to meet the high energy demand of wearable sensors, one, three and six panels of FTENG were designed by considering the size of the human torso, and further evaluated by capacitor charging (

). The output of FTENG is rectified with a full-wave rectifier. These different FTENG layouts are all driven for 30 working cycles to match 10 to 1000 μF (

). For a 1000μF capacitor, one, three and six plates can obtain voltages of 0.03, 0.12, and 0.19 V, respectively, showing a strong charging ability. At a working frequency of 3.3 Hz, the six-panel FTENG shows the largest transfer charge (σ

) Reach 15.73μC in one work cycle (Figure S12A). At the same time, figure. S12B depicts the charging and discharging curves of different capacitors charged to 2 V at a working frequency of 2 Hz using a three-panel FTENG. The three-panel FTENG starts at a working frequency of 1.5 Hz and can recharge a 47μF capacitor within 2 hours from 0 to 2 V (

), indicating that the long-term cycle stability is very high. FTENGs can also be used to charge various capacitors with different cycle lengths (Figure S13). Depending on the specific application, connecting multiple FTENGs in parallel may be a practical and attractive strategy that can greatly increase power output.

A schematic diagram of a dual biosensor array based on ion-selective electrodes (ISE) for sweat analyte analysis is depicted. The laser-engraved microfluidic channels are assembled on the sensor patch. The detailed manufacturing procedures are listed in Materials and Methods and Figure 2. S14. The Ag/AgCl reference electrode is coated with polyvinyl butyral (PVB). Regardless of the ionic strength of the solution, it can maintain a stable potential in the potential measurement of various electrolytes in sweat. Deprotonation of H when used for pH analysis

Measure the atoms on the surface of electrodeposited polyaniline (PANI) layer as H indicator

concentration. Na

Ion selective membranes containing Na facilitate concentration measurement

The ionophore X and the poly(3,4-ethylenedioxythiophene (PEDOT): poly(4-styrene sulfonate) (PSS) layer) between the gold electrode layer and the sodium ion selective membrane serve as ion electrons The converter can minimize potential drift as shown in the figure.

, PH and Na

In the physiologically relevant pH value (4 to 8) and Na, the sensor showed a sensitivity of near nerve energy of 56.28 and 58.63 mV for every ten times the concentration.

Concentration (12.5 to 200 mM). Both sensors have excellent selectivity, repeatability and long-term stability (Figure S15 to S17), and their response remains stable under different physiological temperatures (Figure S18), making them suitable for continuous wearable monitoring .

) Schematic diagram of a flexible biosensor array containing pH sensors and Na

The sensor is patterned on a flexible PET substrate. (

) Open circuit potential response of the pH sensor in standard Mcllvaine buffer solution (B) and Na

Sensor in NaCl solution (C). The illustration shows the corresponding calibration chart for each sensor. Error bars represent SD from six independent tests. (

) Schematic diagram of microfluidic design for dynamic sweat sampling. M tape, medical tape. (

Dynamic response of Na

Sensors at different flow rates when switching solution concentration. (

Repeatability of Na dynamic response

The sensor is realized by continuously switching the inflow solution at a flow rate of 2μlmin

. (

) A schematic diagram of a microfluidic sensor patch attached to human skin. The inset is an optical image of the microfluidic sensor patch under mechanical deformation. Scale bar, 5 mm. (

Na reaction

PH sensor array after 0, 200, and 400 cycles of bending (H) and during bending state (I) (radius of curvature of 2 cm). Data logging pauses for 30 seconds to change conditions and settings. Image courtesy: Yu Song, California Institute of Technology.

The laser-patterned microfluidic layer is connected to a polyethylene terephthalate (PET) sensor substrate in a sandwich structure (medical tape/PDMS/medical tape) for controlled and automatic in-body sweat sampling (

And figure. S19). In order to verify the performance of the microfluidic system, dynamic biosensing was carried out during the continuous flow injection of Na

Physiologically relevant sweat rate (1, 2 and 4μl min

). Donna

The concentration was switched from 50 mM to 200 mM at a flow rate of 2μlmin


It takes about 2 minutes for the sensor to reach a new stable reading. The high time resolution is repeatable in multiple concentration change cycles (

). The flexible microfluidic sensor patch can conformally adhere to human skin (

), and have passed a rigorous bending test (curvature radius of 2 cm) showing excellent mechanical stability, indicating their potential for wearable applications in various sports activities (

As mentioned earlier, FWS

It consists of an interdigital FENG stator, a PMIC, a low-dropout regulator, two low-power instrumentation amplifiers and a BLE PSoC module seamlessly integrated into a polyimide-based FPCB. In addition, the complete platform requires a grating patterned FTENG slider and microfluidic sensor patch. For design compatibility and flexibility, FTENG and electronic circuits are designed on a single PCB design software. The detailed parts list and circuit diagram of the flexible circuit are shown in Figures 1 and 2. S20 and S21 respectively. The block diagram shows the electrical connections between the modules

. In order to achieve the best power management, commercial energy harvesting PMICs are used to manage the power generated by FTENG, while minimizing power waste. With the aid of a bridge rectifier, the rectifier converts the high-voltage AC signal generated by FTENG into a DC signal, and the PMIC stores the power generated by FTENG in two parallel capacitors (220 and 22μF). Three SET_

The resistor sets a programmable threshold and hysteresis voltage to release the stored power through the built-in switch control logic only when absolutely necessary. When the voltage of the storage capacitor (

) Reaches 3.5 V, then the capacitor to the load/output (

) Until

Drop to 2.2V. At 2.2 V, the control unit of the PMIC disconnects the storage capacitor from the load/output until the storage capacitor is charged back to 3.5V. When powered by the storage capacitor, the load/output voltage is adjusted to 2.2 V through the regulator to provide a stable voltage for the precision measurement circuit.

) Schematic diagram of FWS without battery

It consists of FTENG module, biosensor interface, instrumentation amplifier, energy harvesting PMIC, voltage regulator and BLE PSoC module. (

) FWS operation process

Perform signal processing and data transmission. (

) Power consumption of FWS

During the operation. (

) Real-time potential of capacitor during FWS continuous operation (242μF)

Use three-panel FTENG under different operating frequencies. (

) Verify data transmission from FWS

) Long-term stability of capacitor charging process during FWS

It operates at a working frequency of 1.5 Hz. (

Sensor response in human sweat samples collected by FWS

It operates at a working frequency of 1.5 Hz.

Efficient power management matches the low-power measurement performed by a low-power instrumentation amplifier with shutdown mode, and low-power data transmission through connectionless BLE advertising, thus realizing FTENG-powered wearable and wireless sweat analysis. Every time the storage capacitor is charged to 3.5 V, the BLE PSoC module will start a ~510-ms work cycle, as shown in the flowchart (

). After the main processor starts, PSoC pulls the general-purpose input/output (GPIO) pin high to wake up the two instrumentation amplifiers from the shutdown state. After initializing the instrumentation amplifier, PSoC's embedded 12-bit ADC (analog-to-digital converter) samples and averages 32 potential measurements obtained through the instrumentation amplifier. After ADC measurement, the instrumentation amplifier will be turned off to minimize power consumption. PSoC's BLE sub-module requires a 32 kHz watch crystal oscillator (WCO) to operate accurately, and its maximum startup time specification is 500 ms. Therefore, after the ADC measurement, the PSoC main processor starts the WCO, enters a 500 ms deep sleep state, and consumes about 2μA current. Then, the BLE stack is initialized, and the ADC measurement result is notified to nearby BLE observer user equipment. The detailed power consumption breakdown of the circuit including the regulator, BLE PSoC module and two instrumentation amplifiers is shown in the following figure:

. When a 2.2 V power supply is provided, the circuit consumes an average of 330 μA in ~510 ms (168 μC).

Some studies were conducted to verify the robustness of the fully integrated system. The three-panel FTENG is activated by sliding motion at a frequency of 2 to 1 Hz to simulate human arm swing during exercise (

). The final charge and discharge cycle of the storage capacitor is shown in the figure.

. In addition, in order to verify the operation of the low-power wireless sensor circuit, a voltage of 100 to 300 mV (charged every 300 s) was applied to the reference electrode and working electrode pin by using a DC power supply to simulate potential input.

). These analog sensor inputs are accurately measured and transmitted by the FPCB platform, and are powered by the three-panel FTENG, which is activated at different operating frequencies. Long-term stability of the entire FWS

Demonstrate the system by using FTENG to power the FPCB for more than 4 hours, during which pH and Na

Measure the concentration of human sweat collected within one hour (

). In addition, by comparing the ability of FPCB-based FTENG to power the entire platform one month after its first use, its long-term durability was tested (Figure S22). By further improving the power density and efficiency of FTENG, wireless data transmission can be improved in terms of transmission interval.

Common cardiovascular exercises such as running, rowing, and elliptical training can cause sliding between the side of the torso and the inner arm. Using this mechanical movement, the stator of FTENG can be fixed on the side trunk, and the slider of FTENG can be installed inside the arm. For human evaluation, FWS based on six-plate stator FTENG is used

Used to increase power output (as shown in Figure S23). FTENG power output waveforms during various exercises are shown in the figure

. Choose a treadmill as an exercise, and conduct a human body verification experiment on the entire system. During the 60-minute constant speed operation, the FPCB storage capacitor charging and discharging curve shows that up to 18 operating cycles can be achieved (

). The length of the charge/discharge cycle ranges from 2.1 to 3.7 minutes (

). It should be noted that when the stator and slider physically rub against each other, the system generates power. Whenever there is frictional movement, the charge in the capacitor will accumulate without discharging. When the capacitor is charged to the threshold voltage, the capacitor will discharge and power a single measurement event. Although the duration of the capacitor charging/discharging cycle varies due to changes in friction area, force and frequency, FWS

The system proves that it can function normally during normal physical exercise (Figure S24). Human performance of the entire FWS

Healthy subjects were evaluated by a treadmill at a constant speed of 9 km/h

. Two wearable systems charged by FTENG and batteries are placed on the subject's back. The physiological information collected by the two systems is wirelessly transmitted to the user interface via BLE for further analysis (

). Five measurements have been recorded from FWS

During 30 minutes of exercise; stable pH and increased Na

Observe the level from both systems (

), confirm the accuracy of FWS without battery

Used for human body induction. Noise contribution of subjects wearing FWS during various exercises

It is insignificant compared to the sensor signal of interest (Figure S25). These data prove the potential of the self-powered wearable platform to continuously monitor various physiological biomarkers in sweat during exercise.

) The output waveform of FTENG based on six-panel FPCB in various exercises. (

) When the subject runs for 1 hour at a constant speed of 9 km/h on a treadmill, the real-time potential of the capacitor charged by FTENG (B) and the average charging time per package transmission (C)

. The ratio in (C) represents the percentage of charging cycles in all charging cycles (charging duration within a given time range). When the potential reaches 3.3 V, the capacitor discharges due to BLE data transmission. (

) Optical image of an object on a treadmill wearing FWS

And a cell phone. (

Real-time sweat pH and Na

Electricity obtained wirelessly from a wearable system during constant speed operation, which is charged by a lithium battery and FTENG. Image courtesy: Yu Song, California Institute of Technology.

Emerging wearable technology has achieved numerous personalized medical applications. Wearable sweat analysis may achieve non-invasive and continuous monitoring of personal health at the molecular level. Due to multi-function and multi-tasking requirements, wearable sweat biosensors usually have high power consumption. Batteries are the main power source for most wireless electronic skin systems, but they are usually limited by availability, especially when power supplies are limited. Given that the main application of sweat sensing is health and fitness tracking during strenuous exercise, harvesting energy from the human body is a promising method for powering future wearable sweat sensors, especially for moving organisms A sensor that converts mechanical energy into electrical energy.

The emergence of TENG technology has caused great excitement due to its potential application in self-powered systems (especially wearable and implantable electronic products). As an emerging energy conversion technology, TENG faces major challenges that need to be resolved in practical applications. First, the TENG signal is essentially a high-voltage pulse, which is not enough to meet the real-time energy consumption of wearable electronic devices. Second, for continuous use of wearables, due to the stability limitations of organic polymer materials used in device manufacturing, the life span of TENG needs to be increased. Last but not least, the system integration of TENG in wearable devices and the demonstration of their usability in practical applications are far from enough.

Here, we propose a highly durable, mass-produced, fully self-powered battery-less wearable system to meet these challenges. The system can effectively and reliably move from the human body during strenuous exercise through FPCB-enabled FTENG Collect energy in. Compared with traditional TENG, FTENG manufactured using commercial FPCB manufacturing procedures has excellent mechanical and electrical stability even after severe mechanical deformation and repeated cleaning cycles. Through seamless system integration and effective power management, this fully flexible system can provide power for human sweat biosensors and wirelessly send data to the user interface via Bluetooth during human testing. Compared with the previously reported non-wearable wireless sensor system based on TENG (Table S2), the wireless sensor system is either not wearable or requires an extra long charging time to perform the measurement. In contrast, FWS

It represents a breakthrough in the practicality of wearable applications. We envision that with further development, this technology will become a very attractive method for self-powered wireless personalized health monitoring in people's daily activities. It will also find many applications in the environment and defense fields.

EDOT, PSS, ionophore X, bis(2-ethylethylhexyl) sebacate (DOS), PVB, polyvinyl chloride (PVC), tetra[3,5-bis(trifluoromethyl)phenyl ] Sodium borate (Na-TFPB), aniline, sodium thiosulfate pentahydrate (Na



), sodium bisulfite (NaHSO

), calcium chloride dihydrate (CaCl

·2 hours

O), block polymer PEO-PPO-PEO (F127), multi-walled carbon nanotubes (MWCNT), iron (III) chloride (FeCl)

), potassium hydroxide (KOH) and citric acid were purchased from Sigma-Aldrich. Sodium chloride (NaCl), ammonium chloride (NH

Cl), methanol, ethanol, acetone, tetrahydrofuran (THF), hydrochloric acid (HCl), tetrachloroauric acid (HAuCl

) And disodium hydrogen phosphate (Na

high pressure

) Purchased from Thermo Fisher Scientific. PDMS (SYLGARD 184) was purchased from Dow Corning Corporation. Silver Nitrate (AgNO

) Was purchased from Alfa Aesar. Waterproof double-sided medical tape (75μm thick) was purchased from Adhesives Research. The conductive silver paint was purchased from Structure Probe Inc. (SPI) consumables. Moisture-proof PET film (100μm thick) was purchased from McMaster-Carr. PTFE (50μm thick) was purchased from JIAET.

The FPCB module of FTENG and electronic circuit is designed using Eagle CAD (Autodesk). The BLE PSoC module is programmed in the PSoC Creator integrated design environment (Cypress Semiconductor). Figure 2 provides a complete list of components used in circuit design. S20 includes power management unit (MB10S-13, Diodes Incorporated; S6AE101A, Cypress Semiconductor; TPS7A05, Texas Instruments), BLE PSoC module (CYBLE-022001-00, Cypress Semiconductor), potential detection unit (AD8235, Analog Devices) and passive element. Figure 2 shows a detailed circuit diagram. S21.

The flexible circuit and FTENG are manufactured by commercial FPCB manufacturers (the detailed manufacturing process is shown in Figure S1). Two commercial flexible copper clad laminates (120μm thick; Jinghuang Electronics Co., Ltd.) composed of a flexible polyimide substrate and a copper film sandwiched a layer of epoxy adhesive. Pattern the copper film by photolithography and etch with FeCl

The solution is to manufacture the circuit elements and interdigital electrodes of the stator, as well as the complementary grating structure of the slider. The ENIG layer is deposited to protect the stator electrodes. Finally, a layer of PTFE is laminated on the interdigital electrodes of the stator to induce electrification. The total size of FTENG's single-plate stator is 22.6 cm

(Length 5.78 cm; width 3.78 cm). The weight of FTENG's single-plate stator is 0.586 (without PTFE coating) and 0.782 g (with PTFE coating). The total size and weight of FTENG's single-panel slider is 18.22 cm

(Length 4.36 cm; width 4.18 cm) and 0.396 grams.

In order to improve the output performance of FTENG, a digital oscilloscope (Agilent DSO-X 2014A) was used to test the open circuit voltage with a 100 MΩ probe. The short-circuit current is amplified by the SR570 low-noise current amplifier of Stanford Research Systems. COMSOL software is used to simulate and verify the electrostatic stimulation during sliding.

Microstructured PDMS (M-PDMS and W-PDMS) participated in the durability test as a triboelectric material in the contact separation mode. First, mix the PDMS elastomer and crosslinking agent with a ratio of 10:1. For M-PDMS, the vacuum degassed solution is spin-coated on a Si wafer with an inverted pyramid structure (manufactured by photolithography and KOH wet etching). After PDMS is partially cured, a FPCB-based stator (3×4 cm

Coated with ENIG electrodes). After curing at 80°C for 2 hours, the stator was peeled off with the prepared M-PDMS. For W-PDMS, the cured PDMS is pre-stretched at a strain of 30% and subjected to ultraviolet ozone treatment with a commercial ultraviolet lamp (Hangzhou Yaguang Lighting). After the release process, W-PDMS is applied and applied to another FPCB-based stator. During the durability test, the maximum gap between PDMS and another friction material (copper) was fixed at 3 mm at a working frequency of 2 Hz.

For the washing test, first rinse the FPCB-based FTENG with deionized (DI) water (25°C), and then perform bath sonication for 10 minutes. Then, FTENG was fully dried at 60°C for 10 minutes for later use.

For the long-term stability study of FTENG (Figure S22), the test was conducted under the same conditions 1 month before and after. During the 1 month interruption, FTENG was stored in a plastic box in a regular office drawer at room temperature.

An optical microscope (Carl Zeiss AXIO) was used to characterize the morphological microscopic image of the interdigitated stator electrode. The SEM images of PTFE, M-PDMS and W-PDMS were obtained by field emission environment SEM (FEI Quanta 600F).

The fabrication of the electrode array is shown in FIG. 2. S14. After pretreatment of the PET substrate, 20 nm Cr was deposited on the PET substrate using electron beam evaporation, and then 100 nm Au was deposited to form a gold electrode with a diameter of 3 mm. The electrode array is also coated with 1μm Parylene C (ParaTech LabTop 3000 parylene coating machine) and patterned by photolithography. Array made by further etching with O

Reactive ion etching (Oxford III-V System 100 ICP / RIE) removes the parylene layer in the plasma. Then, modify the electrode and deposit different functional materials to form Na

pH electrode and shared Ag/AgCl reference electrode. Carbon monoxide

A laser cutter is used to pattern the microfluidic layer. First, a waterproof double-sided medical tape layer with a cavity of 3mm diameter is pasted on the PET sensor substrate. Then, a container with a diameter of 3 mm, an inlet, an outlet, and a PDMS layer (thickness of 100 μm) with fluid connections were pasted on the medical tape. Finally, another layer of medical tape with an entrance pattern is attached to the PDMS layer.

Electrochemical workstation (CHI 860, CH Instruments) is used for electrochemical deposition and sensor characterization. For Ag/AgCl reference electrode, use Ag deposition solution (0.25 M AgNO

, 0.75 M Na

And 0.43 M NaHSO

) Is used to deposit Ag on Ag electrodes by constant voltage electrodeposition (-0.25 V for 600 s). Next, 0.1 M FeCl

It was drop cast on Ag for 30 seconds to form Ag/AgCl. A total of 6.6 μl of PVB reference mixture (79.1 mg PVB, 50 mg NaCl, 1 mg F127 and 0.2 mg MWCNT in 1 ml methanol) was drip-cast on the Ag/AgCl electrode to dry. First by depositing Au (50 mM HAuCl

Then dissolve in 50 mM HCl at 0 V for 30 s, and then perform 50 cycles of cyclic voltammetry (-0.2 to 1 V at 50 scan rate) on Au electrodes (0.1 M aniline and 0.1 M HCl) in a bath Electropolymerization of PANI. Millivolt second

). For Na

ISE performs constant current electrodeposition (14μA, duration 740 s) in a solution containing 0.01 M EDOT and 0.1 M NaPSS to deposit PEDOT:PSS on an Au electrode. Then, take 15μl Na

The selective membrane mixture was dropped onto the PEDOT:PSS layer and dried overnight. In order to prepare a cocktail, a mixture of 100 mg is required, which contains Na ionophore X (1%, w/w), Na-TFPB (0.55%, w/w), PVC (33%, w/w) and DOS (65.45) %), w/w) dissolved in 660μl of THF.

In order to obtain the best performance for long-term continuous measurement, cover the biosensor with a solution containing 0.1 M NaCl for 1 hour before the measurement to minimize potential drift. For in vitro characterization, unless otherwise specified, NaCl solutions of 12.5, 25, 50, 100, and 200 mM in deionized water and Mcllvaine buffer with a pH of 4 to 8 were used. Considering the difference in the absolute potential value of the ion-selective sensor in the same solution, it is important to perform a single-point calibration in the standard solution. Here, the biosensor was calibrated using a 25 mM NaCl solution before being used in all tests.

The batch biosensor is characterized by changing the solution to verify its repeatability and repeatability. Interference study by continuously adding chloride solution containing 50 mM NH

, 50 mm

And 50 mM Ca

. When changing the solution, all measurements are suspended for 30 s. Long-term stability of pH and Na

The sensor is first continuously tested in a 100 mM NaCl solution for 3 hours, and then evaluated for more than 6 weeks to check the sensitivity change.

Previous work concluded that the pH in sweat remains relatively stable during exercise. Therefore, the sampling ability of the microfluidic sensor patch focuses on the dynamic tracking of Na

. Use a syringe pump to inject different Na solutions

The concentration (50 and 200 mM) was changed at different flow rates through the inlet of the microfluidic channel. The mechanical reliability of the sensor patch was evaluated by repeatedly bending 800 cycles on a three-dimensional printed mold (radius of curvature, 2 cm). The sensor measurement value is obtained every 200 cycles. In another study, continuous sensor measurements were recorded during the active deformation of the sensor.

For the long-term sensor stability test (Figure S17), the biosensor array is tested under the same conditions every week. Before weekly measurements, cover the sensor array with a 0.1 M NaCl solution for 1 hour to minimize potential drift. Store the ion-selective sensor under ambient conditions at room temperature (25°C) for a period of 6 weeks.

Verification and evaluation of FWS

Human subjects were tested in the gymnasium and all ethics requirements under the protocol (ID 19-0892) approved by the California Institute of Technology Institutional Review Board were complied with. Healthy subjects aged 20 to 35 years were recruited from California Institute of Technology. Before participating in the study, all subjects gave written informed consent.

The subjects performed cardiovascular exercises using treadmills (Aeon), elliptical machines (Precor) and rowing machines (Stamina). Before exercise, wipe and clean the subjects’ upper back with alcohol swabs and gauze. Then, use waterproof double-sided medical tape to paste FWS

On the subject. The system containing the FTENG stator is adhered to the side torso and the FTENG slider is fixed to the inner arm. To ensure the accuracy of the data, a new microfluidic sensor patch was used in each human test. To evaluate the power output of FTENG during exercise, connect the output of FTENG or the voltage across the storage capacitor to an oscilloscope. When evaluating the entire system including the microfluidic sensor patch, the subject was asked to run on a treadmill at a constant speed of 9 km

30 minutes (obtain sensor data regularly every few minutes); BLE data is retrieved from mobile phones or personal computers. In addition, sweat samples were collected from the subjects’ foreheads on a regular basis, and then pipetted into a centrifuge tube and centrifuged at 6000 rpm for 15 minutes. The sweat samples were then frozen at -20°C for further testing.

For supplementary materials for this article, please visit:


This is an open access article distributed under the following terms

, It allows use, distribution and reproduction in any medium, as long as the final use is

For commercial interest, and provide the original works appropriately cited.

Volume 6, Number 40

September 30, 2020

Thank you for your interest in promoting the term "scientific progress".

Note: We only ask you to provide your email address so that the person you recommend to the page knows that you want them to see the email and that it is not spam. We will not capture any email addresses.

This question is used to test whether you are a visitor and prevent automatic spam submission.

Please log in to add an alert to this article.


: Eaay9842

The wireless battery-free wearable sensor driven by human motion can analyze sweat biomarkers to achieve personalized medical care.

Volume 371, Issue 6526


. all rights reserved. The American Association for the Advancement of Science is 


ISSN 2375-2548.