Solid-state sensors for touch-based rapid physiological and chemical sensing

Abstract

Methods, materials and devices relating to solid gel-free sensors for touch-based rapid physiological and chemical sensing are disclosed. In some embodiments of the disclosed technology, the sensor device includes a substrate, a plurality of first electrodes and a plurality of second electrodes formed on the substrate, a first current collector formed on the substrate and coupled to the plurality of first electrodes at one end of each of the first electrodes, and a second current collector formed on the substrate and coupled to the plurality of second electrodes at one end of each of the second electrodes, the first electrodes and the second electrodes being arranged in an alternating manner, and adjacent first and second electrodes being spaced apart from each other by a predetermined distance.

Description

[Technical field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This patent application claims priority to and the benefit of U.S. Provisional Application No. 63 / 266,513, entitled "GEL-FREE SENSOR FOR TOUCH-BASED RAPID PHYSIOLOGICAL AND CHEMICAL SENSING," filed on January 6, 2022. The entire contents of the aforementioned patent application are incorporated by reference into this disclosure.

[0002] The disclosed technology relates to methods and devices relating to solid-state sensors for touch-based rapid physiological and chemical sensing. [Background technology]

[0003] Recent fingertip touch-based sensors use a hygroscopic porous hydrogel for analyte collection and an electrolyte membrane covering the electrode surface to enable rapid non-invasive sensing, allowing sweat collection from the fingertip without the need for movement or active extraction. However, such sensors still require the use of hydrogel, which has a tendency to dry out and also poses other problems with the sensing process due to dilution and accumulation of analytes. The use of hydrogel is also difficult to implement and store, making the device and operation less practical and difficult to access for users. Summary of the Invention

[0004] The disclosed technology can be implemented in some embodiments to provide methods, materials, and devices relating to solid-state, gel-free sensors for touch-based, rapid physiological and chemical sensing.

[0005] In some implementations of the disclosed technology, a sensor device includes a substrate, a plurality of first electrodes formed on the substrate and extending in a first direction, a plurality of second electrodes formed on the substrate and extending in the first direction, a first current collector formed on the substrate and coupled to the plurality of first electrodes at one end of each of the first electrodes, and a second current collector formed on the substrate and coupled to the plurality of second electrodes at one end of each of the second electrodes, the first electrodes and the second electrodes being alternately arranged in the second direction, and adjacent first and second electrodes being spaced apart from each other by a predetermined distance.

[0006] In some implementations of the disclosed technology, the sensor device has: a plurality of first electrodes and a plurality of second electrodes; a first current collector coupled to the plurality of first electrodes at one end of each of the first electrodes; a second current collector formed on a substrate and coupled to the plurality of second electrodes at one end of each of the second electrodes. The first and second electrodes are arranged in an alternating manner, and adjacent first and second electrodes are spaced apart from each other by a predetermined distance.

[0007] In some implementations of the disclosed technology, in a sensor device comprising a plurality of electrode arrays for simultaneous or sequential sensing of a plurality of biomarkers of a physiological parameter, each of the electrode arrays comprises: a plurality of first electrodes and a plurality of second electrodes; a first current collector coupled to the plurality of first electrodes at one end of each of the first electrodes; and a second current collector formed on a substrate and coupled to the plurality of second electrodes at one end of each of the second electrodes. The first electrodes and the second electrodes are arranged in an alternating manner, and adjacent first and second electrodes are spaced apart from each other by a predetermined distance.

[0008] In some implementations of the disclosed technology, in a sensor array including multiple sensor devices, the multiple sensor devices are formed on a substrate for simultaneously or sequentially sensing multiple biomarkers in a biological fluid from different locations on the body.

[0009] In some implementations, the method includes placing a sensor device in contact with skin of a subject and measuring a biomarker in a biological fluid from the skin of the subject using the sensor device. In some implementations, the sensor device includes: a plurality of first electrodes and a plurality of second electrodes; a first current collector coupled to the plurality of first electrodes at one end of each of the first electrodes; a second current collector coupled to the plurality of second electrodes at one end of each of the second electrodes. The first electrodes and second electrodes are arranged in an alternating manner, and adjacent first and second electrodes are spaced apart from each other by a predetermined distance.

[0010] These and other aspects and implementations of the disclosed technology are described in more detail in the drawings, description, and claims. [Brief description of the drawings]

[0011] [Figure 1A] 1 illustrates an example of a sensor including at least two electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1B] 1 illustrates an example of a sensor including at least two electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1C] 1 illustrates an example of a sensor including at least two electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1D] 1 illustrates a fingertip in contact with a sensor implemented in accordance with some embodiments of the disclosed technology. [Figure 1E] 1 illustrates an example of a solid-state, gel-free sensor for touch-based rapid physiological and chemical sensing implemented in accordance with some embodiments of the disclosed technology. [Figure 1F] 1 illustrates radially aligned interdigitated electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1G] 1 illustrates interdigitated electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1H]1 illustrates concentric interdigitated electrodes implemented in accordance with some embodiments of the disclosed technology. [Figure 1I] 1 illustrates a multiplexed array of sensors for detecting multiple biomarkers according to some embodiments of the disclosed technology. [Figure 1J] 1 shows multiplexing electrodes for detecting multiple biomarkers and physiological signals on the same sensor. [Figure 1K] 1 shows multiplexing electrodes for detecting multiple biomarkers and physiological signals on the same sensor. [Figure 1L] 4 shows examples of spacing between electrodes.

[0012] [Figure 2A] 1 shows an example of an electrode functionalized with glucose oxidase. [Figure 2B] 1 shows an example of an electrode functionalized with lactate oxidase. [Figure 2C] 1 shows an example of an electrode functionalized with alcohol oxidase.

[0013] [Figure 3A] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating. [Figure 3B] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating. [Figure 3C] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating. [Figure 3D] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating. [Figure 3E] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating. [Figure 3F] We present a solid-state interdigitated electrode (IDE) for rapid glucose sensing via natural fingertip sweating.

[0014] [Figure 4A] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4B] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4C] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4D] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4E] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4F] The design and use optimization of a solid gel-free touch-based IDE sensor is presented. [Figure 4G] The design and use optimization of a solid gel-free touch-based IDE sensor is presented.

[0015] [Figure 5A] Calibration and extended on-body evaluation of a solid gel free-touch based glucose sensor is shown. [Figure 5B] Calibration and extended on-body evaluation of a solid gel free-touch based glucose sensor is shown. [Figure 5C] Calibration and extended on-body evaluation of a solid gel free-touch based glucose sensor is shown. [Figure 5D] Calibration and extended on-body evaluation of a solid gel free-touch based glucose sensor is shown.

[0016] [Figure 6] A dye experiment is presented to visualize the distribution of sweat along a fingertip over time.

[0017] [Figure 7A] 1 shows the in vitro characterization of the IDE sensor. [Figure 7B] 1 shows the in vitro characterization of the IDE sensor.

[0018] [Figure 8A] 1 shows in vitro calibration of an IDE glucose sensor. [Figure 8B] 1 shows in vitro calibration of an IDE glucose sensor.

[0019] [Figure 9] The selectivity of the IDE glucose sensor is shown.

[0020] [Figure 10A] Optimization of electrode spacing is shown. [Figure 10B] Optimization of electrode spacing is shown.

[0021] [Figure 11A] The repeatability of the sensor is characterized. [Figure 11B] The repeatability of the sensor is characterized.

[0022] [Figure 12A] Correlation of touch-based glucose levels to fingerstick capillary blood glucose (CBG) in 3-hour and 12-hour long-term glucose monitoring studies. [Figure 12B] Correlation of touch-based glucose levels to fingerstick capillary blood glucose (CBG) in 3-hour and 12-hour long-term glucose monitoring studies. [Figure 12C] Correlation of touch-based glucose levels to fingerstick capillary blood glucose (CBG) in 3-hour and 12-hour long-term glucose monitoring studies. [Figure 12D] Correlation of touch-based glucose levels to fingerstick capillary blood glucose (CBG) in 3-hour and 12-hour long-term glucose monitoring studies.

[0023] [Figure 13A] Clark Error Grid Analysis (CEGA) with additional calibration and time delay correction is shown. [Figure 13B] Clark Error Grid Analysis (CEGA) with additional calibration and time delay correction is shown.

[0024] [Figure 14] 1 illustrates an exemplary method for measuring biomarkers in a biofluid according to some embodiments of the disclosed technology. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0025] Methods, materials, and devices relating to solid-state gel-free sensors for rapid touch-based physiological and chemical sensing are disclosed. The disclosed technology can be implemented in some embodiments to provide a new class of non-invasive, painless sensors for direct sampling and frequent measurement of biomarkers in sweat and related physiological conditions (e.g., sweat rate). The sensor can operate with direct contact of the user using a fingertip or any other skin surface that emits passive, natural, thermoregulatory eccrine sweat, and can measure the concentration of various ions and biomolecules (e.g., sodium, potassium, chloride, glucose, lactate, urea, uric acid, bilirubin, hydroxybutyrate, vitamins, alcohol, levodopa, caffeine, cortisol, insulin, explosives, narcotics, nerve agents, fluoride, calcium, zinc, lead, cadmium, mercury) in sweat via electrical (e.g., conductive, piezoresistive, thermoresistive, piezocapacitive, thermoelectric, piezoelectric), chemical (e.g., nonspecific adsorption, specific binding, intercalation, insertion), or electrochemical (e.g., catalytic reaction, redox reaction) transduction methods. The sensor uses a closely packed or interdigitated electrode design that ensures a narrow interelectrode distance (<1 mm), which allows an ionic pathway between two or more electrodes for signaling when in contact with the fingertip. Some embodiments of the disclosed technology can eliminate the need for external analyte collection mechanisms such as hydrogels, hydrocolloids, porous materials, microfluidic channels, microneedles, iontophoresis, reverse iontophoresis, transdermal cholinergic delivery, etc., allowing for rapid, maintenance-free, user-friendly, near real-time biochemical and physiological sensing via direct skin contact. For this reason, the device offers reusability, extending operational time and allowing repeated and frequent use of the same sensor after simple cleaning (wiping with tissue or rinsing with water).

[0026] Diabetic patients who require constant glucose monitoring use finger-prick glucometers, which are highly invasive and painful. Alternative continuous glucose monitoring systems have been developed, but still require the insertion of a needle into the body for continuous sensing, which is invasive and requires maintenance. For the sensing of glucose and many other biomarkers important to human health, new wearable sensors have been proposed that feature non-invasive sensing from more accessible biological fluids such as sweat or interstitial fluid. The use of such epidermal sensors often requires challenging analyte extraction methods such as exercise, heating, or iontophoresis, and relies on complex device structures such as microfluidic devices, microneedles, and iontophoretic patches. Such methods are invasive and complex, and require a high degree of maintenance for self-monitoring. Furthermore, these devices either require large amounts of analyte or specialized analyte uptake mechanisms such as microfluidic channels or hydrogels to ensure coverage of the electrode surface.

[0027] Recently, fingertip touch-based sensors that can enable sweat collection from the fingertip without the need for movement or active extraction, using a hygroscopic porous hydrogel for analyte collection and an electrolyte membrane covering the electrode surface to enable rapid non-invasive sensing, have been proposed and demonstrated for sensing of various chemicals. However, this method still requires the use of hydrogel, which tends to dry out, and there are further problems with the sensing method due to dilution and accumulation of analytes. The use of hydrogel is also difficult to implement and store, making the device and operation less practical and less accessible to users.

[0028] The disclosed technology may be implemented in some embodiments to provide a new approach for epidermal sweat sensing that eliminates the need for any difficult sweat collection processes such as exercise, heating, chemical stimulation, or iontophoretic extraction. By using closely packed electrodes functionalized with transducers for biomarkers, users can perform non-invasive, painless, and maintenance-free sensing in a rapid manner. The closely packed or interdigitated electrode design eliminates the need for the use of hydrogels, making the device more accessible, simple, stable, and allowing frequent repeated measurements.

[0029] 1A-1C show examples of sensors including at least two electrodes implemented according to some embodiments of the disclosed technology. FIG. 1D shows a fingertip in contact with a sensor implemented according to some embodiments of the disclosed technology. FIG. 1E shows an example of a gel-free sensor for touch-based rapid physiological and chemical sensing implemented according to some embodiments of the disclosed technology. FIG. 1F shows radially aligned interdigitated electrodes implemented according to some embodiments of the disclosed technology. FIG. 1G shows interdigitated electrodes implemented according to some embodiments of the disclosed technology. FIG. 1H shows concentric interdigitated electrodes implemented according to some embodiments of the disclosed technology. FIG. 1I shows a multiplexed array of sensors for detecting multiple biomarkers according to some embodiments of the disclosed technology. FIGS. 1J and 1K show multiplexed electrodes for detecting multiple biomarkers and physiological signals on the same sensor. FIG. 1L shows examples of spacing between electrodes.

[0030] In some embodiments of the disclosed technology, a sensor device has a plurality of first electrodes extending in a first direction, a plurality of second electrodes extending in the first direction, a first current collector coupled to the plurality of first electrodes at one end of each of the first electrodes, and a second current collector coupled to the plurality of second electrodes at one end of each of the second electrodes.

[0031] In some embodiments of the disclosed technology, the first and second electrodes are disposed very close to each other (e.g., the distance between adjacent first and second electrodes may be less than 1 mm). In one example, the first and second electrodes and the first and second current collectors are disposed in the same direction, as shown in FIG. 1A. In another example, the first and second electrodes and the first and second current collectors are arranged in different directions, as shown in FIG. 1C. As shown in FIG. 1B, the first electrode has a predetermined shape, and the second electrode has a shape that at least partially surrounds the first electrode.

[0032] With reference to FIG. 1E, in some implementations, the first electrode is a working electrode and the second electrode is a reference / counter electrode.

[0033] 1F-1H, the first electrodes (e.g., working electrodes) and second electrodes (e.g., reference / counter electrodes) are arranged in an alternating manner, with adjacent first and second electrodes spaced apart from each other by a predetermined distance. In some implementations, the distance between adjacent first and second electrodes is about 1 mm. In one example, the distance between adjacent first and second electrodes is less than 1 mm. In another example, as shown in FIG. 1L, the distance between adjacent first and second electrodes is greater than 1 mm.

[0034] In one implementation, the first electrode (e.g., working electrode) and the second electrode (e.g., reference / counter electrode) are interdigitated and arranged radially, as shown in FIG. 1F. In another implementation, the first electrode (e.g., working electrode) and the second electrode (e.g., reference / counter electrode) are interdigitated and arranged parallel, as shown in FIG. 1G. In another implementation, the first electrode (e.g., working electrode) and the second electrode (e.g., reference / counter electrode) are arranged concentrically, as shown in FIG. 1H.

[0035] 1F-1H, a current collector is formed, then a reference / counter electrode is formed on the current collector, and a working electrode is formed on the reference / counter electrode. In some implementations, an insulator is formed over the working electrode, a signal transduction layer is formed over the insulator, and a protective layer is formed over the signal transduction layer.

[0036] In some implementations, the sensor device includes multiple electrode arrays, each of which includes a first electrode and a second electrode as discussed herein. In one example, different electrode arrays 142, 144, 146 can be used to detect different biomarkers 1, 2, and 3, as shown in FIG. 1I.

[0037] In some implementations, the first electrode may include a first working electrode WE-1 and a second working electrode WE-2, as shown in Figures 1J and 1K. In some implementations, the sensor device may further include a reference electrode, as shown in Figures 1J and 1K. In some implementations, the sensor device may further include additional sensors for measuring resistance and / or temperature, as shown in Figures 1J and 1K.

[0038] Figure 2A shows an example of an electrode functionalized with glucose, Figure 2B shows an example of an electrode functionalized with lactate, and Figure 2C shows an example of an electrode functionalized with alcohol.

[0039] The device is a sensor composed of at least two electrodes, at least two of which are in a close-packed or interdigitated configuration with a minimum distance of less than 1 mm from the other. The sensor includes a substrate made of fiberglass, silicon, paper, textile, or polymer plastic or elastomer. The electrodes can be manufactured via thin film deposition processes such as inkjet printing, sputtering, chemical / physical vapor deposition, screen printing, spray coating, flexography, computer numerical control milling, laser ablation, 3D printing, hydro or other additive or subtractive manufacturing processes such as electrospinning, etc. The sensor should include conductive electrodes composed of metals (e.g., Cu, Ag, Au, Pt), doped metal oxides (e.g., ITO, FTO), carbonaceous materials (e.g., graphite, graphene, reduced graphene oxide, activated carbon, carbon nanotubes, diamond), conducting polymers (e.g., PEDOT:PSS, polypyrrole, polyaniline, poly-p-phenylene, polythiophene), or 2D materials (e.g., MoS2, WSe2, VO2). The electrodes can be functionalized with a variety of chemical / electrochemical transducers for sensing biomarkers, such as enzymes (e.g., lactate oxidase, lactate dehydrogenase, glucose oxidase, glucose dehydrogenase, bilirubin oxidase, uricase, urea oxidase, alcohol oxidase, alcohol dehydrogenase, tyrosinase, catalase), catalysts (e.g., platinum, ruthenium, palladium, rhodium, silver), redox mediators (Prussian blue, Meldola blue, methylene blue, indigo carmine, 2,2'-bipyridine, 1,4-naphthoquinone, tetrathiafulvalene, tetracyanoquinodimethane, ferrocene), antibodies, ion-selective membranes, silver / silver chloride mixtures, molecularly imprinted membranes, or aptamers.The sensor may include additional closely packed or interdigitated electrodes for physiological sensing with physical transducers such as thermoresistive, thermoelectric, piezoresistive, piezocapacitive, piezoelectric, photovoltaic, physisorptive, or chemisorptive materials that sense physical properties such as temperature, skin moisture level, or pressure. The sensor may also include a protective layer composed of a polymeric material such as Nafion, chitosan, ethyl cellulose, polyvinyl chloride, etc. The sensor may also include an insulating layer composed of a dielectric material. The electrode layer composition may be a single material as listed above, or a mixture or composite of the above materials.

[0040] The fabricated sensors may be used for non-invasive chemical sensing with any complementary physiological sensing. In one example, a user can use a densely packed or interdigitated sensor functionalized with glucose oxidase to quickly and simply sense glucose in sweat and establish a correlation with blood glucose levels. A user can directly press the sensor for about 30-60 seconds, where a potential step is applied and the chronoamperometric signal is read out to obtain a signal. In another example, a user can use a densely packed or interdigitated sensor made with poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) and lactate oxidase, along with a densely packed or interdigitated sensor made with silver, to simultaneously sense sweat lactate concentration and skin moisture level, and the skin moisture data can be used to calibrate the sweat lactate signal. In another embodiment, a user can use a densely packed or interdigitated three-electrode sensor, where one of the electrodes is functionalized with a pH-sensing polyaniline polymer and another electrode is functionalized with an alcohol oxidase enzyme. The user can directly male the sensor for 30-60 seconds, where potential steps for chronoamperometry are applied to the enzyme electrode and the open circuit potential is monitored on the pH detection electrode, sweat pH data and sweat alcohol signal can be obtained simultaneously, and the sweat pH data can be used to calibrate the sweat alcohol signal to improve accuracy. In these examples, the sensor can simply be reused multiple times after a quick and simple cleaning of the electrode surface by rinsing with water or wiping with a tissue. For near real-time measurements over long periods of time, the data obtained can be validated with other measurement methods such as glucometers, blood tests, and respirometers to establish a personal calibration to relate thermoregulatory sweat biomarker levels with blood or other environmental biomarker levels of interest.

[0041] 1E, in some implementations, the gel-free sensor 100 has a plurality of first electrodes 110 extending in a first direction, a plurality of second electrodes 120 extending in the first direction, a first current collector 130 coupled to the plurality of first electrodes 110, and a second current collector 132 coupled to the plurality of second electrodes 120. In one example, the first electrodes 110 and the second electrodes 120 are alternately arranged in the second direction. In one example, the first direction is perpendicular to the second direction. In some implementations, the first current collector 130 is connected to one end of each of the first electrodes 110, and the second current collector 132 is connected to one end of each of the second electrodes 120.

[0042] In some implementations, the first electrode 110 is used as a working electrode and the second electrode 120 is used as a reference / counter electrode. In some implementations, the first electrode 110 includes a PEDOT:PSS-Prussian Blue (PB) cathode and the second electrode 120 includes a poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) anode. In some implementations, the first electrode 110 and the second electrode 120 are solid-state interdigitated electrodes (IDEs) printed on a styrene-isoprene-styrene block copolymer (SIS) substrate. In some embodiments, the solid interdigitated electrodes (IDEs) are decorated with glucose oxidase (GOx) enzymes that selectively react with glucose in fingertip sweat for subsequent detection. In some implementations, the first electrode 110 and the second electrode 120 extend in the same direction (e.g., the second direction) and are arranged alternately at a predetermined distance. In one example, a predetermined distance provides an ion path between adjacent first and second electrodes 110 and 120. In some implementations, the predetermined distance is less than 1 mm.

[0043] In some implementations, the gel-free sensor 100 may further include a substrate configured to support the first and second electrodes 110, 120 and the first and second current collectors 130, 132. In one example, the substrate includes at least one of glass, silicon, paper, textile, polymer plastic, or elastomer. In some implementations, the gel-free sensor 100 may further include a glucose oxidase (GOx) enzyme layer formed on the first and second electrodes and the first and second current collectors to selectively react with glucose in the bodily fluid for detection. In one example, the bodily fluid includes fingertip sweat.

[0044] The disclosed technology can be implemented in some embodiments to provide a solid-state touch sensor that enables frequent, accurate, non-invasive glucose monitoring.

[0045] Figures 3A-F show solid-state interdigitated electrodes (IDEs) for rapid glucose sensing via natural fingertip sweating. Figure 3A shows an exploded view of the structure of a printed PEDOT:PSS-based IDE sensor. Figure 3B shows an illustration of sweat accumulated within the finger grooves establishing a connection between the PEDOT:PSS-PB cathode and the PEDOT:PSS anode for subsequent electrochemical reactions. Figure 3C shows an illustration suggesting the use of a gel-free IDE sensor for rapid and frequent non-invasive sensing of glucose levels during events that cause glucose excursions. Figure 3D shows CA curves demonstrating the reusability and reversibility of the sensor allowing for repeated sensing. Figure 3E shows an illustration suggesting the possible use of a gel-free IDE glucose sensor for continuous glucose monitoring throughout the day. Figure 3F shows Clark's error grid versus reference self-monitoring blood glucose (SMBG) testing, presenting the gel-free IDE sensor for fingertip sweating as a reliable blood glucose sensing mechanism. Data points consisting of 50 measurements taken before and 20 min after a meal over a 5-day period were collected from 5 healthy subjects.

[0046] Figures 4A-G show the design and use optimization of a gel-free touch-based IDE sensor. Figure 4A shows (i) the experimental setup for visualizing sweating on a fingertip when pressed against the sensor, and the corresponding optical image of a fingertip with the indicator turning blue due to sweat in the grooves, (ii) the corresponding % area covered by sweat shown by the color-changing dye, and (iii) an illustration of the resistance between the anode and cathode versus time of pressure application. Figure 4B shows (i) a schematic of a conventional screen-printed electrode (SPE) and IDE design with controlled interelectrode distance d, (ii) the chronoamperometric (CA) response of gel-free electrodes in conventional SPE (left) and IDE (right) designs with an electrode spacing of 0.25 mm to finger sweating, and (iii) a summary of the change in CA current response to finger sweating with different interelectrode distances, corresponding to capillary blood glucose levels of 85 mg / dl (402) and 128 mg / dl (404). FIG. 4C shows (i) the CA curves, (ii) the final current signals after different contact (sweat accumulation) times when a finger is pressed against the sensor. FIG. 4D shows (i) the CA curves, (ii) the final current signals with different pressing pressures. FIG. 4E shows (i) the CA response for a subject with a steady pre-meal glucose level with six repeated contacts on the hydrogel-covered electrode (412) and the solid electrode (414), (ii) bar graphs summarizing the current signals (416, 418), (iii) the CA response for a subject with a post-meal drop in glucose level with six repeated contacts on the hydrogel-covered electrode (420) and the solid electrode (422), and (iv) bar graphs summarizing the current signals (424, 426). Figure 4F shows the repeatability of the sensor with five replicate measurements at 3 minute intervals: in Figure 4F(i), line 428 shows fingertip sensing corresponding to pre-prandial capillary blood glucose (CBG) of 83 mg / dl and line 430 shows fingertip sensing corresponding to post-prandial CBG of 132 mg / dl; in Figure 4F(ii), the corresponding bar graph shows the repeatability of pre-prandial and post-prandial results, resulting in RSDs of 2.4% and 1.9%, respectively. Figure 4G shows (i) 20 replicate measurements performed at 3 minute intervals (total of 60 minutes) on fasting blood glucose levels, (ii) the repeatability of these 20 replicate measurements; RSD 3.9%.

[0047] Figures 5A-5D show the calibration and on-body extended evaluation of the gel-free touch-based glucose sensor. Figure 5A shows the current response of the touch-based glucose sensor to the corresponding fingertip prick SMGB standard and the corresponding linear regression generated from five healthy subjects obtained before and after a 5-day meal. The subjects include both males and females aged 22-36 years. Figure 5B shows: (i) a schematic diagram depicting the time course for continuous glucose monitoring over 170 days, during which the same food and beverages are offered to the volunteers after 20-min and 110-min tests, while the touch-based glucose sensor and fingerstick SMBG signals are recorded every 5 and 10 days, respectively; (ii) the results of the touch-based glucose sensor of three non-diabetic volunteers converted using the corresponding fingertip SMBG along with the calibration on day 1. FIG. 5C shows a schematic diagram depicting the time course for a 12-hour glucose monitoring study, during which a high-carbohydrate meal was provided to a non-diabetic individual 1, 5, and 10 days after the study began: (i) touch-based glucose sensor and fingertip SMGB signals were collected every 5 and 10 minutes, respectively, for 1 hour post-meal and 1 hour inter-meal; (ii) the touch-based glucose sensor results of two non-diabetic volunteers were converted using their corresponding fingertip SMGB reference along with the calibration on day 1. FIG. 5D shows the Clarke Error Grid analysis of the touch-based glucose sensor against SMGB glucose as the reference.

[0048] Figure 6 shows a dye experiment visualizing the distribution of sweat along a fingertip over time.

[0049] Figures 7A and 7B show the in vitro characterization of the IDE sensor. Figure 7A shows: (i) the characterization of the IDE sensor in PBS (pH = 7.3), including two-electrode CV, (ii) CA at different potentials from 0 V to -0.3 V, (iii) current vs. potential plots suggesting an optimum potential (cathode vs. anode) of -0.1 V. Figure 7B shows: (i) the characterization of the IDE sensor in artificial sweat (pH = 7), including two-electrode cyclic voltammetry, (ii) chronoamperometry at different potentials from 0 V to -0.3 V, (iii) current vs. potential plots suggesting an optimum potential of -0.1 to -0.15 V (cathode vs. anode).

[0050] Figures 8A and 8B show the in vitro calibration of the IDE glucose sensor. Figure 8A shows: (i) the CA of IDE at glucose concentrations of 100-500 μM in PBS at pH 7.3, (ii) a sensitivity of 0.97 nA / μM (n=3, r 2 = 0.99). Figure 8B shows the resulting PBS calibration plot with (i) the CA of IDE with glucose concentrations from 0 to 500 µM in artificial sweat at pH 6.0, (ii) a sensitivity of 0.97 nA / µM.

[0051] FIG. 9 shows the selectivity of the IDE glucose sensor: CA response of the sensor in PBS (blank), followed by addition of 100 μM glucose, lactate, ascorbic acid, aminophen acetate, and uric acid, respectively.

[0052] Figures 10A and 10B show the optimization of electrode spacing. Figure 10A shows the CA response of a typical SPE design with different electrode spacing. Figure 10B shows the CA response of an IDE design with different electrode spacing.

[0053] Figures 11A and 11B show the characteristics of sensor reproducibility. Figure 11A shows the overlaid CA responses of five different IDE glucose sensors contacted by the same subject with the same finger. Figure 11B shows the corresponding results with an RSD of 4.3%.

[0054] Figures 12A-12D show the correlation between touch-based glucose levels and fingerstick CBG in 3-hour and 12-hour long-term glucose monitoring studies. Figure 12A shows a 1-day calibration without considering any time delay. Figure 12B shows a 1-day calibration with a 10-minute delay. Figure 12C shows a 3-day calibration with a 10-minute delay. Figure 12D shows a 5-day calibration with a 10-minute delay. Pearson correlation coefficients are labeled in the corresponding subfigures.

[0055] 13A and 13B show CEGA with additional calibration and time delay correction. Fig. 13A shows CEGA with calibration on (i) day 1, (ii) day 3, (iii) day 5 without considering any time delay. Fig. 13B shows CEGA with calibration on (i) day 1, (ii) day 3, (iii) day 5 when considering a 10 minute delay between CBG and fingertip sweat glucose.

[0056] Type 1 diabetes is a chronic disease that requires frequent blood glucose testing to prevent acute and long-term complications. Common glucose monitoring methods, including finger-prick capillary blood testing and subcutaneous continuous glucose monitors, are painful and invasive, respectively, while recent non-invasive epidermal sensors have limited practicality and reliability. The disclosed technology can be implemented in some embodiments to provide a convenient, fast, and accurate approach using interdigitated electrode transducers for frequent touch-based sweat glucose biosensing that utilizes passive sweating from the fingertip. The interdigitated design establishes a solid interface and eliminates the need for sweat-collecting hydrogels, which greatly simplifies the sensing workflow and provides stability and reusability. The sensor can be used repeatedly for 1-day glucose self-monitoring and provides reliable glucose data with a low mean-absolute relative difference of 8.69%, which is comparable to that of commercially available fingerstick and continuous glucose monitors. The new protocol based on some embodiments provides convenient, practical, and painless glucose sensing, promotes frequent self-testing, and improves diabetes self-care.

[0057] Diabetes mellitus is a global health problem and ranks among the leading causes of death. Frequent monitoring of blood glucose levels is critical to understand disease progression and optimize its control. Over the past three decades, diabetic patients have relied on performing finger-prick self-monitoring of blood glucose (SMBG) multiple times daily, but the painful, inconvenient, and invasive nature of such finger-prick procedures undermines patient compliance and significantly reduces testing frequency. Continuous glucose monitoring (CGM) addresses these limitations of SMBG and offers a significant improvement in the management of diabetes. While eliminating the need for finger-pricks and providing more insight from continuous data, CGM still uses invasive (e.g., ~10 mm long), expensive needles, requires daily SMBG calibration, and is challenged by biofouling, long stabilization times, glucose time delay, and limited lifetime. In recent years, alternatives such as noninvasive sweat and interstitial fluid (ISF) glucose sensors have been proposed, but their practical utility is largely hindered by complex biofluid extraction mechanisms (e.g., exercise, sweat-induced drugs, reverse iontophoresis) that may induce non-native metabolic activity or dilution that affects correlation with blood glucose concentrations. Indeed, it has been demonstrated that sweat glucose can accurately reflect blood glucose levels upon proper collection of sweat. Recent advances have ingeniously leveraged direct sampling of passive thermoregulatory sweat from high sweat rate fingertips for chemical sensing, leading to the monitoring of cortisol, levodopa, glucose, caffeine, lactate, and ascorbic acid concentrations. However, such touch-based sensing relies on hydrogels as sweat collection interfaces (made of agarose or polyvinyl alcohol) mounted on the sensor, which introduces significant challenges and errors due to analyte carryover, dilution, and solvent evaporation. Thus, the use of hydrogel interfaces suffers from inconsistent sensing results and requires inconvenient lengthy gel replacement, thus significantly hindering the reliability, reproducibility, durability, simplicity, and overall practicality of such techniques.

[0058] The disclosed technology can be implemented to address the above issues by providing a reusable solid-state touch-based electrochemical sensing protocol for reliable and frequent extended glucose monitoring. To eliminate the need for hydrogels, the sensor relies on a solid-state interdigitated electrode (IDE) design consisting of a poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) anode and a PEDOT:PSS-Prussian blue (PB) cathode printed on a styrene-isoprene-styrene block copolymer (SIS) substrate. These printed IDE electrodes are decorated with glucose oxidase (GOx) enzyme that selectively reacts with glucose in fingertip sweat for subsequent detection (Figure 3A). The interdigitated PEDOT:PSS electrode design gives the capability of direct contact-based measurement of sweat glucose without hydrogels or ion-conducting interfaces, and passive sweat rapidly spreads along the fingertip grooves to establish an ionic pathway between the anode and cathode of the IDE (Figure 3B). The use of conductive PEDOT:PSS polymers results in ion transport and low electrode impedance, which are advantageous for solid-state bioelectronic interfaces. Relying on fast natural sweating on the fingertip, the solid-state IDE biosensor can provide accurate noninvasive sweat sensing within 90 seconds, which involves 60 seconds of sweat accumulation followed by 30 seconds of chronoamperometry (CA) at a low voltage of -0.1 V, providing a highly selective current response that can be easily calibrated to blood glucose. Such a simple, rapid, user-friendly, and painless glucose self-test offers a high sensing frequency that is comparable to state-of-the-art CGM technology (1 data point per 1.5-15 min), but in a completely noninvasive manner. Thus, this protocol allows users to closely track blood glucose levels and capture dynamic events, including rapidly fluctuating concentrations (Figure 3C). Furthermore, the elimination of hydrogels and associated chemical buildup allows for the reuse of the same sensor throughout the day (after a quick and gentle wipe with a tissue), eliminating the need for any stabilization time.The sensor produces highly reproducible signals upon repeated use and is highly reversible to follow rapid fluctuations in blood glucose concentration (Figure 3D). One sensor strip can be reliably used for continuous glucose testing for one day, allowing users to track glucose levels and conveniently and closely detect potential glycemic abnormalities (Figure 3E). A simple two-data-point individualized calibration (at different blood glucose levels) on the first day is used to fully address human variability (e.g., sweat rate). The sensor delivers highly accurate blood glucose concentration data, closely matching commercial fingertip capillary blood glucose (CBG) levels with a short time delay, with a low mean absolute relative difference (MARD) of 8.69% and a zone-A ratio (Z-A ratio) of 89.4% in Clark Error Grid Analysis (CEGA) (Figure 3F). This is comparable to the accuracy of commercial CGM and fingertip SMBG.

[0059] To take full advantage of the low sweat rate on fingertips, we closely examined the sweat distribution on the fingertip surface using a color-changing bromocresol green dye that turns blue on contact with sweat. A SIS substrate uniformly coated with a thin layer of dye, as shown in Figure 4A(i), was placed on a glass slide and a video was taken to record the color change of the dye when a clean, dry finger was pressed against the substrate (Figure 6). Over time, most of the grooves are covered by sweat within 60 seconds, establishing a conductive path between the anode and cathode (Figure 4A(ii)). Although the dye color darkens over time, additional areas are not covered by sweat. This is consistent with measurements of the internal resistance between the two electrodes, which is mostly stable after 60 seconds (Figure 4A(iii)).

[0060] The performance of the gel-free touch-based sensor for glucose monitoring was first characterized in vitro using glucose in phosphate buffer solution (PBS) at pH 7. The optimal potential was optimized at -0.1 V between the cathode and anode, yielding the highest CA response. Similarly, the sensor was evaluated in artificial sweat (which has a lower pH of 6), showing that the optimal potential and response were not affected by the lower pH (Figures 7A, 7B, 8A and 8B). The benefits of the solid-state IDE electrode, along with a hydrogel sweat collection interface, were then tested on the fingertips of human subjects in comparison to a traditional screen-printed electrode (SPE) sensor design. Compared to the SPE disk design, the IDE layout maximizes the number of electrical connections between the anode and cathode while reducing the spacing between the electrodes (Figure 4B(i)). As a result, when sensing from the same finger, the sensitivity of the IDE is considerably higher than the SPE, even with a similar interelectrode distance (ii in Figure 4B). In general, the sensitivity increased with decreasing electrode spacing to 0.25 mm, which was used in subsequent measurements (Fig. 4B iii, Figs. 10A and 10B). Different sweat accumulation times before CA measurements were evaluated and showed that the response gradually increased with contact time up to 45 s, then it started to plateau (Fig. 4C), which is consistent with the results in Fig. 4A. The pressing force on the electrodes was also optimized; the results (shown in Fig. 4D) indicate that 5 N per finger was sufficient to reach a stable signal.

[0061] Thus, the IDE design enables solid-state contact-based fingertip sensing, offering significant advantages over common hydrogel-based sweat collection mechanisms in terms of simplicity, reusability, testing frequency, and data reproducibility. As a comparison, the same finger was repeatedly tested on the same sensor with and without hydrogel at steady and declining glucose levels (before and 30 min after a meal, respectively). As shown in Figure 4Ei-ii, for constant glucose levels, the hydrogel-covered sensor shows an increase in current signal due to glucose carryover and accumulation with repeated contacts, while the solid-state IDE sensor shows good stability and the differences due to these repeated contacts are negligible. In contrast, for declining glucose levels (Figure 4Eiii-iv), the solid-state IDE sensor can accurately capture the dynamically declining glucose concentration, while the hydrogel-covered sensor shows a slowly increasing response due to the combination of competitive effects of the declining glucose concentration and its accumulation in the gel, resulting in an unrealistically increasing response associated with such carryover effects. The reproducibility of the sensor at different glucose concentrations was investigated from five replicate tests in the fasting state (CBG = 83 mg / dl) and 20 min after a meal (CBG = 132 mg / dl) (Figure 4F). Highly reproducible CA signals were observed in these replicate touch-based measurements, with RSD values ​​of 2.4% and 1.9%, respectively. Furthermore, the reproducibility of the glucose sensor was evaluated with fasting blood glucose levels over an extended period of 60 min including 20 replicate tests spaced 3 min apart (Figure 4G). The highly reproducible current signal over this long series with a low RSD value of 3.9% (n = 20) revealed good reproducibility over multiple replicate frequent measurements. Furthermore, the selectivity and repeatability of the sensor were characterized (Figures 9, 11A and 11B).

[0062] Following system optimization and characterization, the accuracy of the touch sensor was evaluated in five healthy nondiabetic male and female subjects between the ages of 22 and 36 years. Due to individual differences (e.g., sweat rate, fingertip size), each subject was first tested before and 20 min after a meal to build an individualized calibration between the CA signal of the gel-free touch-based glucose sensor and the CBG amount obtained from fingertip SMBG (Figure 5A). Through 10 data points generated over the first five days, the five subjects obtained calibrations with Pearson values ​​above 0.98, 0.99, 0.96, 0.96, and 0.99, respectively, suggesting a consistent correlation between glucose in passive sweat from the fingertip and fingertip blood levels across different days. Such consistency eliminated the need for repeated recalibration, and therefore, only the data points from day 1 were used for subsequent individual calibrations.

[0063] [Table 1]

[0064] The ability to use touch-based, gel-free glucose sensors to replace fingertip-worn SMBG for frequent, noninvasive, and rapid glucose sensing was first evaluated over a 3-hour period, during which all subjects consumed a high-calorie meal followed by a sugary juice, during which touch-based sensing (with the same sensor) was performed every 5 minutes and fingertip-worn CBG reference measurements were performed at 10-minute intervals (Figure 5Bi). As shown in Figure 5Bii, blood glucose values ​​(estimated from calibration on day 1) showed good agreement with the reference blood glucose values ​​for all three subjects. To further explore the solid-state IDE glucose sensor as a potential alternative to CGM, long-term glucose monitoring was performed over a 12-hour period. During this time, subjects consumed breakfast, lunch, and dinner, and in parallel, touch-sensing and fingerstick CBG measurements were performed every 5 and 10 minutes, respectively, within 1 hour after the meal, in addition to hourly measurements between meals (Figure 5Ci). Results from two subjects show a continuous dynamic profile of glucose concentrations over the entire day, with high values ​​and temporal trends consistent between both measurement methods (Figure 5Cii). Notably, a ~10 min lag between fingertip glucose levels and CBG was observed among all subjects in both the 3-h and 12-h trials, consistent with previous reports and reflecting diffusive transport of glucose from capillaries to apocrine sweat glands. Similar time delays are observed in commercially available CGM devices and are generally addressed via various predictive algorithms that improve accuracy. By taking into account the 10 min time lag, touch-based glucose calibration can be adjusted, further improving the correlation of the five trials, as shown in Figures 12A-12D, 13A, and 13B, and Table 1.

[0065] Mean absolute relative difference (MARD) and Clarke error grid analysis (CEGA) are common methods to evaluate the accuracy and reliability of glucose sensing technologies. Commercially available CGM systems are generally characterized by MARD values ​​of 9-14%. The gel-free touch-based sensor was characterized by a MARD of 8.69% (data size n=160), reflecting its high accuracy comparable to commercially available glucose monitoring devices. Additional calibration (against blood; over the first 3-5 days instead of the initial calibration) did not show a significant improvement in MARD, but considering the time delay, the MARD of the sensor could be further reduced to 4.76% (Table 1). CEGA evaluates the reliability of the blood glucose measurement method by dividing the grid into intervals A-E. Interval A corresponds to within 20% of the criterion, meaning that it has no impact on medical action. As shown in Figure 5D, the solid-gel free-touch based glucose sensor has 87.9% of data points landing in region A and 100% of all values ​​falling in the combined A+B region, indicating high accuracy with very low chance of misdiagnosis of hypoglycemia and hyperglycemia. Notably, considering a time delay of 10 min, with 98% of the data in region A, yielding even higher sensor reliability, although additional calibration data points did not significantly improve accuracy (Figures 13A and 13B).

[0066] The disclosed technology can be implemented in some embodiments to provide a highly accurate, simple, and rapid glucose sensing protocol that uses interdigitated solid-state PEDOT:PSS-based electrodes. It utilizes passive sweating of the fingertip to enable reliable, near real-time, non-invasive monitoring of glucose levels. Eliminating the need for a sweat-collecting hydrogel interface has greatly simplified operation to allow frequent repeated measurements throughout the day while providing superior analytical performance compared to conventional hydrogel-based SPE sensors. Compatible with a variety of subjects, the sensor can rapidly establish an individualized calibration (from just the first two fingertip measurements), showing considerable promise for replacing painful and frequent finger-prick SMBG and invasive CGM techniques for long-day continuous glucose monitoring. The new protocol provides high accuracy with low MARD and offers good CEGA metrics, which have a painless (blood-free and needle-free) and rapid operation, comparable to those of commercially available glucose sensing technologies. The same low-cost sensor can be used without restabilization to perform more than 50 measurements throughout the day. Such convenient touch-based sensing would significantly increase the frequency of self-testing compared to conventional SMBG to improve diabetes control. The disclosed technology may also be implemented in some embodiments to enable large-scale validation with diverse subjects and to further speed and simplify advanced bloodless calibration processes and prediction algorithms. This will be accompanied by improved understanding of the fingertip passive sweat phenomenon and the role of electrode geometry in hydrogel-free IDE setups. The translation of such contact-based sensing into a wearable device for daytime and nighttime passive and continuous monitoring of glucose would further promote its practical use as a true CGM alternative. Combined with engineering efforts to create user-friendly sensor prototypes, these developments can achieve highly reliable, painless, rapid and frequent glucose self-testing for home and other distributed environments for improved management of diabetes, as well as for easy and accurate non-invasive monitoring of other important sweat biomarkers.

[0067] According to some embodiments of the disclosed technology, graphite, toluene, acetone, ethanol, glutaraldehyde, D-(+)-glucose, glucose oxidase (GOx), Ag flakes, potassium chloride (KCl), sodium chloride (NaCl), anhydrous sodium phosphate, Prussian blue, and sodium dodecylbenzene sulfonate (DBSS) can be used for the fabrication of interdigitated electrodes. According to some embodiments of the disclosed technology, styrene-ethylene-butylene-styrene (e.g., SEBS G1645) triblock copolymer can be used for the fabrication of interdigitated electrodes. According to some embodiments of the disclosed technology, screen-printable PEDOT:PSS paste can be used for the fabrication of interdigitated electrodes.

[0068] Fabrication of interdigitated electrodes

[0069] In some implementations, the IDE sensor can be fabricated using layer-by-layer screen printing with customized four inks: electrochromic poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) ink, silver ink for interconnects, and insulating resin composed of SEBS. The PEDOT:PSS ink can be formed using a mixture of inks prepared using 1 g of PEDOT:PSS paste, 0.2 mL of toluene, 0.15 mL of DBSS (75 mg / mL in DI water), and 0.0135 mL of fluorosurfactant FS-65. Similar ingredients can be used to create the PEDOT:PSS-PB ink, with an additional 20 mg of PB powder added per 1 g of mixture. The stretchable silver ink can be synthesized. In some implementations, the silver flakes, toluene, and SEBS can be mixed in a weight ratio of 4: 2.37: 0.63. All inks can be homogenized by mixing in a dual asymmetric centrifugal mixer for example at a speed of 1900 RPM for 5 minutes. The insulating resin is prepared by dissolving SEBS in toluene solution (weight ratio 4:10). The solution is then mixed at 1900 RPM for 20 minutes or until the SEBS is completely dissolved in the solution.

[0070] After the ink preparation, a flexible substrate for printing the fingerprint sensor is fabricated. With the help of stainless steel wire, a 1000 μm SIS layer is formed on a plastic sheet of polyethylene terephthalate. The thin layer is dried at 60 °C for 30 min. The printed electrode pattern (electrode area 0.02 cm) is then printed on a 1000 μm SIS substrate. 2 ) was designed with the software and a stainless steel plate (12 x 12 in 2) etching to produce a metal stencil. The electrochemical system consisting of a working electrode and a reference electrode was screen-printed using an MPM-SPM semi-automatic screen printer. The printing process first printed the reference electrode using PEDOT:PSS ink on the SIS substrate, followed by a curing step at 120 °C for 30 min. Then, the working electrode was printed using PEDOT:PSS-PB ink, and similar drying conditions were applied after printing the electrodes. Then, a silver interconnect pattern was printed on top of the electrode system, followed by a curing step at 90 °C for 15 min. The wiring is insulated using SEBS resin. Finally, the insulator is dried at 90 °C for 10 min to obtain the printed IDE electrodes.

[0071] The IDE electrodes are modified by drop-casting a mixture of 6 µL of glucose oxidase (20 mg / mL in PBS 0.1 M pH 7.3) and 3 µL of glutaraldehyde (1% in DI water) onto the exposed electrode surface. After modification, store the electrodes in a refrigerator at 4 °C overnight.

[0072] In vitro sensor characterization

[0073] In vitro characterization of the touch-based glucose sensor was carried out using both 0.1 M PBS (pH 7.3) and artificial sweat (AS, pH 6.0). Two-electrode cyclic voltammetry between -0.35 V and 0.55 V was performed to observe the redox peaks of PB molecules at 0.02 V and -0.08 V in PBS, and 0.12 V and -0.04 V in AS. Optimization of the applied potential was studied at different potentials decreasing from 0 V to -0.3 V at 50 mV intervals. From the current responses at different potentials, -0.1 V was selected and used in subsequent in vitro and on-body experiments (Figures 7A and 7B). Under the optimized conditions, glucose levels varied from 0 to 500 μM with the addition of 100 μM, showing good linearity for both PBS and AS (Figures 8A and 8B). Additionally, selectivity studies were performed in PBS followed by the addition of 100 μM glucose, lactate (LA), ascorbic acid (AA), acetaminophen (AP), and uric acid (UA) (Figure 9).

[0074] On-body chemical fingerprint sensing

[0075] Fingerprint sensing was performed in healthy consenting subjects in strict compliance with the clinical trial protocol approved by the Institutional Review Board of the University of California, San Diego. Volunteers were asked to place their index finger on the sensor and perform all physical assessments. Glucose levels were verified using a commercially available blood glucose meter. Volunteers were asked to clean their hands with water and soap before using the sensor. Physical examination results were obtained using a Metrohm benchtop Autolab potentiostat / galvanostat 204. A chronoamperometric potential step of -0.1 V for 30 s was used during all experiments. To collect blood glucose levels, each individual's finger was pricked and a small drop of blood was analyzed. The fingertip glucose signal was then measured by touching the sensor for 60 s. Two chronoamperometric steps were then performed to quantify the current signal. The final current value from the second scan was taken as the corresponding sweat glucose level. After each sensing session, the sensor was first wetted with deionized water and then gently wiped with a paper tissue and dried for the next measurement.

[0076] Personal calibration in physical experiments

[0077] Blood and fingertip glucose signals were obtained from five volunteers for five consecutive days in a fasting state and 20 minutes after ingesting a sugar-rich meal. Current and blood glucose values ​​were plotted together and, using linear fitting, a slope b and an intercept a were obtained for each subject using data from a given number of days. To obtain the predicted glucose concentration for each subject, the following equation is used: BG=a+bi where "i" represents the current value obtained after the second chronoamperometric scan.

[0078] Augmented Body Experiment

[0079] Three subjects were recruited to conduct a 3-hour monitoring experiment. Fasting glucose levels were recorded for 20 minutes, followed by the ingestion of a meal. After 90 minutes, a high sugar content drink was provided to each subject. At the same time, blood and fingertip glucose were monitored every 10 and 5 minutes, respectively. Two subjects were recruited to conduct a 12-hour experiment. The experiment started by monitoring both blood and fingertip glucose every hour. The regular monitoring of both parameters was only changed after the ingestion of a meal 1, 5, and 10 hours after the start of the experiment. In these time slots, blood and fingertip glucose were monitored every 10 and 5 minutes, respectively.

[0080] Frequent or continuous measurements of glucose are necessary to achieve tight glycemic control for effective management of diabetes. Noninvasive glucose monitoring has always been considered the “holy grail” of next-generation biosensing technologies because of its significant impact on large populations and huge commercial markets. However, after decades, diabetic patients who rely on multiple blood glucose self-monitoring daily to prevent life-threatening complications still rely on painful and inconvenient finger-prick glucometers and expensive and invasive continuous glucose monitoring (CGM). Noninvasive glucose sensing technologies based on sweat and interstitial fluid have been proposed, but their practical implementation remains limited due to their long and complicated biofluid extraction processes. To address these challenges, our group recently explored the use of passive sweating from the fingertip for chemical sensing due to its high natural sweat rate. This fingertip-based contact sensing uses a hydrogel to extract sweat from the fingertip to enable chemical monitoring, but faces various limitations such as analyte carryover and dilution, solvent evaporation, and inconvenient gel placement and storage, hindering its practical utility for tracking dynamically changing concentrations.

[0081] To address the above issues, the disclosed technology may be implemented in some embodiments to provide a reusable solid-state touch-based electrochemical sensing protocol for reliable, convenient, frequent, extended glucose monitoring. In addition, the disclosed technology may be implemented in some embodiments to provide a unique solid-state interdigitated electrode transducer for direct contact-based glucose monitoring without the need for a sweat extraction mechanism. Upon touching the fingertip, the solid interface is rapidly covered by a sweat layer across the interdigitated electrodes, which rapidly records glucose levels without preconditioning or incubation. Such a painless biosensor features reusability and speed, and thus can significantly increase the testing frequency compared to traditional finger pricks, toward tracking dynamic glucose changes throughout the day. Unlike earlier studies based on stimulated sweating via exercise or iontophoresis, which reported low correlation with blood glucose, the current use of natural sweating provides high accuracy in predicting blood glucose levels in conjunction with a single, day-long, personalized calibration. Over 160 tests performed on subjects with diverse backgrounds across multiple days of glucose monitoring sessions, the sensor provides highly accurate blood glucose concentration data as indicated by a low mean absolute relative difference (MARD) of 8.15%, which compares favorably with the accuracy of commercially available blood glucose strips and CGMs (typically 9-14%). Due to the attractive features of the new user-friendly non-invasive biosensing method and its high quality data, the disclosed technology, in some embodiments, can be implemented to enable non-invasive glucose monitoring that can be a critical component of diabetes self-care.

[0082] FIG. 14 illustrates an exemplary method 1400 for measuring biomarkers in a biofluid in accordance with some embodiments of the disclosed technology.

[0083] In some implementations, a method 1400 includes placing a sensor device in contact with skin of a subject at 1402, and measuring a biomarker in a biological fluid from the skin of the subject using the sensor device at 1404. The sensor device has a plurality of first electrodes and a plurality of second electrodes, a first current collector coupled to the plurality of first electrodes at one end of each first electrode, and a second current collector coupled to the plurality of second electrodes at one end of each second electrode, the first electrodes and the second electrodes being arranged in an alternating manner, and adjacent first and second electrodes being spaced apart from each other by a predetermined distance.

[0084] Thus, various implementations of the features of the disclosed technology may be made based on the above disclosure, including the examples listed below.

[0085] Example 1. A sensor device comprising: a substrate; a plurality of first electrodes and a plurality of second electrodes formed on the substrate; a first current collector formed on the substrate and coupled to the plurality of first electrodes at one end of each first electrode; and a second current collector formed on the substrate and coupled to the plurality of second electrodes at one end of each second electrode, wherein the first electrodes and the second electrodes are arranged in an alternating manner, and adjacent first and second electrodes are spaced apart from each other by a predetermined distance.

[0086] Example 2. The sensor device of example 1, wherein the plurality of first electrodes comprises a working electrode and the plurality of second electrodes comprises a reference or counter electrode.

[0087] Example 3. The sensor device of example 1, wherein the plurality of first and second electrodes comprise at least one of a metal, a carbonaceous material, a doped conductive metal oxide, a conductive polymer, or a metal salt.

[0088] Example 4. The sensor device of example 3, wherein the metal comprises at least one of silver, gold, platinum, copper, titanium, or brass.

[0089] Example 5. The sensor device of example 3, wherein the carbonaceous material comprises at least one of graphite, carbon nanotubes, graphene, laser-induced graphene, glassy carbon, or reduced graphene oxide.

[0090] Example 6. The sensor device of example 3, wherein the doped conductive metal oxide comprises indium tin oxide (ITO).

[0091] Example 7. The sensor device of example 3, wherein the conductive polymer comprises at least one of polyaniline, polypyrrole, polythiophene, or poly(3,4-ethylenedioxythiophene):polystyrenesulfonate (PEDOT:PSS).

[0092] Example 8. The sensor device of example 3, wherein the metal salt comprises silver chloride.

[0093] Example 9. The sensor device of Example 3, wherein the plurality of first and second electrodes comprise an electroactive redox mediator comprising at least one of a quinone, Prussian blue, an osmium-containing redox mediator, a ruthenium-containing mediator, an organic dye, tetrathiafulvalene, tetrathiafulvalene-tetracyanoquinodimethane, or ferrocene.

[0094] Example 10. The sensor device of example 9, wherein the quinone comprises at least one of benzoquinone, naphthoquinone (NQ), anthraquinone, hydroquinone, or chlorohydroquinone.

[0095] Example 11. The sensor device of example 9, wherein the osmium-containing redox mediator comprises an osmium-containing redox mediator.

[0096] Example 12. The sensor device of Example 9, wherein the ruthenium-containing mediator comprises a ruthenium bipyridine.

[0097] Example 13. The sensor device of example 9, wherein the organic dye comprises at least one of methylene blue, toluidine blue O, methylene green, azure A and B, or thionin.

[0098] Example 14. The sensor device of Example 1, wherein the substrate comprises a polymeric substrate and the plurality of first and second electrodes comprises one or more solid interdigitated electrodes (IDEs) deposited on the polymeric substrate. In some implementations, deposition can be performed using a planar deposition method. In some implementations, the first and second electrodes comprise a material deposited on the substrate using at least one of screen printing, inkjet printing, flexography, sputtering, lithography, electrodeposition, or direct ink writing.

[0099] Example 15. The sensor device of example 1, wherein the substrate comprises at least one of glass, silicon, paper, a textile, a polymer plastic, or an elastomer.

[0100] Example 16. The sensor device of Example 15, wherein the substrate comprises the polymeric plastic or the elastomer and comprises one or more polymeric layers, the one or more polymeric layers comprising at least one of polyethylene, polypropylene, polyethylene terephthalate, polyvinyl chloride, polystyrene, polysaccharide-based polymers, polystyrene-based block copolymers, polysaccharides, polyvinyl alcohol, or polyethylene vinyl acetate.

[0101] Example 17. The sensor device of example 16, wherein the polystyrene-based block copolymer comprises at least one of polystyrene-polyethylene-polybutylene-polystyrene (SEBS), polystyrene-polyisoprene-polystyrene (SIS), or polystyrene-polybutylene-polystyrene (SBS).

[0102]

[0103] Example 18. The sensor device of example 16, wherein the polysaccharide comprises at least one of a starch polymer, cellulose, nitrocellulose, chitosan, ethylcellulose, or methylcellulose.

[0104] Example 19. The sensor device of claim 1, wherein the substrate comprises at least one of a fluorinated polymer, a copolymer of the fluorinated polymer, poly(vinyldifluoroethylene), tetrafluoropropylene, or hexafluoropropylene.

[0105] Example 20. The sensor device of claim 19, wherein the copolymer of the fluorinated polymer comprises poly(tetrafluoroethylene).

[0106] Example 21. The sensor device of Example 1, wherein the sensor device is configured to be placed directly on the skin surface to directly interact with natural sweating without a bio-fluid collection mechanism including microfluidics or hydrogels.

[0107] Example 22. The sensor device of Example 21, wherein the biological fluid comprises fingertip sweat.

[0108] Example 23. A sensor device as described in Example 1, wherein the predetermined distance between the adjacent first and second electrodes includes an ionic pathway for signal transmission upon contact with a user's skin surface due to natural sweating, and the ionic pathway is established only upon contact with the skin and is removed upon removal of the sensor from the skin surface.

[0109] Example 24. The sensor device of example 1, wherein the predetermined distance is less than 1 mm.

[0110] Example 25. The sensor device of Example 1, wherein the plurality of first electrodes and the plurality of second electrodes are interdigitated with each other, and the interdigitated electrodes are arranged in parallel, radial, or concentric circles.

[0111] Example 26. The sensor device of Example 1, further comprising at least one of an enzymatic or non-enzymatic conversion layer formed on the first and second electrodes and the first and second current collectors for selectively reacting with a biomarker in the bodily fluid for detection.

[0112] Example 27. The sensor device of example 26, wherein the enzyme conversion layer comprises an enzyme material comprising at least one of glucose oxidase, glucose dehydrogenase, lactate oxidase, lactate dehydrogenase, alcohol oxidizer, alcohol dehydrogenase, tyrosinase, ascorbic acid oxidase, urease, uricase, xanthine oxidase, tyrosinase, glutamate oxidase, laccase, hydroxybutyrate dehydrogenase, catalase, or bilirubin oxidase.

[0113] Example 28. The sensor device of Example 26, wherein the non-enzymatic transduction layer comprises an electroactive non-enzymatic material comprising at least one of a metal nanoparticle, a metal, a nanozyme, or an ion-selective material.

[0114] Example 29. The sensor device of example 28, wherein the metal nanoparticles include at least one of gold nanoparticles, platinum nanoparticles, or silver nanoparticles.

[0115] Example 30. The sensor device of example 28, wherein the metal comprises at least one of copper or nickel.

[0116] Example 31. The sensor device of example 28, wherein the ion-selective material comprises at least one of a hydrogen ionophore, a sodium ionophore, and a potassium ionophore.

[0117] Example 32. The sensor device of Example 26, wherein at least one of the enzyme conversion layer or the non-enzyme conversion layer includes at least one of an enzyme cofactor or a cross-linking chemical deposited under or on at least one of the enzyme conversion layer or the non-enzyme conversion layer.

[0118] Example 33. The sensor device of example 32, wherein the enzyme cofactor comprises at least one of nicotinamide adenine dinucleotide, flavin adenine dinucleotide, flavin mononucleotide, heme, or ascorbic acid.

[0119] Example 34. The sensor device of example 32, wherein the cross-linking chemicals include at least one of glutaraldehyde, epoxy resin, dihydrazide, bisacrylamide, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide / N-hydroxysuccinimide (EDC / NHS), stabilizers, polymers, or surfactants.

[0120] Example 35 The sensor device of example 34, wherein the stabilizing agent comprises at least one of bovine serum albumin, human serum albumin, glycerol, or polyphenylenediamine.

[0121] Example 36. The sensor device of example 34, wherein the polymer comprises at least one of polyethyleneimine, polyurethane, polyvinyl alcohol, polyvinylpyrrolidone, chitosan, or Nafion.

[0122] Example 37. The sensor device of example 34, wherein the surfactant comprises at least one of sodium dodecyl sulfate, sodium dodecylstyrene sulfonate, Triton X-100, Triton X-114, or Tween 80.

[0123] Example 38. The sensor device of Example 1, further comprising a signal conversion layer disposed on the first and second electrodes and functionalizing the first and second electrodes for a biomarker, and a protective layer disposed on the signal conversion layer and protecting the signal conversion layer.

[0124] Example 39. The sensor device of Example 38, wherein the protective layer comprises at least one of Nafion, chitosan, polyethyleneimine, polyurethane, polyvinyl alcohol, polyvinyl chloride, poly(vinyldifluoroethylene), poly(tetrafluoroethylene), tetrafluoropropylene, hexafluoropropylene, starch polymers, cellulose, nitrocellulose, chitosan, ethylcellulose, or methylcellulose.

[0125] Example 40. The sensor device of example 1, wherein the sensor device is configured to selectively react with a biomarker in a bodily fluid for detection, the biomarker comprising at least one of glucose, lactate, alcohol, levodopa, creatinine, urea, uric acid, bilirubin, hydroxybutyrate, vitamins, oxygen, ions, hormones, opioids, or cannabinoids.

[0126] Example 41. The sensor device of example 40, wherein the vitamin comprises ascorbic acid.

[0127] Example 42. The sensor device of example 40, wherein the ions include at least one of protons, sodium, potassium, chloride, fluoride, calcium, zinc, lead, cadmium, or mercury.

[0128] Example 43. The sensor device of example 40, wherein the hormone comprises at least one of cortisol, adrenaline, or insulin.

[0129] Example 44. The sensor device of Example 1, wherein the sensor device is integrated with an additional sensor for measuring a physical parameter when in contact with the skin, the physical parameter including at least one of temperature, moisture, or pressure.

[0130] Example 45. The sensor device of Example 1, wherein the sensor device is integrated with an additional sensor for measuring a physiological parameter when in contact with the skin, the physiological parameter including blood oxygen level, heart rate, fingerprint pattern, or blood pressure.

[0131] Example 46. A sensor device comprising a plurality of electrode arrays for simultaneous or sequential sensing of a plurality of biomarkers of a physiological parameter, each of the electrode arrays comprising: a plurality of first electrodes and a plurality of second electrodes; a first current collector coupled to the plurality of first electrodes at one end of each first electrode; and a second current collector coupled to the plurality of second electrodes at one end of each second electrode, the first electrodes and the second electrodes being arranged in an alternating manner, with adjacent first and second electrodes being spaced apart from each other by a predetermined distance.

[0132] Example 47. An array of sensors comprising a plurality of sensor devices according to any of Examples 1-46, the plurality of sensor devices being formed on a substrate for simultaneously or sequentially sensing a plurality of biomarkers in a biological fluid from different locations on the body.

[0133] Example 48. A method comprising: placing a sensor device according to any one of claims 1 to 47 in contact with the skin of a subject; and using the sensor device to measure a biomarker in a biological fluid from the skin of the subject.

[0134] Example 49. The method of Example 48, wherein the biomarker comprises at least one of glucose, lactate, alcohol, levodopa, creatinine, urea, uric acid, bilirubin, hydroxybutyrate, vitamins, oxygen, ions, hormones, opioids, or cannabinoids.

[0135] Example 50. The method of Example 48, further comprising measuring at least one of a physical parameter or a physiological parameter together with the biomarker, the physical parameter comprising at least one of temperature, moisture, or pressure, and the physiological parameter comprising at least one of skin resistance, temperature, blood oxygen level, heart rate, fingerprint pattern, or blood pressure.

[0136] Implementations of the subject matter and functional operations described in this patent document may be implemented in various systems, digital electronic circuits, or computer software, firmware, or hardware, including the structures disclosed herein and their structural equivalents, or in one or more combinations thereof. Implementations of the subject matter described herein may be implemented as one or more computer program products, i.e., one or more modules of computer program instructions encoded on a tangible and non-transitory computer-readable medium for execution by or for controlling the operation of a data processing device. The computer-readable medium may be a machine-readable storage device, a machine-readable storage substrate, a memory device, a composition of matter that provides a machine-readable propagated signal, or a combination of one or more of them. The term "data processing unit" or "data processing device" encompasses all apparatus, devices, and machines for processing data, including, by way of example, a programmable processor, a computer, or multiple processors or computers. In addition to hardware, an apparatus may include code that creates an execution environment for the computer program in question, such as code constituting a processor firmware, a protocol stack, a database management system, an operating system, or one or more combinations of them.

[0137] A computer program (also known as a program, software, software application, script, or code) can be written in any form of programming language, including compiled or interpreted languages, and can be deployed in any form, including as a stand-alone program or as a module, component, subroutine, or other unit suitable for use in a computing environment. A computer program does not necessarily correspond to a file in a file system. A program can be stored as part of a file that holds other programs or data (e.g., one or more scripts stored in a markup language document), in a single file dedicated to the program in question, or in multiple cooperating files (e.g., files that store one or more modules, subprograms, or portions of code). A computer program can be deployed to be executed on one computer, or on multiple computers located at one site or distributed across multiple sites and interconnected by a communication network.

[0138] The processes and logic flows described herein may be performed by one or more programmable processors executing one or more computer programs to perform functions by operating on input data and generating output. The processes and logic flows may also be performed by, and an apparatus may be implemented as, special purpose logic circuitry, such as an FPGA (field programmable gate array) or an ASIC (application specific integrated circuit).

[0139] Processors suitable for executing computer programs include, by way of example, both general-purpose and special-purpose microprocessors, as well as any one or more processors of any kind of digital computer. Generally, a processor receives instructions and data from a read-only memory or a random-access memory or both. The essential elements of a computer are a processor for executing instructions and one or more memory devices for storing instructions and data. Generally, a computer is also operatively coupled to receive data from, transfer data to, or both of, one or more mass storage devices for storing data, e.g., magnetic, magneto-optical, or optical disks. However, a computer need not have such devices. Computer-readable media suitable for storing computer program instructions and data include, by way of example, all forms of non-volatile memory, media, and memory devices, including semiconductor memory devices, e.g., EPROM, EEPROM, and flash memory devices. The processor and memory may be supplemented by, or incorporated in, special-purpose logic circuitry.

[0140] It is intended that the specification, together with the drawings, be considered merely illustrative, with exemplary means being illustrative. As used herein, the singular forms "a," "an," and "the" are intended to include the plural forms as well, unless the context clearly dictates otherwise. Additionally, the use of "or" is intended to include "and / or" unless the context clearly dictates otherwise.

[0141] Although this patent document contains many details, these should not be construed as limiting the scope of any invention or what may be claimed, but rather as descriptions of features that may be specific to certain embodiments of a particular invention. Certain features described in this patent document in the context of separate embodiments may also be implemented in combination in a single embodiment. Conversely, various features described in the context of a single embodiment may also be implemented in multiple embodiments separately or in any suitable subcombination. Furthermore, features may be described above as acting in a particular combination, and even as initially claimed, one or more features from a claimed combination may in some cases be carved out of the combination, and the claimed combination may be subject to subcombinations or variations of the subcombination.

[0142] Similarly, although operations are depicted in the figures in a particular order, this should not be understood as requiring that such operations be performed in the particular order shown, or in sequential order, or that all of the illustrated operations be performed, to achieve desired results. Further, the separation of various system components in the embodiments described in this patent document should not be understood as requiring such separation in all embodiments.

[0143] Only some implementations and examples are described; other implementations, extensions, and variations can be made based on what is described and illustrated in this patent document.

Claims

1. 1. A glucose monitoring device for detecting glucose in a biological fluid sample from a skin surface, comprising: a) a substrate including a conductive layer, the substrate comprises at least one of glass, silicon, paper, textile, polymer, or elastomer; the conductive layer comprises platinum, silver, gold, copper, titanium, ruthenium, palladium, rhodium, brass, or a combination thereof; substrate, b) a first electrode comprising a first electrode material deposited on the conductive layer, the first electrode material comprises one or more of glutaraldehyde, epoxy, dihydrazide, bisacrylamide, 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide / N-hydroxysuccinimide (EDC / NHS)), stabilizer, polymer, or surfactant; the first electrode material further comprising a transducer configured to generate an electrochemical signal from a reaction with the glucose from the biological fluid sample. a first electrode; c) a second electrode comprising a second electrode material deposited on the conductive layer, the first electrode includes a first plurality of engagement portions, and the second electrode includes a second plurality of engagement portions; the first plurality of fitting portions are alternately arranged with the second plurality of fitting portions; the electrochemical signal is used to provide a blood glucose concentration of the user. a second electrode, 1. A device comprising:

2. 10. The device of claim 1, wherein the polymer comprises at least one of polyethylene, polypropylene, polyethylene terephthalate, polyvinyl chloride, polystyrene, a polysaccharide-based polymer, a polystyrene-based block copolymer, a polysaccharide, polyvinyl alcohol, or polyethylene vinyl acetate.

3. The device of claim 1 , wherein the transducer is glucose oxidase.

4. The second electrode material is (a) an iron-based redox couple, or (b) one or more of polyaniline, polypyrrole, polythiophene, or poly(3,4-ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS); Including, The device of claim 1.

5. The device of claim 1 , wherein each of the first and second plurality of mating portions is spaced apart by a predetermined distance of about 1 mm or less.

6. The device of claim 1 , wherein the alternating first and second pluralities of mating portions are interdigitated with each other.

7. The device of claim 1 , wherein the biological fluid comprises sweat and the biological fluid sample from the skin surface is collected by directly contacting a fingertip with the device.

8. The device of claim 1 , wherein the device comprises a second layer, the second layer comprising one or more of chitosan, agarose, or polyethyleneimine.

9. The device of claim 1 , wherein the device comprises a third layer, the third layer comprising one or more of Nafion or a cellulosic material.

10. 10. The device of claim 1, wherein the device comprises a protective layer, the protective layer comprising one or more of Nafion, polyethyleneimine, polyurethane, polyvinyl alcohol, polyvinyl chloride, poly(vinyldifluoroethylene), poly(tetrafluoroethylene), tetrafluoropropylene, hexafluoropropylene, starch polymers, cellulose, nitrocellulose, chitosan, ethylcellulose, or methylcellulose.

Images