Electrochemical analysis of sweat using soft bioelectronics on human skin provides a new route for noninvasive glucose monitoring without painful blood collection. However, sweat-based glucose sensing still faces many challenges, such as difficulty in sweat collection, activity variation of glucose oxidase due to lactic acid secretion and ambient temperature changes, and delamination of the enzyme when exposed to mechanical friction and skin deformation. Precise point-of-care therapy in response to the measured glucose levels is still very challenging. We present a wearable/disposable sweat-based glucose monitoring device integrated with a feedback transdermal drug delivery module. Careful multilayer patch design and miniaturization of sensors increase the efficiency of the sweat collection and sensing process. Multimodal glucose sensing, as well as its real-time correction based on pH, temperature, and humidity measurements, maximizes the accuracy of the sensing. The minimal layout design of the same sensors also enables a strip-type disposable device. Drugs for the feedback transdermal therapy are loaded on two different temperature-responsive phase change nanoparticles. These nanoparticles are embedded in hyaluronic acid hydrogel microneedles, which are additionally coated with phase change materials. This enables multistage, spatially patterned, and precisely controlled drug release in response to the patient’s glucose level. The system provides a novel closed-loop solution for the noninvasive sweat-based management of diabetes mellitus.

Keywords

Various kinds of flexible and stretchable devices based on an ultrathin ( 16 – 18 ) and stretchable design ( 19 , 20 ) have been developed for monitoring individual health status and delivering the corresponding feedback therapy ( 21 , 22 ). Here, we develop a patch-based wearable/strip-type disposable system for noninvasive sweat glucose monitoring and microneedle-based point-of-care therapy ( 23 ). Key novel advantages of this system include (i) mass production–compatible porous metal-based electrodes and fabrication processes, (ii) miniaturized sensor design that allows for reliable sweat analysis even with an infinitesimal amount of sweat, (iii) patch- and disposable-type design for enhancing practical applicability, (iv) multiple sweat control and uptake layers for efficient sweat collection, (v) a porous gold nanostructure for maximizing the electrochemically active surface area for detection of a small amount of glucose in sweat with high sensitivity, (vi) multiple glucose sensing devices for enhanced accuracy, and (vii) multistage and precisely controlled transdermal drug [metformin or chlorpropamide (type 2 diabetes drugs)] delivery through biocompatible hyaluronic acid hydrogel microneedles containing drug-loaded phase change nanoparticles (PCNs). Metformin is one of the first-line drugs for treating type 2 diabetes ( 24 ). Because drug delivery through the skin can bypass the digestive system, transdermal delivery of metformin requires a lower dosage of drugs than oral delivery and prevents gastrointestinal side effects ( 23 ). This novel system for high-fidelity sweat glucose measurement and feedback-controlled drug delivery enables efficient management of blood glucose concentration without pain and stress.

Extensive efforts have been made to develop noninvasive sweat-based biomarker monitoring methods. Wearable biosensors enable continuous monitoring of metabolites [including glucose ( 3 – 5 ), lactate ( 8 ), and alcohol ( 9 )] and electrolytes [including potassium ( 3 ), calcium ( 10 ), and heavy metal ions ( 11 )] in sweat. The estimation of the blood glucose concentration based on the sweat-based glucose measurement is a potential solution ( 12 , 13 ). However, many challenges still exist for the accurate sweat-based glucose measurement ( 14 , 15 ). For example, the sweat collection procedure is tedious, and the sweat collection times vary depending on environmental conditions. In addition to the difficulty of measuring glucose levels in sweat because of its much smaller concentration than that in blood, lactic acid in sweat, ambient temperature changes, and various medications can induce errors in enzyme-based glucose sensing. Mechanical friction and deformation of devices on soft human skin can delaminate the enzyme from the glucose sensor and cause mechanical fractures in devices. A closed-loop system with the feedback delivery of a precisely controlled amount of drugs in response to the glucose monitoring result is another important unaccomplished goal for maintaining homeostasis ( 6 , 7 ).

Diabetes is one of the most prevalent chronic diseases, causing uncontrollable blood glucose levels ( 1 ). Patients with diabetes are advised to check their blood glucose level daily and to take periodic insulin shots for continuous management of their blood glucose level ( 2 ). However, patients often do not follow this recommendation because of the pain and accompanying intense stress of repetitive blood collection and insulin shots. This often leads to various severe diabetic complications, such as cardiovascular and kidney diseases, stroke, blindness, and nerve degeneration. In addition, insulin overtreatment causes an abrupt drop in the blood glucose concentration, which may cause seizures, unconsciousness, and even death. Therefore, a novel method for painless and stress-free glucose monitoring ( 3 – 5 ) and precise maintenance of homeostasis through controlled drug delivery ( 6 , 7 ) is highly desirable.

RESULTS

System design of the wearable diabetes patch and disposable strip The patch-based wearable devices in Fig. 1A have an ultrathin and stretchable design (fig. S1A), which enables conformal contact with the skin for efficient sweat collection and high performance under physical deformation. The same sensors can be further miniaturized and fabricated as a disposable strip (Fig. 1B). Depending on the sweat glucose concentration, an appropriate amount of metformin (or chlorpropamide) is transdermally delivered through microneedles. The drug delivery can be thermally controlled in a multistage manner. The microneedles assembled on the multichannel thermal actuator can be periodically replaced with new ones (Fig. 1C and fig. S1, B and C). These patch-type devices are fabricated on a handle substrate and then transfer-printed onto a thin silicone patch (fig. S2). The strip-type sensor is fabricated on a thin polyimide (PI) substrate. Patterned electrochemical functionalization completes the fabrication. Fig. 1 Wearable/disposable sweat monitoring device and microneedle-based transdermal drug delivery module. (A) Optical camera image (top; dotted line, edges of the patch) and schematic (bottom) of the wearable sweat monitoring patch. A porous sweat-uptake layer is placed on a Nafion layer and sensors. (B) Optical camera image (top) and schematic (bottom) of the disposable sweat monitoring strip. (C) Optical camera image (top; dotted line, edges of the patch) and schematic (bottom) of the transdermal drug delivery device. Replacement-type microneedles are assembled on a three-channel thermal actuator. (D) Schematic drawing of the glucose sensor in an exploded view. PB, prussian blue. (E) Minimum volume of the artificial sweat required for sensing with different sizes of the glucose sensor. (F) Scanning electron microscope (SEM) images before (left) and after (right) immobilization of the enzyme (enz.) on the porous gold electrode. (G) Comparison of the H 2 O 2 sensitivity in the planar and porous gold electrode deposited with Prussian blue at different H 2 O 2 concentrations. (H) Schematic of the drug-loaded microneedles. The right inset describes details of different PCNs. (I) SEM image of the microneedles. (J) Confocal microscope images of the released dye from microneedles (MN) (top view) into the 4% agarose gel (green, agar; red, dye). (K) Infrared (IR) camera image of the three-channel (ch) thermal actuator. The system operation sequence is schematically described in fig. S3. After wearing the patch, sweat accumulates in the porous sweat-uptake layer. A waterproof band behind the silicone patch aids in sweat collection and prevents delamination of the patch from the skin. A porous and negatively charged Nafion layer between the sensors and the sweat-uptake layer helps in the immobilization of the enzyme (25) and screens out negatively charged molecules that may affect the glucose sensing (for example, drug molecules contained in sweat, such as acetaminophen or acetylsalicylic acid) (25, 26). A humidity sensor monitors the critical amount of sweat for reliable glucose sensing by measuring impedance change by sweat generation. Above the critical humidity, the glucose, pH, and temperature sensors begin taking measurements to determine correlated blood glucose level. The pH and temperature sensors correct potential errors of the enzyme-based glucose level measurement in real time. Upon hyperglycemia, thermal actuation controls the feedback transdermal delivery of metformin loaded in PCNs.

Materials and key device design A series of stretchable sensors (humidity, glucose, pH, and temperature) are monolithically integrated (Fig. 1A) for efficient sweat-based glucose sensing. Multipoint sensing for glucose (triple) and pH (quadruple) improves the detection accuracy (fig. S1A). In the glucose and pH sensors, the reference and counter electrodes are designed to be packed as closely as possible to minimize the required amount of sweat (Fig. 1D). As the working electrode diameter (D work ) decreases, the required sweat amount can be reduced to as small as 1 μl (Fig. 1E), which is a 20-fold decrease from that previously reported by Lee et al. (5). The working electrode consists of porous gold formed by electrodeposition and enzymes [glucose oxidase (GOx)] drop-casted on it and then covered by Nafion and sequentially cross-linked by glutaraldehyde (Fig. 1F). The porous structure allows for a larger electrochemically active surface area (27, 28) and stronger enzyme immobilization (29, 30). Cyclic voltammograms (fig. S4A) and ac impedance measurements (fig. S4, B and C) using the Fe(CN) 6 3−/4− redox couple show a high charge storage capacity and low interfacial impedance in the porous structure. The high H 2 O 2 reducing catalytic activity of the porous structure (Fig. 1G and fig. S4D) confirms the enhanced sensitivity. Robustly cross-linked enzymes on the porous metal structure also enhance the reliability of the sensor under mechanical friction and deformation. Film-type microneedles (fig. S1B) are integrated on top of the stretchable heater (fig. S1C), and these needles can be periodically replaced (Fig. 1C). For the multistep and precisely controlled drug delivery, two types of metformin-loaded PCNs [PCN1, melting transition temperature (T m,1 ) = 38°C; PCN2, T m,2 = 43°C] are embedded in the hyaluronic acid hydrogel microneedles (Fig. 1H). The hyaluronic acid hydrogel is a widely used biocompatible material (31). An additional coating with phase change materials (PCMs) prevents unwanted dissolution of the hyaluronic acid matrix in contact with interstitial fluids (Fig. 1H). Controlled thermal actuation by the integrated heater activates either PCN1 alone or both PCNs (fig. S5). An SEM image of the fabricated microneedles is shown in Fig. 1I. Confocal microscope images show the dye release from microneedles into 4% agarose gel (tissue-like gel) (Fig. 1J) (32). The heater is designed with three channels (Fig. 1K) and triggers the multistage drug delivery that can be tuned with thermal patterns.

Sweat control optimization Efficient use of the generated sweat is an important issue in the sweat-based monitoring system, which is enabled by the sweat-uptake and waterproof layers. A porous and hydrophilic sweat-uptake layer (Fig. 2A, inset) (33, 34), which is placed on the Nafion layer and sensors, absorbs sweat and delivers it to the sensors through the Nafion layer (fig. S6A). The waterproof band added on the elastomeric silicone patch covers the outside to separate the sweat from external humidity (Fig. 2A), suppresses sweat evaporation (fig. S6B), and prevents delamination of the patch during skin deformation (Fig. 2B). The strip-type sensor absorbs sweat by the capillary force formed by the gap (fluidic channel) between the sweat-drawing layer and substrate (Figs. 1B, bottom inset, and 2C). The miniaturized sensor design (Fig. 1, D and E) and the fast collection of sweat enable rapid sensing. The miniaturized glucose sensor can be stably operated with only 1 μl of sweat (Fig. 2D and fig. S6, C and D). Fig. 2 Optimization of the sweat control and characterization of individual sensors. (A) Optical image of the wearable sweat analysis patch with a sweat-uptake layer and a waterproof band. The inset shows the magnified view of the porous sweat-uptake layer. (B) Optical image of the sweat analysis patch under deformation. (C) Optical image of the disposable sweat analysis strip on human skin with perspiration. (D) Glucose (glu.) concentration measurement at different sweat volumes (0.3 mM glucose in artificial sweat). (E) Optical image (top) and calibration curve (bottom) of the humidity sensor. Inset shows the image before and after wetting of the sensor. (F) Optical image (top) and calibration curve (bottom) of the glucose sensor. (G) Optical image (top) and calibration curve (bottom) of the pH sensor. (H) Optical image (top) and calibration curve (bottom) of the temperature sensor. (I) Changes of the relative sensitivity of the uncorrected glucose sensor at different pH levels. The relative sensitivity (S/S o ) is defined as measured sensitivity divided by sensitivity at pH 5. (J) In vitro monitoring of glucose changes with (green) and without (red) correction using simultaneous pH measurements (blue). (K) Calibration curves of the glucose sensor at different temperatures.

Electrochemical and electrical characterization of individual sensors The humidity sensor monitors the sweat collection procedure through impedance changes of poly(3,4-ethylenedioxythiophene) (PEDOT) interdigitated electrodes to estimate an appropriate starting point of sweat analysis (Fig. 2E). The glucose sensor based on a Prussian blue–deposited porous gold electrode is calibrated to the glucose concentration range between 10 μM and 1 mM (typical glucose concentrations in human sweat) by the galvanostatic method (Fig. 2F and fig. S7, A and B). These glucose concentrations correspond to typical sweat glucose concentrations of both hypoglycemic and hyperglycemic patients as well as healthy people (12, 13). The glucose sensor is stable under mechanical deformation (fig. S7C) and selectively detects glucose in the presence of other biomolecules [including ascorbic acid, uric acid, and lactic acid (3, 35)] and drugs [including acetaminophen, acetylsalicylic acid, and metformin (fig. S7D) (36–38)], whose concentrations are in values commonly found in the human sweat. The glucose sensor works for one full day without additional calibration under ambient conditions and for several days depending on storage conditions (fig. S7E). The enzyme-based glucose sensor shows deviations at different pH levels (fig. S7F), and thus simultaneous monitoring of pH levels is important. The pH sensor measures changes in the open circuit potential (OCP) between a polyaniline (PANi)–deposited working electrode and an Ag/AgCl electrode. The pH sensor is calibrated using standard pH buffer solutions (Fig. 2G and fig. S8A) and maintains stable operation under deformation (fig. S8B). Because the dependence of measured OCP and the pH was not linear, the calibration curves were obtained for every pH difference of 0.5. The pH sensor is reliable over repeated use at different pH levels (fig. S8C) and different temperatures (fig. S8D). The resistor-based temperature sensor is also calibrated for skin temperature monitoring (Fig. 2H). The simultaneous use of these co-integrated sensors can enhance the accuracy of glucose sensing. For example, metabolic secretion of lactic acid in sweat lowers the pH level to 4 to 5. The simultaneous pH sensing can correct this pH-dependent deviation (Fig. 2I and fig. S7F) of the enzyme-based glucose sensor. Simulated hyperglycemia—0.3 mM sweat glucose—is monitored in vitro under two different pH variation settings (5→4→5 and 5→6→5; Fig. 2J, left and right). In both cases, real-time correction increases glucose measurement accuracy. The glucose sensor with the current device setting does not show significant deviations at different temperatures (Fig. 2K).

Thermoresponsive PCNs PCMs have melting temperatures that are low but above human body temperature, and thus they are widely used in the thermoresponsive drug delivery (39). PCNs are made of PCMs, drugs, and biocompatible ligands (Fig. 3A). Different PCNs enable temperature-dependent stepwise drug delivery. We use palm oil (PCN1, T m = 38°C) and tridecanoic acid (PCN2, T m = 43°C), which melt above the skin temperature (30°C; fig. S9A). The drugs (metformin or chlorpropamide) are embedded in the PCM matrix (Fig. 3A and fig. S9, B to D). Pluronic F-127 and 3,4-dihydroxyl-l-phenylalanine (DOPA)–conjugated hyaluronic acid (fig. S10, A and B) are used as ligands, which make an oil-in-water emulsion (fig. S11A). Cryogenic transmission electron microscopy (cryo-TEM) images show that PCNs below the melting temperature are solid (Fig. 3, B and C, left). Above the melting temperature, PCNs change to liquid and aggregate (Fig. 3, B and C, right). Cytotoxicity tests show that both PCNs are nontoxic and suitable for biomedical applications (fig. S11B). The hydrodynamic diameters and negative zeta potentials of the PCNs slightly increase and decrease, respectively, as the temperature increases from the skin temperature (30°C; fig. S9A) to above melting temperatures (40° and 45°C; Fig. 3D and fig. S11C). The PCM matrices can block the drug release before thermal actuation. Their stepwise temperature-dependent melting controls the amount of the drug release. When the temperature reaches 40°C, only the drugs contained in PCN1 are released, whereas at 45°C, the drugs in both PCN1 and PCN2 are released (fig. S11D). Fig. 3 Characterization of PCNs and PCN-loaded microneedles. (A) Schematic illustration of the PCN. HA, hyaluronic acid. (B) Cryo-TEM image of PCN1 (palm oil) below the melting temperature (left) and TEM image of PCN1 above the melting temperature (right). (C) Cryo-TEM image of PCN2 (tridecanoic acid) below the melting temperature (left) and TEM image of PCN2 above the melting temperature (right). (D) Dynamic light scattering size measurement of PCNs at 30° (skin temperature), 40°, and 45°C. (E) Optical image (left) and confocal fluorescence (FL) microscope image (right) of PCM-coated microneedles. Microneedles contain dye-loaded PCNs for imaging. (F) Microneedle dissolution test in PBS before (left) and after (right) the PCM coating. (G) Mechanical strength test of microneedles in their dry and wet states. (H) Confocal microscope image of microneedles penetrating into 4% agarose gel (left) and IR camera image of the thermal actuation on the gel (right). (I) Drug-release profile from microneedles. (J) IR camera images of eight different spatiothermal profiles using the three-channel thermal actuator for multistage drug delivery. (K) Multistage drug-release profile. a.u., arbitrary units.

Microneedle fabrication and multistage drug delivery The microneedles are fabricated by molding the hyaluronic acid hydrogel matrix containing drug-loaded PCNs, followed by an additional PCM spray coating (Fig. 3E, left, and fig. S12). A confocal laser fluorescence microscope image shows that the dye-loaded PCNs are well embedded in the microneedles (Fig. 3E, right). Hydrogel-based microneedles dissolve when they come in contact with body fluids (Fig. 3F, left), whereas the PCM coating prevents the dissolution before the controlled melting of the PCM (Fig. 3F, right). The microneedles should be stiff enough to penetrate into the skin (40); compression tests confirm their mechanical strength (Fig. 3G). The microneedles successfully penetrate into 4% agarose gel (tissue-like gel) and generate pores (Fig. 3H, left). After poration by microneedles, vertical heat transfer dissolves the outside PCM coating and embedded PCNs. Although accidental peel-offs create small defects in the PCM coating, the drug contained in the PCNs is not released. The vertical temperature distribution of the agarose gel imaged by an IR camera confirms the successful heat transfer (Fig. 3H, right). To investigate the stepwise drug release, we heated the microneedles containing dyes from 25° to 45°C (Fig. 3I). There is a negligible release under and around the skin temperature (30°C) due to the PCM coating and PCNs. At the elevated temperature (40°C), the PCM coating on microneedles and PCN1 are dissolved, and metformin in PCN1 is released. At the higher temperature (45°C), metformin in both PCN1 and PCN2 is released. Spatial patterning ofthe embedded heater subdivides the release steps further, which isprecisely controlled (fig. S13A) by co-integrated temperature sensors (fig. S13, B and C). Heaters consisting of the three channels together with the two types of PCNs result in eight different spatiothermal patterns (Fig. 3J) and six-stage programmed dye-release profiles (Fig. 3K).

Sweat-based glucose monitoring on human subjects Wearable/disposable sweat-based glucose sensors are used for human sweat analysis. The wearable patch is connected to a portable electrochemical analyzer (Fig. 4A and fig. S14). The monitoring starts with humidity sensing to determine the optimum point to start sweat analysis (Fig. 4B). When the sweat-uptake layer absorbs a sufficient amount of sweat, the glucose and pH sensors detect the sweat glucose and pH levels, respectively. The triple glucose and quadruple pH sensing (fig. S1A) improve the detection accuracy (Fig. 4C). The correction using the measured pH enables more accurate glucose sensing (Fig. 4D). The patch works reliably under different skin temperatures (temperature range, 30° to 37°C) before, during, and after physical movements (fig. S15, A and B). The changes of skin temperature with the current experimental setting are small because of homeostatic functions of the nervous system (41). The patch is reusable and/or reattachable multiple times (fig. S15, C and D). Fig. 4 Sweat-based glucose monitoring and feedback therapy in vivo. (A) Optical camera image of the subject using a cycle ergometer for sweat generation with the wearable patch on the subject’s arm. (B) Real-time humidity monitoring to check the accumulation of sweat. (C) Multimodal glucose and pH sensing to improve detection accuracy. (D) Measured sweat glucose concentrations (n = 3), pH levels (n = 4), and corrected sweat glucose level (n = 3) based on the averaged pH (dotted line, glucose concentration measured by a commercial glucose assay). (E) Optical camera image of the disposable strip-type sensors connected to a zero insertion force (ZIF) connector. (F) Optical camera images of the sweat uptake via the fluidic channel of the strip. (G) Humidity monitoring of the disposable strip using impedance measurements. (H) Sweat glucose and pH monitoring using the disposable strip. (I) Ratio of sweat glucose concentrations (n = 3) measured by the patch and a commercial glucose assay with and without the pH-based correction before and after a meal. (J) Comparison of the sweat and blood glucose concentrations before and after a meal. (K) Optical camera image of the transdermal drug delivery device on the db/db mouse. (L) Blood glucose levels of the db/db mice for the treated groups (microneedles with the drugs) and control groups (without the patch, microneedle without the drugs) (*P < 0.05, **P < 0.01 versus control, Student’s t test). The disposable strip-type sensor is more convenient for the sweat analysis. The strip-type sensor first absorbs sweat (Fig. 2C) because of the capillary effect, and it is then connected to the hardware for sweat analysis through a ZIF connector (Fig. 4E). The absorbed sweat should cover the surface of both pH and glucose sensors (Fig. 4F). The amount of absorbed sweat can be monitored by measuring the impedance between the electrodes of the glucose sensor (Pt counter electrode and Ag/AgCl reference electrode) and those of the pH sensor (two PANi-deposited working electrodes) (Fig. 4G). The strip-type sensor can be stably operated with 4 μl of sweat (fig. S16A). When the sweat covers both glucose and pH sensors, the measurements begin (Fig. 4H). The calibration curves for a strip-type device are shown in fig. S16 (B and C). The sweat pH levels vary among subjects and depending on the physiological conditions of each subject (fig. S17A). The sweat glucose levels corrected by pH measurements before and after a meal agree well with the sweat glucose levels measured among subjects using a commercial glucose assay kit (Fig. 4I). Statistical analysis confirms the reliable correlation of sweat glucose levels measured by the wearable and disposable sweat glucose sensors with the blood glucose levels measured by a commercial glucose meter (fig. S17B). For the accurate estimation of the blood glucose level based on the sweat glucose level, the correlation factor between the glucose concentration of blood and sweat should be found for each individual subject with enough data (fig. S17C). With this data, it is clear that blood glucose concentration tends to vary before and after a meal, similarly to the measured sweat glucose concentration (Fig. 4J).