Patent Application
20180263539
Any and all applications for which a foreign or domestic priority claim is identified in the Application Data Sheet as filed with the present application are hereby incorporated by reference under 37 CFR 1.57.
The present disclosure was made by or on behalf of the below listed parties to a joint research agreement. The joint research agreement was in effect on or before the date the present disclosure was made and the present disclosure was made as a result of activities undertaken within the scope of the joint research agreement. The parties to the joint research agreement are: 1) The Regents of the University of California, and 2) The Board of Trustees of the Leland Stanford Junior University.
The present invention relates to devices that measure physiological parameters of a user.
Wearable electronics have been developed that can be worn by user's to continuously and closely monitor an individual's activities, such as walking and running, Such wearable electronics may include physiological sensors configured to sense certain physiological parameters of the wearer, such as heart rate, as well as motion sensors, GPS radios, and altimeters.
The following presents a simplified summary of one or more aspects in order to provide a basic understanding of such aspects. This summary is not an extensive overview of all contemplated aspects, and is intended to neither identify key or critical elements of all aspects nor delineate the scope of any or all aspects. Its sole purpose is to present some concepts of one or more aspects in a simplified form as a prelude to the more detailed description that is presented later.
An aspect of the disclosure relates to wearable biometric monitoring system comprising: a first sensor configured to sense a first sweat analyte; a second sensor configured to sense a second sweat analyte at substantially the same time as the first sensor is measuring the first sweat analyte; a signal conditioner coupled to the first sensor and the second sensor, the signal conditioner configured receive and condition sensor signals from the first sensor and the second sensor, the signal conditioner comprising one or more amplifiers and one or more filters; and an interface configured to transmit information corresponding to the conditioned sensor signals to a remote computing device.
An aspect of the disclosure relates to wearable biometric monitoring system comprising: a flexible substrate; a plurality of sweat analyte sensors affixed to the flexible substrate, the plurality of sweat analyte sensors configured to sense a plurality of different sweat analytes of a wearer at substantially the same time, the plurality of sweat analyte sensors comprising at least a first sweat analyte sensor configured to sense a metabolite and a second sweat analyte configured to sense an electrolyte; a temperature sensor configured to measure skin temperature of the wearer; a signal conditioner affixed to the flexible substrate, the signal conditioner coupled to the plurality of sweat analyte sensors, the signal conditioner configured receive and condition sensor signals from the plurality of sweat analyte sensors, the signal conditioner comprising one or more amplifiers and one or more filters; an analog and digital converter configured to convert the conditioned sensor signals from an analog domain to a digital domain, and a digital processor configured to digitally process the converted sensor signals in the digital domain, the analog and digital converter and the digital processor affixed to the flexible substrate; an interface configured to transmit information corresponding to the conditioned sensor signals to a remote computing device, the interface affixed to the flexible substrate; and a battery configured to power at least portions of the wearable biometric monitoring system.
An aspect of the disclosure relates to a method of fabricating a sweat analyte sensing system, the method comprising: patterning a flexible substrate with a sweat analyte sensor array; depositing a metal on the sweat analyte sensor array; depositing an insulating layer on the sweat analyte sensor array; defining electrode areas using photolithography and etching of the insulator; patterning a metal on the electrode areas; and forming reference electrodes corresponding to the electrode areas.
Embodiments will now be described with reference to the drawings summarized below. These drawings and the associated description are provided to illustrate example embodiments, and not to limit the scope of the invention.
Wearable sensor technologies may play a significant role in realizing personalized medicine through continuously (or periodically) monitoring an individual's health state. Disclosed herein are various example devices and sensors that can be used to sense various aspects of a user's physiological state. A wearable sensing platform is disclosed that may include some or all of the different sensors and circuits disclosed herein to sense, analyze, and report various aspects of a user's state.
To this end, human sweat is an excellent candidate for non-invasive monitoring as it contains physiologically rich information. Conventional sweat-based and other non-invasive biosensors either can only monitor a single analyte at a time or lack on-site signal processing circuitry and sensor calibration mechanisms for accurate analysis of the physiological state. Given the complexity of sweat secretion, simultaneous and multiplexed screening of target biomarkers and full system integration advantageously ensures the accuracy of measurements.
Disclosed herein is an optionally mechanically flexible and fully-integrated perspiration analysis system, including a wearable sensing platform, that simultaneously and selectively measures multiple sweat analytes, such as, by way of example, sweat metabolites (e.g. glucose and lactate) and electrolytes (e.g. sodium and potassium ions), optionally as well as the skin temperature to calibrate the sensors' response. Other sweat analyte sensors may be included as well or instead. For example, calcium, heavy metal, pH, and/or protein sensors may be included.
As discussed elsewhere herein, the panel of target analytes and skin temperature may be selected based on their informative role in understanding an individual's physiological state. Measuring and analyzing certain analytes (e.g., sodium, potassium, glucose, lactate, skin temperature, heavy metals, pH, etc.) may then be used to detect and monitor various physiological conditions. For example, excessive loss of sodium and potassium in sweat could result in hyponatremia, hypokalemia, muscle cramps or dehydration. Sweat sodium and potassium could be useful biomarkers for electrolyte imbalance and Cystic Fibrosis diagnosis. Sweat glucose comes from blood glucose. Thus, glucose monitoring is desirable in managing diabetes, and several studies have reported that sweat glucose levels are correlated with blood glucose levels. As such, sweat glucose sensing may serve as a non-invasive way for blood glucose monitoring.
Sweat lactate analysis may be helpful for many potential clinical applications. For example, sweat lactate has been shown to potentially be a very useful early warning indicator of pressure ischemia. Sweat lactate may also be used to monitor physical performance since lactate is a product of anaerobic metabolism. If there is an adequate correlation between blood and sweat lactate levels, the detection of sweat lactate may offer a non-invasive way for blood lactate monitoring. There are also reports on using sweat lactate as a biomarker for panic disorder or Frey's syndrome. Skin temperature is clinically informative of a variety of diseases and skin injuries such as pressure ulcers. Skin temperature is an effective indicator of human sensations and provides significant clinical information about cardiovascular health, cognitive state and malignancy. Additionally, skin temperature measurements may be used to compensate for and to reduce or eliminate the influence of temperature variation on the chemical sensors' readings, optionally through a built-in signal processing functionality, as discussed elsewhere herein.
Aspects of the disclosure bridges the technological gap between signal transduction, conditioning, processing and wireless transmission in wearable biosensors by merging sensors (e.g., plastic-based sensors), that interface with the skin, and silicon integrated circuits consolidated on a circuit board (e.g., a flexible circuit board, which optionally be configured to be worn around or on a wrist, arm, ankle, leg, head, chest, or other body party) for complex signal processing.
The disclosed wearable system may be used to measure the detailed sweat profile of human subjects engaged in prolonged indoor and/or outdoor physical activities, and infer real-time assessment of the physiological state of the subjects. The platform enables a wide range of personalized diagnostic and physiological monitoring applications.
Wearable electronics comprise devices that can be worn or mated with human skin to continuously and closely monitor an individual's activities, without unduly interrupting or limiting the user's motions. Accordingly, as noted above, wearable biosensors may play a significant role in realizing personalized medicine due to their capability in real-time and continuous monitoring of an individual's physiological biomarkers. Current commercially available conventional wearable sensors are only capable of tracking an individual's physical activities and vital signs (e.g. heart rate), and fail to provide insight into the user's health state at molecular levels.
To gain such insight, human sweat is an excellent candidate, as similarly discussed above, as it contains physiologically and metabolically rich information that can be retrieved non-invasively. Sweat analysis is currently used for applications such as disease diagnosis, drug abuse detection, and athletic performance optimization. Disadvantageously, for these applications, the sample collection and analysis are conventionally performed separately, failing to provide a real-time profile of sweat content secretion, while requiring extensive lab analysis using bulky instrumentations. Development of wearable sweat sensors has recently been explored where a variety of biosensors were used to measure analytes of interest (see, e.g., Supplementary Table 1 below).
Given the complex nature of sweat and the multivariate mechanisms that are involved in its secretion process, an attractive strategy would be to devise a fully-integrated multiplexed sensing system to extract insightful information from sweat. Aligned with this vision, disclosed herein is a wearable flexible integrated sensing array (FISA) for simultaneous and selective screening of a panel of biomarkers in sweat (see, e.g.,
With reference to
For example, the glucose sensor (with current output), has the current output amplified using a trans-impedance amplifier (1), whose output is inverted by an inverter (1), and the output of the inverter is filtered using a low-pass filter (2). By way of further example, the lactate sensor (with current output), has the current output amplified using a trans-impedance amplifier (3), whose output is inverted by an inverter (3), and the output of the inverter is filtered using a low-pass filter (4). The temperature sensor, which has a resistance that various in accordance with temperature, provides a voltage output that is divided down using a voltage divider. The sodium sensor (with voltage output), has its output buffered using a voltage buffer (5), the output of the voltage buffer (5) is amplified using a differential amplifier (6), and the output of the differential amplifier (6) is filtered using a low-pass filter (7). Potassium sensor (with voltage output), has its output buffered using a voltage buffer (8), the output of the voltage buffer (8) is amplified using a differential amplifier (6), and the output of the differential amplifier (6) is filtered using a low-pass filter (8). The analog outputs of the low-pass filters (2), (4), (7), (9), are feed into an analog-to-digital converter (ADC), which may be integrated into a processing device, such as microcontroller (10). The ADC converts the analog signals to digital signals, and the processing device may then process the digitized signals. For example, the processing device may calculate physiologic data using some or all of the data from one or more of the sensors. Optionally, the processing device and signal conditioning circuitry may be integrated into a single device.
A wireless interface (e.g., Bluetooth transceiver 10) may be used to wirelessly communicate or facilitate communication (e.g., of the processed sensor readings) to a remote device (e.g., a mobile device, such as a cell phone, tablet computer, laptop, etc., or a non-mobile device, such as a desktop computer or large screen networked television). Thus, for example, the wireless interface may facilitate or provide connectivity to, for example, relatively local external devices and/or remote devices via the Internet. The remote device may provide user interfaces that display (e.g., in real time and/or at later time) the sensor data and the remote device may upload the sensor data to a cloud system comprising one or more cloud servers or to other devices in association with a user and/or device identifier. The cloud system (or other device) may then store the sensor data (which may have been first processed by the wearable device processing system) in a data store in an account record associated with the user and/or the wearable device for later access and/or for further processing by the cloud system.
In particular, as illustrated in
The measurement of Na+ and K+ levels are facilitated through the use of ion selective electrodes (ISEs), coupled, in this example, with a polyvinyl butyral (PVB) coated reference electrode to maintain a stable potential in solutions with different ionic strengths (see, e.g.,
Optionally, the FISA may include a display (e.g., an LCD, OLED, e-ink, or other display). The display may be coupled to the processor and may display various sensor measurements, alerts, and/or information derived from the sensor measurements. For example, the processor may utilize the display to present information on wearer dehydration, and the likelihood or presence of hyponatremia, hypokalemia, muscle cramps, ischemia, and/or pressure ulcers. The information may be displayed in an association with corresponding icons (e.g., alert icons indicating that the user is or is about to suffer an adverse physiological condition). Similarly, the mobile application may be configured to determine and/or display, via the mobile device, similar information. The display may be touch sensitive and/or the FISA may include physical controls, such as physical buttons or knobs. The wearable device may include a microphone configured to receive voice commands and may include a speaker to provide audible confirmation of the voice commands. The wearer may command the FISA to display the information measured and/or determined via analysis via the touch display, voice commands, and/or the physical controls.
Testing of an example implementation of the FISA was performed. The performance of each sensor was monitored separately with respective analyte solutions.
In particular,
It is desirable for wearable devices to have the ability of withstanding stress from daily human wear and physical exercise in order to be utilized on a daily basis by typical users. A study on mechanical deformation conducted by monitoring the performance of both the sensor array and the FPCB before, during, and after bending (radii of curvature are 1.5 cm and 3 cm, respectively) (see, e.g.,
The FISAs can be configured to be comfortably worn on various body parts, including, for example, the forehead, wrists, and/or arms.
In particular,
Real-time physiological monitoring was performed on a subject during constant-load exercise on a cycle ergometer. In this example, the protocol involved a 3-minute ramp up, a 20-minute cycling at 150 W, and a 3-minute cool down. During the exercise, the heart rate (HR), oxygen consumption (VO2), and minute ventilation (VE) were measured using external monitoring instruments, and were found to increase proportionally with increasing power output (PO) as shown in
Sweat analyte levels on the wrist follow similar trends but with different concentrations from the forehead (see, e.g.,
The physiological response of the subjects due to a sudden change in exercise intensity was also investigated in a graded-load exercise which involved a 5-minute rest, a 20-minute cycling at 75 W followed by a cycling at 200 W PO until volitional fatigue, and a 10-minute recovery (
Monitoring hydration status is highly desirable for athletes, as fluid deficit impairs endurance performance and increases carbohydrate reliance. To evaluate the utility of FISA in identification of the dehydration status as an effective and non-invasive approach, real-time sweat [Na+] and [K+] measurements were conducted simultaneously on a group of subjects engaged in prolonged outdoor running trials (
Here, in the illustrated example implementation, skin-conforming plastic-based sensors (5 different sensors in this example, although additional or fewer sensors may be used) and IC components (e.g., conventional commercially available or custom IC components, including more than 10 chips in this example) are merged at high level of integration, to not only measure the output of an array of multiplexed and selective sensors, but to also through signal processing obtain accurate assessment of physiological state of the human subjects. The envisioned application could not have been realized by either of the technologies alone due to their respective inherent limitations. The plastic-based device technologies lack the ability to implement sophisticated electronic functionalities for critical signal conditioning and processing. On the other hand, the silicon IC technology does not provide sufficiently large active areas nor intimate contact to skin needed to achieve stable and sensitive on-body measurements. Importantly, the entire system may be mechanically flexible and self-sustained, thus delivering a practical wearable sensor technology that can be used for prolonged indoor and outdoor physical activities. The same platform can be configured for in-situ analyses of other biomarkers within sweat and other human fluid samples to facilitate personalized and real-time physiological and clinical investigations. The large data sets that are collected through such studies along with voluntary community participation would enable application of data mining techniques to generate predictive algorithms for understanding the health status and clinical needs of individuals and the society as a whole.
Certain example optional techniques and processes, and related optional example materials, circuits, dimensions, and amounts, will now be described. In addition certain measured results of fabricated components and devices will also be described. It is understood that other techniques and processes, and related optional example materials, circuits, dimensions, and amounts, may be used.
Example Materials. Selectophore™ grade sodium ionophore X, bis(2-ethylehexyl) sebacate (DOS), sodium tetrakis[3,5-bis(trifluoromethyl)phenyl] borate (Na-TFPB), high-molecular weight polyvinyl chloride (PVC), tetrahydrofuran (THF), valinomycin (potassium ionophore), sodium tetraphenylborate (NaTPB), cyclohexanone (CHA), polyvinyl butyral resin BUTVAR B-98 (PVB), sodium chloride (NaCl), 3,4-ethylenedioxythiophene (EDOT), poly(sodium 4-styrenesulfonate) (NaPSS), glucose oxidase (GOx) (from Aspergillus niger), chitosan (CS), single-walled carbon nanotubes (SWCNTs), iron (III) chloride, potassium ferricyanide (III), multiwall carbon nanotubes (MWCNTs), block polymer PEO-PPO-PEO (F127), L-Lactate oxidase (LOx) (activity, >80 U/mg), phosphate buffered saline (PBS)—pH 7.2, moisture-resistant polyethylene terephthalate (PET)—100 μm thick.
Example Fabrication of electrode arrays. An example fabrication process of the electrode arrays is demonstrated in
Example Design of electrochemical sensors. For amperometric glucose and lactate sensors, a two-electrode system where Ag/AgCl acts as both reference and counter electrode was chosen to simplify circuit design and to facilitate system integration. The two-electrode system is a common strategy for low current electrochemical sensing. Other sensor configurations may be used. The output currents (between working electrode and Ag/AgCl reference/counter electrode) of the glucose and lactate sensors could be converted to a voltage potential through a transimpedance amplifier. amperometric sensors with larger area provide larger current signal. Considering the low concentration of glucose in sweat, the sensors are designed to be 3 mm in diameter to obtain a high current, although other diameters may be used.
Preparation of Na+ and K+ selective sensors. The Na+ selective membrane cocktail in this example comprises Na ionophore X (1% w/w), Na-TFPB (0.55% w/w), PVC (33% w/w), and DOS (65.45% w/w). In this example, 100 mg of the membrane cocktail was dissolved in 660 μL of THF17. The K+ selective membrane cocktail was composed of valinomycin (2% w/w), NaTPB (0.5%), PVC (32.7% w/w), and DOS (64.7% w/w). In this example, 100 mg of membrane cocktail was dissolved in 350 μL of CHA. The ion selective solutions were sealed and stored at 4° C. The solution for the PVB reference electrode was prepared by dissolving 79.1 mg PVB and 50 mg of NaCl into 1 mL methanol 36. 2 mg F127 and 0.2 mg MWCNTs were added into the reference solution to reduce or minimize the potential drift.
Poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) was chosen for this example implementation as the ion-electron transducer to minimize the potential drift of the ISEs37 and deposited onto the working electrodes by galvanostatic electrochemical polymerization with an external Ag/AgCl reference electrode from a solution containing 0.01 M EDOT and 0.1 M NaPSS. A constant current of 14 μA (2 mA/cm2) was applied to produce polymerization charges of 10 mC onto each electrode.
Ion-selective membranes were then prepared by drop-casting 10 μL of the Na+ selective membrane cocktail and 4 μL of the K+ selective membrane cocktail onto their corresponding electrodes. The common reference electrode for the Na+ and K+ ISEs was modified by casting 10 μl of reference solution onto the Ag/AgCl electrode. The modified electrodes were left to dry overnight. The sensors could be used without pre-conditioning (with a small drift of ˜2-3 mV/h). However, to obtain the best performance for long-term continuous measurements such as dehydration studies, the ion-selective sensors were covered with a solution containing 0.1 M NaCl and 0.01 M KCl through microinjection (without contact to glucose and lactate sensors) for 1 h before measurements. This conditioning process was important to further minimize the potential drift.
Example preparation process for preparation of lactate and glucose sensors. 1% CS solution was first prepared by dissolving CS in 2% acetic acid and magnetic stirring for about 1 h; next, the CS solution was mixed with SWCNTs (2 mg/mL) by ultrasonic agitation over 30 minute to prepare a viscous CS/CNTs solution. To prepare the glucose sensors, the CS/CNTs solution was mixed thoroughly with GOx solution (10 mg/mL in PBS—pH 7.2) by a ratio of 2:1 (v/v). A Prussian blue (PB) mediator layer was deposited onto the Au electrodes by cyclic voltammetry from 0 V to 0.5 V (vs. Ag/AgCl) for 1 cycle at a scan rate of 20 mV/s in a fresh solution containing 2.5 mM FeCl3, 100 mM KCl, 2.5 mM K3Fe(CN)6, and 100 mM HCl. A thinner PB layer can provide better sensitivity which is essential for the low glucose level measurements in sweat. The glucose sensor was obtained by drop-casting 3 μL of the GOx/CS/CNTs solution onto the PB/Au electrode. For the lactate sensors, the PB mediator layer was deposited onto the Au electrodes by cyclic voltammetry from −0.5 V to 0.6 V (vs. Ag/AgCl) for 10 cycles at 50 mV/s in a fresh solution containing 2.5 mM FeCl3, 100 mM KCl, 2.5 mM K3Fe(CN)6, and 100 mM HCl. The thicker PB layer can provide a wider linear response range which is advantageous for the lactate measurement in sweat. 3 μL of CS/CNTs solution was casted onto the PB-Au electrode and dried under ambient environment; the electrode was later covered with 2 μL of LOx solution (40 mg/mL) and finally 3 μL CS/CNTs solution. The sensor arrays were allowed to dry over-night at 4° C. with no light. The solutions were stored at 4° C. when not in use.
Example signal conditioning, processing and wireless transmission circuit design. An example circuit diagram of the analog signal-conditioning block of the flexible integrated sensor array (FISA) is shown in
The example conditioning path for each sensor is implemented in relation to the corresponding sensing mode. In the case of the amperometric-based glucose and lactate sensors, the originally generated signal is in the form of electrical current. Therefore, in the respective signal conditioning paths, a transimpedance amplifier stage is used to convert the signal current into voltage. In the electrical current measurements, the direction of the current is from the shared Ag/AgCl reference/counter electrode toward the working electrode of each of the glucose and lactate sensors, which would result in a negative transimpedance output voltage. Hence, for both glucose and lactate paths, the transimpedance amplifiers are followed by inverter stages to make the respective voltage signals positive, since the example ADC stage only is configured to receive positive input values (although other ADCs may be used the are configured to received negative input values, or both negative and positive input values). The feedback resistors in each of the transimpedance sections was chosen (1 MΩ for the glucose path and 0.5 MΩ for the lactate path) such that the converted voltage signal could be finely resolved, while staying within the input voltage range of the ADC stage of the microcontroller. The current sensing signal paths are capable of measuring current levels as low as 1 nA, which was significantly lower than the minimum signal in the measurements (˜10s of nA).
In this example implementation, with the transimpedance amplifier at the front-end, the Ag/AgCl reference/counter electrode of the amperometric-based sensors needed to be grounded, which prevent grounding the shared PVB reference electrodes in the potentiometric-based sensors, as the potential difference between the Ag/AgCl reference and PVB electrodes changes in the presence of different chloride ions concentrations (
Also, the high impedance nature of the ISE-based sensors makes advantageous the use of high impedance voltage buffers to ensure accurate open voltage measurement as intended.
The analog signal conditioning paths include a corresponding unity gain four-pole low pass filters, each with a −3 dB frequency at 1 Hz to minimize the noise and interference in the measurements. Utilizing active filters in in the system also provides flexibility in tuning the gain in the signal-conditioning path if needed or desired. The low pass filters are connected to the ADC stage of the microcontroller, to facilitate the conversion of the filtered analog signals to their respective digital forms. In an example implementation, each of the analog signal conditioning paths were electrically characterized to validate the linear output response of the channels with respect to the corresponding electrical input signals mimicking the sensor output signals. For this characterization, electrical current was applied as an input to the glucose and lactate channel terminals to model the respective amperometric-based sensor output and differential voltage was applied at the terminals of the sodium and potassium channels to model the corresponding potentiometric-based sensor output. As illustrated in
The example power delivery to the FISA. The FISA is powered, in this example, by a single rechargeable lithium-ion polymer battery with a nominal voltage of 3.7 V of a desired capacity (a representative 105 mAh battery is illustrated in
The example mobile application design. A mobile application (Perspiration Analysis App) may accompany and communicate with the FISA system to provide a user-friendly interface for data display and aggregation (see, e.g.,
The characterization of the sensors. A set of example electrochemical sensors was characterized to explore their reproducibility in solutions of target analytes.
For continuous use, all the sensors displayed excellent stability over the entire exercise period. The sensor array could be repeatedly used for continuous temperature and sweat electrolyte monitoring. However, the glucose and lactate responses degraded beyond the exercise period (after two hours) due to decreased enzyme activity. The devised sensor-FPCB interface allows for convenient replacement of the fresh sensor arrays for subsequent use.
Analysis of the effect of mechanical deformation on the sensors was performed by repeatedly bending Na+, glucose sensors, and temperature sensors (radius of curvature, 1.5 cm) as well as FPCB (radius of curvature, 3 cm) for a total of 60 cycles (
Ex-situ evaluation of the sweat samples. Ex-situ sensor performance was also conducted by testing subjects' sweat samples collected from their forehead. Sweat samples were collected every 2 to 4 minute by scratching their foreheads with microtubes, and subjects' foreheads were wiped and cleaned with gauze after every sweat collection 19. The changes of [Na+] and [K+] during euhydration and dehydration trials were also studied ex-situ in the same manner. The calibration of the sensor arrays was performed prior to ex-situ measurements using artificial sweat containing 22 mM urea, 5.5 mM lactic acid, 3 mM NH4, 0.4 mM Ca2+, 50 μM Mg2+, and 25 μM uric acid with varying [glucose] from 0 to 200 μM, [K+] from 1 to 16 mM and [Na+] from 10 to 160 mM.
The example setup of FISA for on-body testing. A water absorbent thin rayon pad was placed between the skin and the sensor array during on-body experiments to absorb and maintain sufficient sweat for stable and reliable sensor readings, and to prevent direct mechanical contact between the sensors and skin. The pad could only absorb ˜10 μL sweat which was sufficient to provide stable sensor readings. During on-body tests, the newly generated sweat would refill the pad and ‘rinse away’ the old sweat. The on-body measurement results were also consistent with ex-situ tests using freshly collected sweat samples. The refill time was estimated to be less than 1 minute based on the sweat rate (˜3-4 mg/min/cm2) and the pad size (1.5 cm×2 cm). The intrinsic response time of FISA was smaller than body's response time to the changes in physiological conditions. An increase in temperature was observed when the smart headband or smart wristband was worn due to the use of the plastic substrate on skin. While this may result in a small error in measuring the actual skin temperature, it should be noted that this does not have an impact on the measurement of the electrolytes and metabolites due to the on board temperature calibration. To further ensure fidelity of sensor readings, the data collection of each channel took place when sufficient sweat sample was present, as evident by stabilization of sensor readings (within 10% variations between the continuous 5 data points) within the physiologically relevant range ([Na+]: 20-120 mM, [K+]: 2-16 mM, [glucose]: 0-200 μM, [lactate]: 2-30 mM).
Example on-body sweat analysis. The example on-body evaluation of the FISA was performed in compliance with the protocol that was approved by the institutional review board (IRB) at the University of California, Berkeley (CPHS 2014-08-6636). Twenty six healthy subjects (4 females and 22 males), aged 20-40, were recruited. The study was conducted as three trials: constant workload cycle ergometry, graded workload cycle ergometry, and outdoor running. Constant workload cycle ergometry was conducted on 14 volunteers (4 females and 10 males between the ages of 20 and 40). The graded cycle ergometry was conducted on 7 male volunteers (who were also involved in the constant workload cycle study). 12 male volunteers between the ages of 20 and 40 were recruited for outdoor running study. An electronically braked leg-cycle ergometer (Monark Ergomedic 839E, Monark Exercise AB, Vansbro, Switzerland) was used for cycling trials which included real-time monitoring of heart rate (HR), oxygen consumption (VO2), and pulmonary minute ventilation (VE). The power output (PO) was calibrated and monitored through the ergometer. HR was measured using a Tickr heart rate monitor (Wahoo fitness), and VO2 and VE were continuously recorded throughout trials via an open-circuit, automated, indirect calorimetry system (TrueOne metabolic system; ParvoMedics, Sandy, Utah).
The FISAs were packaged in traditional sweatbands during the indoor and outdoor trials. The sensor arrays were calibrated, and the subjects' foreheads and wrists were cleaned with alcohol swabs and gauze before sensors were worn on body. For the constant workload cycling trial subjects were cycling at 50 W with 50 W increments every 90 s to 150 W, and 20 minutes of cycling at 150 W. The PO was then decreased by 50 W every 90 s. The graded workload trial consisted of 5 minutes of seated rest followed by cycling at 75 W for 20 minutes and then cycling at 200 W until fatigue followed by a 10 minutes rest. The outdoor running trial was conducted with a group of 12 subjects in which six were instructed to drink 150 mL water every 5 minutes and six did not drink water throughout the trial. Subjects consented to run until volitional fatigue at a self-selected pace (˜12 km/h) and the Na+ and K+ sensors responses (from their foreheads) were recorded.
The wearable sensing platform described herein can be configured with additional or different features. As discussed herein, optionally the wearable sensing platform can simultaneously and selectively measure detailed profiles of Ca2+ and pH in real-time through a fully integrated wearable sensing system that can be worn during the course of normal daily activities.
Homeostasis of ionized calcium in biofluids is critical for human biological functions and organ systems. However, conventionally measurement of ionized calcium for clinical applications is not easily accessible due to its strict procedures and dependence on pH. Further, pH balance in body fluids greatly affects metabolic reactions and biological transport systems. A wearable electrochemical device is disclosed for monitoring (e.g., continuous monitoring) of ionized calcium and pH of body fluids using an array of Ca2+ and pH sensors (e.g., a disposable and flexible array of Ca2+ and pH sensors) that interfaces with a printed circuit board (e.g., flexible printed circuit board).
The disclosed platform enables real-time quantitative analysis of these sensing elements in body fluids such as sweat, urine, and tears. Accuracy of Ca2+ concentration and pH measured by the wearable sensors is validated through inductively coupled plasma-mass spectrometry technique and a commercial pH meter, respectively. Test results show that the wearable sensors have high repeatability and selectivity to the target ions. Real-time on-body assessment of sweat is also performed, and test results indicate that calcium concentration increases with decreasing pH. The disclosed platform can optionally be used in noninvasive continuous analysis of ionized calcium and pH in body fluids for disease diagnosis such as primary hyperparathyroidism and kidney stones.
Calcium is an essential component for human metabolism and minerals homeostasis. Indeed, about 1%-2% of human body weight is made up of calcium. Excessive alternation of ionized calcium levels in biofluids can have detrimental effects on the function and structure of many organs and systems in the human body, including myeloma, acid-base balance disorder, cirrhosis, renal failure, and normocalcaemic hyperparathyroidism. Free Ca2+ is conventionally measured in body fluids, such as urine for estimating kidney stone-forming salts. A person's pH can be another significant component for potential disease diagnosis.
For example, kidney stone patients with type II diabetes are reported to have a lower pH than normal individuals. Change in pH of skin, which is due to sweat, has been reported to take part in the development of skin disorders such as dermatitis, ichthyosis, and fungal infections. Additionally, free Ca2+ level in biofluids is dependent on pH. Therefore, rigorous processes and rapid analysis of Ca2+ with pH correction are conventionally performed in special laboratories within hours of samples extraction for accurate analysis of biofluids. Such applications can become easier by in situ measurement of Ca2+ and pH in body fluids through an in-depth data analysis performed using a reliable wearable sensing platform, such as that disclosed herein. However, conventionally, wearable Ca2+ sensors for real-time health assessment via body fluids has not been performed. On the other hand, careful analysis of the pH of body fluids is needed for more accurate in situ measurement. The use of flexible electronics and Ca2+ sensors having conformal contact with the human body, as disclosed herein, provides a more accurate and reliable epidermal quantitative analysis.
The disclosed integrated wearable sensing system performs real-time multiplexed sensing of human perspiration which enables accurate measurement of sweat analytes through signal processing and calibration. Considering the importance of Ca2+ and pH and their relationship in body fluids, it is desirable to simultaneously and selectively measure detailed profiles of Ca2+ and pH through an integrated wearable sensing platform during the course of normal daily activities with real-time feedback.
The disclosed wearable sensing system is configured to monitor in real-time Ca2+ concentration and pH of body fluids as well as skin temperature (see, e.g.,
Certain figures will now be summarized.
The example wearable sensing system includes an electrochemical platform comprising a Ca2+ sensor, a pH sensor, and/or a skin temperature sensor. The sensors may be plastic-based biosensors that are fabricated on a flexible polyethylene terephthalate (PET) substrate by common physical evaporation and electrochemical deposition methods as illustrated in
Surface membrane compositions of the electro-chemical electrodes are demonstrated in
Concentration of Ca2+ in human body fluids commonly varies from 0.5 to 3 mM. Due to the limited Ca2+ concentration range, sensitivity is important to ensure accurate measurements. ETH 129 is utilized as the Ca2+-selective ionophore due to its ability to translocate Ca2+ across biological membranes.
Dynamic response of the Ca2+ sensor under consecutive change from high to low and then to high Ca2+ concentrations is performed two times. Since Ca2+ is a divalent ion, the ideal sensitivity of an electrochemical Ca2+ sensor at standard temperature is 29.6 mV/decade of ion concentration, which is half of a monovalent ion, based upon the Nernst equation. The Ca2+ sensor shows a near-Nernstian response with an average of 32.7 mV/decade in two complete cycles. The senor shows fast response to changes in Ca2+ level with a 3.0% relative standard deviation (RSD) of sensitivity. This indicates that Ca2+ detection by the sensor is reproducible and durable under repetitive testing.
Body fluids contain a variety of electrolytes such as Ca2+, Mg2+, Na+, K+, H+, and NH4+. One desirable aspect of a wearable electrochemical sensor is its ability to selectively discriminate and measure target ions. Thus, the influence of these major electrolytes on sensor's performance is examined. In this study, interfering ions with physiological relevant concentrations (2 mM H+, 2 mM NH4+, 1 mM Mg2+, 8 mM K+, and 20 mM Na+) are subsequently added into 1 mM Ca2+ solution, and measurements are performed after 20 seconds waiting time. The change in potential due to addition of such ions, as demonstrated in
Additionally, it is beneficial for sensors to be reproducible such that reliable analysis can be attained from individual sensors. Six sample sensors were tested in a solution containing 0.125-2 mM of Ca2+ concentration range. As displayed in
These sensors show sensitivity ranging from 29.8 to 34.2 mV/decade of concentration, with an average sensitivity of 32.2 mV/decade and a RSD of 1.5%. The value of average sensitivity is used as a standard slope for calibration in later studies of biofluids. The variations in the absolute potentials of different sensors (resulted from sensor preparations and manually dropcasting method) are resolved by one-point calibration, as shown in inset of
This is similar to commercial pH meters where a standard solution is used to calibrate the measurement before actual measurement is undertaken. In the case of long-term analysis on Ca2+ concentration in body fluids, variation due to potential drift can easily conceal the actual measurement results. To test this, the Ca2+-selective sensor is kept under 0.25-1 mM Ca2+ solutions for a total of 90 minute and under 1 mM solution for 4 hours, as presented in
Similar to the Ca2+ sensors, aspects of the performance of the pH sensors are evaluated. PANI is one medium for pH measurement in body fluids due to its ease of fabrication, reproducibility, and biocompatibility. In this study, H+-selective PANI film is electrochemically deposited onto a Au electrode by cyclic voltammetry. The resulting PANI-based pH sensor presented in
The pH sensor is also selective to H+ with a potential variation of approximately 3.1% compared to its sensitivity as shown in
Conventional wearable pH sensors are not sufficiently accurate for detailed quantitative analysis in body fluids due to Cl− influence on solid-state Ag/AgCl RE. Here, pH sensing with Ag/AgCl and PVB-based Ag/AgCl REs is compared under a constant pH 5.0 with varying Cl− concentrations.
Skin temperature is an effective marker of the thermal state of individuals and is also informative for many skin related diseases (such as ulceration). Performance aspects of Cr/Au-based temperature sensors are discussed above. Such resistive temperature sensors have a sensitivity of 0.18% per ° C. with respect to its baseline resistance at room temperature, although other temperatures sensors with different sensitivity may be used. To investigate the influence of temperature on Ca2+ and pH sensors, sensors are tested in temperatures ranging from 23 to 37° C. in McIlvaine's buffer of pH 5.0 containing 0.5 mM Ca2+. Unlike enzymatic sensors, in which the performance is greatly influenced by the change in temperature, both Ca2+ and pH sensors show no significant response to temperature change as illustrated in
In this example, measurements of sweat and urine [Ca2+] acquired by sensors vary by a maximum of 7.0% and 10%, respectively, from the ICP-MS results. On the other hand, in this example pH sensors show <2.2% and 3.6% variations in sweat and urine from a commercial pH meter. These variations are relatively small, compared to normal range of [Ca2+] and pH of body fluids.
To further confirm the accuracy of sensor readings, additional studies were made by adding fixed amounts of Ca2+ and H+ into raw sweat, urine, and tear samples, and the change in potential with concentration was examined.
Following the ex situ analysis of sweat, urine, and tear, real-time on-body evaluation in human perspiration using the flexible integrated wearable device was also performed. As illustrated in
In this example, temperature then remains stable in the rest of the cycling time, as similarly discussed above.
These on-body results further affirm the utilization of the wearable system in personal health care. Such real-time continuous analysis can alert the wearer regarding excessive loss or rise of electrolytes.
Thus, a fully integrated wearable electrochemical platform for simultaneous in situ analysis of Ca2+ and pH in body fluids is disclosed. The wearable system, containing flexible sensors coupled with integrated circuits and a wireless transceiver, enables accurate measurements of characteristics of biofluids, including urine, tear, and sweat with real-time feedback. The disclosed wearable sensing systems offers many advantages over the traditional extensive laboratory analysis for accurate measurement of analytes in complex biofluids. The disclosed sensors' capabilities for long-term quantitative analysis and real-time on-body monitoring can also provide insightful information about Ca2+ and pH homeostasis in the human body. Owing to its miniaturization, system integration, and measurement simplification, the disclosed platform manifests a useful wearable sensing system that can be exploited for disease diagnosis where rapid analysis is desired for Ca2+ and pH in body fluids.
Example Materials. Calcium ionophore II (ETH 129), bis(2-ethylehexyl) sebacate (DOS), sodium tetrakis[3,5-bis(trifluoromethyl)phenyl] borate (Na-TFPB), high-molecular-weight polyvinyl chloride (PVC), [tetrahydrofuran (THF), polyvinyl butyral resin BUTVAR B-98 (PVB), sodium chloride (NaCl), 3,4-ethylenedioxythiophene (EDOT), poly-(sodium 4-styrenesulfonate) (NaPSS), aniline, and moisture-resistant 100 μm-thick PET.
Example Fabrication of Electrode Array. The fabrication process may be the same or similar to that discussed above. The PET may be cleaned with isopropyl alcohol and O2 plasma etching. An electrode array of 3.2 mm in diameter may be patterned via photolithography and may be thermally evaporated with 30/50 nm of Cr/Au, followed by lift-off in acetone. The electrode array may be additionally coated with 500 nm parylene C insulation layer (e.g., in a SCS Labcoter 2 Parylene Deposition System), and the 3 mm-diameter sensing electrode area may be defined via photolithography. The fabricated array may be further etched with O2 plasma to remove the parylene layer at the defined sensing area. Then, 200 nm Ag may be deposited via thermal evaporation and lift-off in acetone. It is understood that other processes, dimensions, and materials may be used to fabricate the electrode array.
Preparation of Ca2+ Selective Sensors and pH Sensors. Ca2+-selective cocktail was prepared by dissolving 100 mg of 33:0.5:65.45:1 wt % ratio of PVC:NaTFPB:DOS:ETH129 in 660 μL THF. The surface of the Ca2+-selective electrodes was modified by galvanostatic electrochemical polymerization of 0.01 M EDOT with 0.1 M NaPSS at a constant current of 2 mA·cm-2 to produce polymerization charges of 10 mC. Ten μL (1.4 μL·cm-2) of Ca2+-selective cocktail was then drop-casted onto a PEDOT:PSS coated electrode and left to dry overnight in a dark environment. Aniline was distilled at a vapor temperature of 100° C. and a pressure of 13 mmHg before usage. PANI was polymerized in a 0.1 M aniline/0.1 M HCl solution. Au surface was first modified by depositing Au (50 mM HAuCl4 and 50 mM HCl) for 30 s at 0 V, followed by PANI deposition using cyclic voltammetry from −0.2 to 1 V for 25 cycles at 100 mV/s. It is understood that other processes, dimensions, and materials may be used to prepare the sensors.
Evaluation of Ca2+ and pH Sensors General Performances. General performance of Ca2+ sensors was tested under a 0.01 M acetate buffer solution (pH 4.6) containing varying Ca2+ concentrations unless stated otherwise. Interference study was performed by subsequent addition of chloride solutions containing various cations (2 mM H+, 2 mM NH4 +, 1 mM Mg2+, 8 mM K+, and 20 mM Na+) into a 1 mM Ca2+ solution. pH sensors were tested using McIlvaine's buffer with varying pH to characterize general performances of the sensors.
Interference study was conducted by subsequent addition of chloride solutions containing 1 mM Ca2+, 1 mM NH4 +, 1 mM Mg2+, 8 mM K+, and 20 mM Na+ into a Mcllvaine's buffer solution of pH 4.0. All measurements were paused while changing solutions, and measurements were done after 20 s waiting period.
Ex Situ Evaluation of Body Fluids. Urine, sweat, and tear samples were collected from volunteer subjects for off-body evaluation.
Sweat and urine were initially tested with ICP-MS to measure [Ca2+], and the results were compared with the sensor readings of same sweat and urine samples. Sweat samples were diluted four times with deionized water for ex situ evaluations using ICP-MS and wearable sensors. The results were converted back in Table 2 to reflect raw sweat Ca2+ concentrations. [Ca2+] measured by the sensor was computed using a calibration curve. The calibration curve was obtained from artificial body fluids containing 50 mM NaCl and 4 mM KCl with 0.25, 0.5, and 1 mM CaCl2 in 0.01 M acetate buffer. pH of the samples was measured with a commercial pH meter (Horiba LAQUA Twin pH meter B-713) and PANI-based pH sensors. PANI-based pH sensor measurement was obtained by using similar methods as the [Ca2+] measurement. pH values were computed from a calibration curve obtained from solutions containing 50 mM NaCl and 4 mM KCl with McMaine buffer of pH varying from 4 to 7. To further confirm sensor readings, raw sweat, urine, and tear samples were subsequently added with a fixed amount of Ca2+, and initial [Ca2+] was back-calculated based on the change in potential with concentration.
The relationship between potential change and logarithmic concentration was analyzed by comparing with a standard calibration curve obtained from
In Situ Assessment of Sweat [Ca2+], pH, and Skin Temperature. On-body evaluation of sweat [Ca2+] and pH was performed in compliance with the protocol that was approved by the institutional review board at the University of California, Berkeley (CPHS 2014-08-6636). Five healthy male subjects, aged 20-30, were recruited. An electronically braked leg-cycle ergometer (Kettler E3 Upright Ergometer Exercise Bike) was used for stationary cycling trials. Subjects were told to bike for 30 minute at a constant workload cycle ergometry. Subject's forehead was wiped and cleaned with alcohol swab and gauze prior to wearing the sensor. Cycling protocol included a 5 minute ramp-up and a 20 minute biking at a power of 150 W, followed by a 5 minute cool-down session. Data are directly recorded in a mobile phone via a customized application. Sweat was simultaneously collected every 5 minute during cycling to compare on-body data with measurements from the ICP-MS and a pH meter. Collected sweat was diluted four times for ICP-MS measurements.
The wearable sensing platform described herein can be configured with still additional or different features. For example, the wearable sensing platform may be adapted for heavy metal monitoring of body fluids. An aspect of the disclosure relates to a flexible and wearable microsensor array for simultaneous multiplexed monitoring of heavy metals in human body fluids, such as, by way of example, Zn, Cd, Pb, Cu, and Hg ions. The target analytes may be detected, by way of example, via electrochemical square wave anodic stripping voltammetry (SWASV) on Au and Bi microelectrodes.
In an example process, the oxidation peaks of these metals are calibrated and compensated by incorporating a skin temperature sensor. The wearable sensing platform sensor arrays may provide high selectivity, repeatability, and flexibility. Urine samples are collected for heavy metal analysis, and measured results from the microsensors are validated through inductively coupled plasma mass spectrometry (ICP-MS). Real-time on-body evaluation of heavy metal (e.g., zinc and copper) levels in sweat of human subjects by cycling is performed to examine the change in concentrations with time. The wearable sensing platform is configured to provide insightful information about an individual's health state such as heavy metal exposure and aid the related clinical investigations.
By way of background, human body fluids are composed of various electrolytes, proteins, metabolites, as well as heavy metals. A variety of heavy metals can be found in human body fluids (such as blood, sweat, and urine) and are closely related to human health conditions. For example, Cu and Zn are essential trace elements that can have detrimental effects on an individual's health when there is an excess or deficiency. High copper accumulation in human body can lead to Wilson's disease, heart and kidney failure, liver damage, brain disease and disorder, and even death in extreme cases, whereas low levels of copper can cause anemia and osteoporosis. A lethal form of diarrhea and pneumonia can occur when a body has low zinc concentrations, whereas high levels of zinc can be toxic enough to cause liver damage, and even decrease cardiac functionality and pancreatic enzyme count in cases of prolonged exposure.
Additionally, cadmium, lead, and mercury exhibit toxic effects on human body systems including the nervous, immunological, and cardiovascular systems. High levels of cadmium exposure can lead to fatal respiratory tract, liver, and kidney problems. On the other hand, lead poisoning can slow down growth and cause other developmental delay as well as irritability, increased violent behavior, learning difficulties, fatigue, loss of appetite, and hearing loss for children and cause memory loss, infertility, high blood pressure, and decline in mental functioning for adults. Further, mercury poisoning leads to many diseases such as Hunter-Russell syndrome, Minamata disease, and acrodynia, to name a few. Therefore, determining one's exposure to such heavy metals can offer important insights into a person's health. Human sweat and urine are known to be the most important sources for detoxification of heavy metals; therefore, examination of sweat and urine heavy metals can assist toxicological and therapeutic studies.
Conventionally, detection of heavy metals in body fluids is challenging due to their extremely low concentrations (on the order of μg/L). Conventional heavy metals analysis procedures, including atomic absorption spectroscopy (AAS) or inductively coupled plasma mass spectrometry (ICP-MS), rely on inconvenient, bulky and expensive analytical instruments. Further, such instruments may be unavailable in many regions of the world.
The effective preconcentration/deposition step and advanced electrochemical measurements of the accumulated analytes make anodic stripping analysis a highly sensitive and effective electroanalytical technique. The stripping-voltammetric measurements of trace metals in sweat have been reported using collected sweat samples. Disadvantageously, such use of collected sweat may be subject to inaccuracy due to sample contamination and sweat evaporation. In addition, such techniques do not yield real-time information on dynamic events. Given the importance or toxicity of a variety of heavy metals to the individual's health states, it is very attractive to perform simultaneous multiplexed screening of heavy metals in sweat and do proper signal calibrations to ensure accurate measurements.
Disclosed herein is a flexible multiplexed trace metals monitoring device to extract useful information on heavy metal levels in body fluids such as sweat and urine. The device can also be used as a wearable device for real-time monitoring of the heavy metals in human sweat. A microsensor array is utilized to simultaneously and selectively measure multiple heavy metals (e.g., Zn, Cd, Pb, Cu, and/or Hg) using, by way of example, square wave anodic stripping voltammetry (SWASV), as well as skin temperature to calibrate heavy metal sensors' readings in real-time (
Furthermore, on-body measurements of sweat trace metals during exercise are performed by implementing the integrated sensors directly on human skin. Such real-time assessment of heavy metals in sweat can give early warnings of heavy metal exposure. As similarly discussed elsewhere herein (see, e.g.,
The working electrode is selected to provide for successful stripping analysis. The ideal material for working electrode should offer an effective preconcentration, a favorable redox reaction of the target metal, reproducible and renewable surface, and a low background current over a wide potential range. Although mercury has been the most explored electrode for many stripping applications, it is not desired for wearable biosensors given its toxicity and volatility. However, bismuth and gold electrodes, by way of example, provide good stripping voltammetric performance and biocompatibility and so are employed in developing the disclosed wearable biosensors for heavy metals analysis.
The microsensors arrays are optionally fabricated on a flexible polyethylene terephthalate (PET) substrate through a procedure involving multiple steps of photolithography, evaporation (Cr/Au, Ag, Bi), and lift-off as illustrated in
As aforementioned, Au microelectrodes offer excellent biocompatibility and a wide operational potential window owing to their high stability. Au is an excellent electrode material for Pb, Cu, and Hg stripping, although other materials may be used. In order to investigate the relationship between the concentration of heavy metals and the response of SWASV, the voltammograms are recorded using a 0.01 M acetate buffer solution (pH 4.6) containing 50 mM NaCl (to mimic human sweat) with an addition of 50-100 μg/L heavy metals after every trial. As illustrated in
A linear relationship between the peak height (current amplitude measured from the baseline as illustrated in
The selectivity of the Bi and Au based microsensors are advantageous for the analysis in biofluids. Given the relatively high concentration of sweat Cu and Zn (on the order of hundreds μg/L), an interference study on Au and Bi based microsensors is implemented by varying Cu and Zn concentrations, respectively. As illustrated in
Another significant aspect is the influence of temperature on the responses of biosensors. Skin temperature and environmental temperature can have direct influence on metabolite sensors; hence, temperature compensation advantageously ensures accurate readings of the sensors. To this end, SWASV responses of the microsensors are investigated by gradually increasing the temperature of the solution containing Cu(II) and Zn(II) from 20 to 40° C. As shown in
The integration of a temperature sensor into the microsensor array enables real-time temperature compensation to ensure an accurate and a reliable heavy metal detection. The repeatability of Au and Bi based microsensors for stripping analysis of Cu and Zn was examined by recording the stripping voltammograms under the same condition mentioned above in a solution containing 150 μg/L Cu and Zn, respectively.
In addition, the reproducibility of different microsensor arrays for heavy metal analysis was examined. As demonstrated in
The wearable microsensor array preferably are able to withstand mechanical deformation during vigorous physical exercise. The flexibility was investigated by monitoring the peak heights of stripping performance of the sensor array after mechanical bending (radii of curvature is 3.2 mm) (
A stable and a reliable performance of biosensors in biofluids is desirable for practical usage. To demonstrate such capability of the microsensor array (not only for on-body real-time monitoring but also for off-body analysis of different biofluids), sweat and urine samples are collected from volunteer subjects for off-body measurements. The physiological levels of heavy metals in human sweat and urine are relatively low (<1 mg/L). Specifically, human sweat contains 100-1000 μg/L of Zn and Cu while the concentrations of Pb, Cd, and Hg usually fall below 100 μg/L. Because of the relatively high concentrations of free Cu and Zn ions, off-body measurements showed visible oxidation peaks in sweat and urine samples of all the subjects. For Cu stripping, 15 seconds deposition time was used on the Au microelectrodes to minimize peak distortion and electrode fouling. In addition, the permselective/protective Nafion coating is found to be beneficial in addressing the challenge of biofouling due to the surface-active compounds in complex human biofluids. It helps to enhance oxidation peaks of targeted trace metals and allows direct detection in human sweat and urine samples.
Cu and Zn levels in sweat (
On-body heavy metals monitoring was performed during a constant-load exercise on a cycle ergometer. The protocol involved a 5 minute ramp-up, 30 minute cycling at 150 W, and a 5 minute cool-down. The microsensor arrays were packaged in a wristband (
Thus, a wearable and flexible microsensor array is disclosed that can perform simultaneous and selective detection of multiple heavy metals (e.g., Zn, Cd, Pb, Cu, and Hg) noninvasively. The flexible microsensor arrays display very good repeatability and stability for heavy metal analysis. A temperature sensor is utilized for real-time compensation of the signals to ensure accurate and reliable measurements. The microsensor array has been successfully used to accurately and selectively monitor heavy metal levels in human body fluids such as sweat and urine. The disclosed microsensor array device greatly expands the panel of analytes for noninvasive wearable biosensing. For example, the microsensor array device may be used to monitor heavy metal exposure and aid in related clinical investigations.
Example Materials. Moisture-resistant polyethylene terephthalate (PET), 100 μm thick, zinc, cadmium, lead, copper, and mercury standard AAS solutions (1000 mg/L in nitric acid), acetate buffer, sodium chloride (NaCl), and Nafion 117 solution (5 wt %).
Fabrication of Electrode Arrays. An example fabrication process of the electrode arrays is illustrated in
Characterizations of the Microsensor Arrays. Square wave anodic stripping voltammetry (SWASV) was employed to characterize the electrochemical stripping of heavy metals on the microsensor arrays. In order to evaluate the performance of Bi electrodes, a deposition potential of −1.5 V (vs Ag+/Ag) was applied for 180 s, followed by a SWASV scan to a final potential of −0.5 V (vs Ag+/Ag) at a frequency of 60 Hz, an amplitude of 40 mV, and a potential step of 4 mV in 0.01 M acetate buffer (pH 4.6) containing 50 mM NaCl. In order to evaluate the performance of Au electrodes, a deposition potential of −0.7 V (vs Ag+/Ag) was applied for 120 s (15 s for repeatability tests and detection in biofluids), followed by a SWASV scan to a final potential of 0.8 V (vs Ag+/Ag) at a frequency of 60 Hz, an amplitude of 40 mV, and a potential step of 4 mV in 0.01 M acetate buffer (pH 4.6) containing 50 mM NaCl. Mechanical deformation was tested by repeatedly bending (radii of curvature is 3.2 mm) the microelectrodes array for 200 times.
Off-Body Calibration and Validation in Human Sweat and Urine. Sweat samples were collected directly from the forehead and the arm of volunteer subjects during their constant load (150 W) cycling exercise. The subjects' skin was cleaned with alcohol swabs and gauze before the exercise and after every sweat collection. Urine samples were collected from the same volunteer subjects. And 50 μL human sweat and urine samples were used for the off-body measurement. The level of heavy metals was estimated by SWASV through standard addition (1-2 μL each time) of 100 mg/L Zn or Cu standard solutions. It should be noted that, in some cases, the peak positions slightly shifted after the standard addition due to the greatly changed heavy metal concentrations. The measurement results of biofluid heavy metal levels from the microsensors were compared with the data measured directly through an ICP Optima 7000 DV instrument.
Real Time On-Body Heavy Metal Analysis. On-body evaluation of the microsensor arrays was performed in compliance with the protocol that was approved by the institutional review board at the University of California, Berkeley (CPHS 2014-08-6636). Five healthy subjects (all males), aged 20-40, were used. An electronically braked leg-cycle ergometer (Kettler E3 Upright Ergometer Exercise Bike) was used for cycling trials. The subjects' skin was cleaned with alcohol swabs and gauze before sensors were worn on-body. A constant workload cycling regimen was used in which subjects were cycling at 50 W with 50 W increments every 150 s up to 150 W, and then cycling at 150 W for 30 min. The 5 minute cool down section involved cycling with decreased power output by 50 W every 150 s. During the exercise, on-body analysis was recorded using a Gamry electrochemical potentiostat (PCI4/G300). The on-body measurement results were calibrated using the measured skin/environment temperature at the same time. Such calibration eliminated the errors of stripping signals resulted from temperature variations. The heavy metal concentrations from on-body tests were roughly estimated using a coefficient factor obtained from an off-body standard addition method shown in
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The methods and processes described herein may have fewer or additional steps or states and the steps or states may be performed in a different order. Not all steps or states need to be reached. The methods and processes described herein may be embodied in, and fully or partially automated via, software code modules executed by one or more general purpose computers, microcontrollers, and/or other processing devices. The code modules may be stored in any type of computer-readable medium or other computer storage device. Some or all of the methods may alternatively be embodied in whole or in part in specialized computer hardware. The systems described herein may optionally include displays, user input devices (e.g., touchscreen, keyboard, mouse, voice recognition, etc.), network interfaces, etc.
The results of the disclosed methods may be stored in any type of computer data repository, such as relational databases and flat file systems that use volatile and/or non-volatile memory (e.g., magnetic disk storage, optical storage, EEPROM and/or solid state RAM).
The various illustrative logical blocks, modules, routines, and algorithm steps described in connection with the embodiments disclosed herein can be implemented as electronic hardware, computer software, or combinations of both. To clearly illustrate this interchangeability of hardware and software, various illustrative components, blocks, modules, and steps have been described above generally in terms of their functionality. Whether such functionality is implemented as hardware or software depends upon the particular application and design constraints imposed on the overall system. The described functionality can be implemented in varying ways for each particular application, but such implementation decisions should not be interpreted as causing a departure from the scope of the disclosure.
Moreover, the various illustrative logical blocks and modules described in connection with the embodiments disclosed herein can be implemented or performed by a machine, such as a general purpose processor device, a digital signal processor (DSP), an application specific integrated circuit (ASIC), a field programmable gate array (FPGA) or other programmable logic device, discrete gate or transistor logic, discrete hardware components, or any combination thereof designed to perform the functions described herein. A general purpose processor device can be a microprocessor, but in the alternative, the processor device can be a controller, microcontroller, or state machine, combinations of the same, or the like. A processor device can include electrical circuitry configured to process computer-executable instructions. In another embodiment, a processor device includes an FPGA or other programmable device that performs logic operations without processing computer-executable instructions. A processor device can also be implemented as a combination of computing devices, e.g., a combination of a DSP and a microprocessor, a plurality of microprocessors, one or more microprocessors in conjunction with a DSP core, or any other such configuration. Although described herein primarily with respect to digital technology, a processor device may also include primarily analog components. A computing environment can include any type of computer system, including, but not limited to, a computer system based on a microprocessor, a mainframe computer, a digital signal processor, a portable computing device, a device controller, or a computational engine within an appliance, to name a few.
The elements of a method, process, routine, or algorithm described in connection with the embodiments disclosed herein can be embodied directly in hardware, in a software module executed by a processor device, or in a combination of the two. A software module can reside in RAM memory, flash memory, ROM memory, EPROM memory, EEPROM memory, registers, hard disk, a removable disk, a CD-ROM, or any other form of a non-transitory computer-readable storage medium. An exemplary storage medium can be coupled to the processor device such that the processor device can read information from, and write information to, the storage medium. In the alternative, the storage medium can be integral to the processor device. The processor device and the storage medium can reside in an ASIC. The ASIC can reside in a user terminal. In the alternative, the processor device and the storage medium can reside as discrete components in a user terminal.
Conditional language used herein, such as, among others, “can,” “may,” “might,” “may,” “e.g.,” and the like, unless specifically stated otherwise, or otherwise understood within the context as used, is generally intended to convey that certain embodiments include, while other embodiments do not include, certain features, elements and/or steps. Thus, such conditional language is not generally intended to imply that features, elements and/or steps are in any way required for one or more embodiments or that one or more embodiments necessarily include logic for deciding, with or without other input or prompting, whether these features, elements and/or steps are included or are to be performed in any particular embodiment. The terms “comprising,” “including,” “having,” and the like are synonymous and are used inclusively, in an open-ended fashion, and do not exclude additional elements, features, acts, operations, and so forth. Also, the term “or” is used in its inclusive sense (and not in its exclusive sense) so that when used, for example, to connect a list of elements, the term “or” means one, some, or all of the elements in the list.
Disjunctive language such as the phrase “at least one of X, Y, Z,” unless specifically stated otherwise, is otherwise understood with the context as used in general to present that an item, term, etc., may be either X, Y, or Z, or any combination thereof (e.g., X, Y, and/or Z). Thus, such disjunctive language is not generally intended to, and should not, imply that certain embodiments require at least one of X, at least one of Y, or at least one of Z to each be present.
While the phrase “click” may be used with respect to a user selecting a control, menu selection, or the like, other user inputs may be used, such as voice commands, text entry, gestures, etc. User inputs may, by way of example, be provided via an interface, such as via text fields, wherein a user enters text, and/or via a menu selection (e.g., a drop down menu, a list or other arrangement via which the user can check via a check box or otherwise make a selection or selections, a group of individually selectable icons, etc.). When the user provides an input or activates a control, a corresponding computing system may perform the corresponding operation. Some or all of the data, inputs and instructions provided by a user may optionally be stored in a system data store (e.g., a database), from which the system may access and retrieve such data, inputs, and instructions. The notifications and user interfaces described herein may be provided via a Web page, a dedicated or non-dedicated phone application, computer application, a short messaging service message (e.g., SMS, MMS, etc.), instant messaging, email, push notification, audibly, and/or otherwise.
The user terminals described herein may be in the form of a mobile communication device (e.g., a cell phone), laptop, tablet computer, interactive television, game console, media streaming device, head-wearable display, networked watch, etc. The user terminals may optionally include displays, user input devices (e.g., touchscreen, keyboard, mouse, voice recognition, etc.), network interfaces, etc.
While the above detailed description has shown, described, and pointed out novel features as applied to various embodiments, it can be understood that various omissions, substitutions, and changes in the form and details of the devices or algorithms illustrated can be made without departing from the spirit of the disclosure. As can be recognized, certain embodiments described herein can be embodied within a form that does not provide all of the features and benefits set forth herein, as some features can be used or practiced separately from others. The scope of certain embodiments disclosed herein is indicated by the appended claims rather than by the foregoing description. All changes which come within the meaning and range of equivalency of the claims are to be embraced within their scope.
This invention was made with government support under Number P01 HG000205 awarded by the National Institute of Health. The government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US16/53988 | 9/27/2016 | WO | 00 |
| Number | Date | Country | |
|---|---|---|---|
| 62233955 | Sep 2015 | US |
