Epidermal biosensor
By designing epidermal biosensors that include diffusion layers, enzyme bioreceptors, transducers, processors and substrates, the problem of difficulty in achieving continuous and low-cost epidermal biomarkers in the prior art is solved, and the continuous and real-time detection of epidermal solid phase biomarkers is achieved, and a non-invasive and low-cost health monitoring solution is provided.
Patent Information
- Application Number
- CN202380032654.X
- Authority / Receiving Office
- CN · China
- Patent Type
- Applications(China)
- Current Assignee / Owner
- Priority Date
- 2022-03-08
- Filing Date
- 2023-03-02
- Publication Date
- 2025-04-15
AI Technical Summary
The prior art is difficult to achieve continuous, low-cost monitoring of epidermal biomarkers, and traditional methods are highly invasive, uncomfortable and unable to provide continuous monitoring.
An epidermal biosensor was designed, including a diffusion layer, enzyme biological receptor, transducer, processor and substrate. Through the diffusion layer, solid-phase epidermal analytes are dissolved, enzyme biological receptors oxidize these analytes, transducer and diffusion layer work at the interface, and the processor processes electrochemical data to realize solid-phase electrochemical sensing.
Continuous and real-time detection of epidermal solid phase biomarkers is achieved, and a non-invasive and low-cost health monitoring program is provided, which eliminates the need for sweat-induced, can be worn at home, and supports remote health monitoring.
Smart Images
Figure CN119836257A_ABST
Abstract
Description
Technical Field
[0001] The present invention relates generally to biochemical data monitoring, and more particularly, to epidermal biosensors. Background Art
[0002] Monitoring biochemical data is critical for medical diagnosis, disease monitoring, and health management. The invasiveness and form factor of sensors are two (2) limiting factors that hinder the continuous monitoring of important biomarkers. For example, blood tests are invasive, non-continuous, expensive, and require a professional medical setting.
[0003] Recent research efforts on non-invasive sensing modalities have focused on alternative biofluids, such as human sweat. It has been proposed that continuous measurement of biomarker concentrations in sweat could be used for fitness tracking and chronic disease monitoring. However, obtaining sweat from patients and physically inactive individuals is difficult. Typical sweat induction methods include iontophoresis or administration of drugs (e.g., pilocarpine). However, long-term use of sweat induction methods can cause patient discomfort and other side effects, including headaches, changes in skin pH, lesions, diarrhea, and flu-like symptoms.
[0004] Solid-phase biomarkers can be used to diagnose and monitor chronic diseases. For example, epidermal solid-phase glucose is positively correlated with diabetes and may be more sensitive than urine and capillary blood obtained by finger pricking. Similarly, epidermal solid-phase cholesterol can be used as an important biomarker for hyperlipoproteinemia, coronary artery disease, and atherosclerosis. Lactic acid from insensible perspiration provides valuable health information for cardiovascular disease. However, current detection methods rely on offline skin sample collection and require the use of advanced instruments such as mass spectrometry (MS) or high-performance liquid chromatography (HPLC), which are difficult to miniaturize for remote patient monitoring and cannot provide continuous monitoring.
[0005] In view of the above, it is desirable to provide low-cost epidermal biosensors capable of sustained solid-phase electrochemical sensing. Summary of the invention
[0006] Thus, in a first aspect, the present invention provides an epidermal biosensor. The epidermal biosensor comprises: a diffusion layer operable to dissolve a solid-phase epidermal analyte; an enzyme bioreceptor operable to oxidize the dissolved epidermal analyte from the diffusion layer; a transducer having an interface with the diffusion layer; a processor configured to process electrochemical data from the transducer; and a substrate to which the enzyme bioreceptor and the transducer are attached.
[0007] Other aspects and advantages of the present invention will become apparent from the following detailed description taken in conjunction with the accompanying drawings, which illustrate, by way of example, the principles of the invention. BRIEF DESCRIPTION OF THE DRAWINGS
[0008] Embodiments of the present invention will now be described, by way of example only, with reference to the accompanying drawings, in which:
[0009] Figure 1 is an exploded schematic diagram showing an epidermal biosensor according to an embodiment of the present invention;
[0010] Figure 2 It is shown Figure 1 A cross-sectional schematic diagram of the operation of the epidermal biosensor;
[0011] Figure 3 is a photograph of an epidermal biosensor according to an embodiment of the present invention attached to human skin;
[0012] Figure 4 It is shown Figure 1 Schematic diagram of the application of epidermal biosensors;
[0013] Figure 5A is a graph showing sensor sensitivity and operating time at different glycerol / water ratios for characterizing a solid-phase lactic acid interface;
[0014] Figure 5B is a graph showing measurements of measured current over time using a lactate oxidase based sensor;
[0015] Figure 5C is a graph showing a calibration curve of current density versus area density for solid phase lactate;
[0016] Figure 5D is a diagram of selective characterization of solid-phase lactic acid interface using chronoamperometry;
[0017] Fig. 6A is a graph showing sensor sensitivity and operating time at different gelatin ratios for characterizing solid-phase cholesterol interfaces;
[0018] Figure 6B is a graph showing measurements of measured current over time using a cholesterol oxidase based sensor;
[0019] Figure 6C is a graph showing a calibration curve of current density versus area density of solid-phase cholesterol.
[0020] Fig.6D is a graph showing the selective characterization of a solid cholesterol interface using chronoamperometry.
[0021] Fig. 7Aand 7B is a graph providing a comparison of lactate predicted by a solid phase electrochemical sensor and lactate measured by a commercial colorimetric assay kit; and
[0022] Figure 7C and 7D is a graph providing a comparison of cholesterol predicted by a solid phase electrochemical sensor and cholesterol measured by a commercial colorimetric assay kit.
[0023] Detailed Description of Exemplary Embodiments
[0024] The detailed description given below in conjunction with the accompanying drawings is intended to describe the present preferred embodiments of the present invention, and is not intended to represent the only form in which the present invention can be implemented. It should be understood that the same or equivalent functions can be achieved by different embodiments intended to be included within the scope of the present invention.
[0025] As used herein, the term "diffusion layer" refers to the thickness of a porous material that allows species to move through the porous material based on concentration differences.
[0026] As used herein, the term "enzyme bioreceptor" refers to a biological substance that is capable of selectively catalyzing a reaction.
[0027] As used herein, the term "hydrogel material" refers to a three-dimensional (3D) network of hydrophilic polymers that swells in water and holds large amounts of water while maintaining its structure due to chemical or physical cross-linking of individual polymer chains.
[0028] As used herein, the term "about" refers to both a number within a numerical range and a value including the standard deviation of error for the device or method used to determine the value. As used herein, the term "about" can allow for a certain degree of variation in a value or range, for example, within 10%, within 5%, or within 1% of a specified value or a specified range limit.
[0029] Reference now Figure 1 , an epidermal biosensor 10 is shown. The epidermal biosensor 10 includes a diffusion layer 12 operable to dissolve a solid-phase epidermal analyte, an enzyme bioreceptor 14 operable to oxidize the dissolved epidermal analyte from the diffusion layer 12, a transducer 16 having an interface with the diffusion layer 12, a processor 18 configured to process electrochemical data from the transducer 16, and a substrate 20 to which the enzyme bioreceptor 14 and the transducer 16 are attached. In the illustrated embodiment, an adhesive layer 22 encapsulates a portion of the substrate 20. A plurality of interconnects 24 may be provided to electronically connect the transducer 16 to the processor 18.
[0030] The various elements of the epidermal biosensor 10 can be made of flexible and / or stretchable or elastic materials. The epidermal biosensor 10 provides a stretchable biochemical interface for epidermal solid phase biomarkers. The fully integrated and wearable stretchable electrochemical sensor 10 is capable of continuous detection of biochemical substances in the solid phase. The epidermal biosensor 10 can be further integrated with batteries and mobile applications.
[0031] When in use, the diffusion layer 12 is in direct contact with human skin and solvates the solid-phase biomolecules (epidermal biomolecules) present on human skin. The solvation-diffusion process of the solid-phase analyte occurs in the diffusion layer 12—the solvation-diffusion layer 12 dissolves the solid analyte and allows the dissolved analyte to diffuse to the enzyme bioreceptor 14 and the transducer 16.
[0032] The diffusion layer 12 may comprise a matrix or three-dimensional network of a hydrophilic polymer such as a hydrogel material. In order to allow electrochemical reactions to occur, the hydrogel material is designed and modified to solvate the solid phase molecules, thereby allowing subsequent diffusion. More specifically, the hydrogel material has a dual function, namely, solvating the solid phase analyte and promoting the transmission from the skin to the sensing interface; the solid analyte is solvated and diffused in the hydrogel material. A phase change from solid to liquid occurs in the hydrogel material. The hydrogel material can be formulated and designed for the solvation process from solid to liquid and the diffusion process of the target solid phase analyte. The solvation-diffusion layer 12 made of the hydrogel material allows the solid-to-liquid conversion of the solid analyte within the hydrogel matrix, thereby enabling the electrochemical quantification of the surface density of the solid phase biomolecules (epidermal biomolecules) present on the human skin. Advantageously, due to the use of the hydrogel material, the epidermal biosensor 10 is less affected by motion artifacts. The stability of the electrochemical interface provided by the epidermal biosensor 10 with the hydrogel is also higher than the liquid interface (sweat) on the dynamically stretched skin. In one embodiment, the hydrogel material 12 may be composed of an agarose hydrogel at between about 0.1% weight / volume (% w / v) and about 4% w / v. If the proportion of agarose in the hydrogel is too high, diffusion of ions and analytes may be hindered, resulting in poor sensing results.
[0033] The working time of the diffusion-solvation layer 12 can be extended by introducing a high boiling point biocompatible additive (e.g., glycerol). Glycerol can be added to the hydrogel to reduce or slow the water evaporation rate and achieve a long working time without seriously reducing the sensor performance. In order to improve the mechanical strength of the hydrogel and reduce the water evaporation rate of the hydrogel, gelatin can be added in different weight percentages. Therefore, in order to provide a longer working time, the diffusion layer 12 can further include glycerol and / or gelatin. In one embodiment, the hydrogel material 12 can be composed of a gelatin hydrogel between about 0.1% w / v and about 10% w / v.
[0034] For water-insoluble biomarkers, additives such as ethanol and / or 2-[4-(2,4,4-trimethylpentan-2-yl)phenoxy]ethanol (Triton X-100) may be added to increase solubility in the diffusion layer 12, thereby promoting solvation and transport of hydrophobic analytes. Therefore, the diffusion layer 12 may further include a surfactant. For hydrophobic species, a surfactant may be added to the hydrogel. The surfactant may include 2-[4-(2,4,4-trimethylpentan-2-yl)phenoxy]ethanol (Triton X-100) at between about 0.1% w / v and about 10% w / v. The diffusion layer 12 may further include ethanol at between about 0.1% w / v and about 10% w / v. For example, the hydrogel may be modified with a certain amount of ethanol (2%) and a surfactant (2% Triton-X 100) for water-insoluble biomarkers.
[0035] If the diffusion layer 12 is too thick, diffusion may take too long. However, if the diffusion layer 12 is too thin, water tends to evaporate more quickly from the diffusion layer 12 at body temperature. In order to optimize the epidermal biosensor 10 to allow for fast response and extended operating time at the same time, the thickness of the diffusion layer 12 can be between about 300 micrometers (μm) and about 1.5 millimeters (mm).
[0036] Through the design of the solvation-diffusion layer 12 , the epidermal biosensor 10 can be used to detect water-soluble analytes (eg, lactate) and water-insoluble analytes (eg, cholesterol) to monitor chronic diseases, such as cardiovascular disease.
[0037] The enzyme bioreceptor 14 may be provided as an enzyme functionalized layer where a native enzyme oxidizes the analyte and produces hydrogen peroxide. The enzyme layer or enzyme bioreceptor 14 provides a layer of enzymes that convert target biomolecules into other molecules allowing measurement of electron transfer.
[0038] The transducer 16 may include screen-printed electrodes. The electrochemical reaction occurs at the hydrogel-electrode interface. To provide a screen-printed stretchable electrode, the electrode may include graphite and poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS). The electrode may further include at least one selected from the following: hexacyanoferrate (II, III) iron (II, III) (Prussian blue (PB)), aqueous polyurethane, dimethyl sulfoxide (DMSO) and (3-glycidyloxypropyl) trimethoxysilane (GPTMS). Advantageously, Prussian blue (PB) can be used as a redox mediator to reduce the overpotential required for detecting hydrogen peroxide. Advantageously, the inclusion of viscoelastic WPU as a secondary polymer network helps to improve the stability and sensitivity of the electrode, and also helps to improve the stretchability of the epidermal biosensor 10. The introduction of stretchable WPU makes the printed electrode ink / paste stretchable, thereby enabling the stretchable version of the electrode to withstand 30% strain without degrading performance. DMSO helps to improve the conductivity of the electrode material, while GPTMS helps to improve the stability of the electrode material in an aqueous environment.
[0039] The processor 18 can be integrated with a printed circuit board (PCB) or a flexible printed circuit board (fPCB), and the front-end circuit on the PCB can collect and process electrochemical data.
[0040] Substrate 20 is elastic in nature and may be made from a polymer such as styrene-ethylene-butylene-styrene (SEBS).
[0041] The interconnect 24 may be a silver-based conductive trace formed by screen printing, and may be made of a polymer material having stretchability, thereby enabling the interconnect 24 to withstand a large amount of strain without breaking.
[0042] Reference now Figure 2 , showing a cross section of the working electrode, illustrating the three (3) steps of the epidermal solid analyte, specifically solvation, diffusion, and electrochemical reaction. Figure 2 It can be seen that the electrochemical reaction occurs at the hydrogel-electrode interface.
[0043] Reference now Figure 3 , the figure shows an epidermal biosensor 10 placed on human skin and conforming to the curved surface of the user's forearm. The epidermal biosensor 10 is stretchable and can adapt to body movements without causing discomfort to the user and severe motion artifacts. Figure 3 As shown in , the solvation-diffusion layer surrounded by the dashed circle is in direct contact with the stratum corneum to allow the delivery of solid-phase analytes to the electrode surface.
[0044] Reference now Figure 4, shows a use case of the epidermal biosensor 10 providing a stretchable biochemical interface for solid epidermal analytes. Figure 4 As can be seen in the figure, the epidermal biosensor 10 can be placed on the skin for detecting epidermal solid-phase biomarkers, including water-insoluble cholesterol and water-soluble lactic acid. The epidermal biosensor 10 can be wirelessly connected to a smartphone 50 to read out the surface density of the solid-phase biomarkers. The electrochemical data from the epidermal biosensor 10 can be sent to the smartphone 50 via Bluetooth technology. The data can then be displayed to the user, family members, and caregivers to track chronic diseases. Advantageously, the epidermal biosensor 10 is capable of in situ detection of solid-phase biomarkers including dry sweat and secreted sebum. More advantageously, the stretchable solid-phase sensor 10 eliminates the need to induce sweating through drugs or exercise, and provides a comfortable and non-invasive interface to reliably obtain practical health signals outside of a hospital or clinical environment. Example
[0045] Example 1
[0046] The sensing performance of the electrochemical sensor solid-phase lactic acid interface was experimentally characterized. A thin layer of hydrogel made of 2wt% low-melting-point agarose was selected to constitute the solvation-diffusion layer due to its fast diffusion response, low cost, and enzyme compatibility. Glycerol was added to the hydrogel to reduce the water evaporation rate and achieve a long working time.
[0047] Reference now Figure 5A , showing the optimization of sensor sensitivity and operating time (limited by water evaporation) at different glycerol / water ratios for characterizing solid-phase lactic acid interfaces. Figure 5A As can be seen in the graph, the working time increases from 1.8 h in pure water-based hydrogels (without glycerol) to 4.2 h in hydrogels with a glycerol / water ratio of 0.1:1. However, glycerol has an adverse effect on sensitivity, probably due to the hindrance of lactic acid diffusion to the hydrogel-electrode interface. As a trade-off, the sensitivity of the sensor increases from 32.5 μA μmol -1 cm -2 Reduced to 10.3 μA μmol at a glycerol / water ratio of 0.1:1 -1 cm -2 Excess glycerol further hampered the sensitivity of the sensor, and the optimal glycerol / water ratio adopted throughout the subsequent experiments was 0.1:1.
[0048] Reference now Figure 5B , shows the measurement of the measured current using a lactate oxidase based sensor. Prussian blue was used as a redox mediator.
[0049] Reference now Figure 5C, shows the calibration curve of current density versus area density of solid lactic acid. Figure 5C As shown in , the Pearson correlation coefficient r is 0.98. Figure 5C The inset of Figure 2 shows the sensing interface at the solvation-diffusion layer and solid lactic acid. The area density was calculated by dividing the contact area of solid lactic acid by the sensing interface.
[0050] Figure 5B and 5C The detection of the measured current of solid-phase lactic acid is shown. 2 and 2593.6nmol / cm 2 The linearity of lactate between the two groups was good, with a Pearson correlation coefficient of 0.98 and a detection limit (LOD) of 2.5 nmol / cm 2 .
[0051] Reference now Figure 5D , showed selectivity characterization using chronoamperometry, where interfering molecules (i.e., urea, uric acid (UA), glucose, ascorbic acid (AA), cholesterol) in the solid phase were applied to the solid phase lactate sensor. The solid phase lactate sensor showed excellent selectivity over other solid surface analytes (i.e., urea, uric acid (UA), glucose, ascorbic acid (AA), and cholesterol) in physiologically relevant amounts.
[0052] Example 2
[0053] For water-insoluble biomarkers, additives such as ethanol and Triton X-100 were used to increase solubility in the hydrogel. For the solid-phase cholesterol sensor, 2v / v% ethanol and 2v / v% Triton X-100 were dissolved into the low-melting point agarose solution. In order to further improve the mechanical strength of the hydrogel and reduce the water evaporation rate of the hydrogel, different weight percentages of gelatin were added.
[0054] Reference now Fig. 6A , showing the optimization of sensor sensitivity and operating time (limited by water evaporation) at different gelatin ratios for characterizing solid-phase cholesterol interfaces. Fig. 6A As can be seen in Figure 2, the sensitivity of the cholesterol sensor is stable from 0% to 2%, and drops sharply at 4%. Therefore, the use of 2% gelatin achieved a relatively long working time (4.8 hours of continuous amperometric detection) and good sensitivity in subsequent experiments.
[0055] Reference now Figure 6B , shows the measurement of the measured current using a cholesterol oxidase based sensor. Prussian blue was used as a redox mediator.
[0056] Reference now Figure 6C, shows the calibration curve of current density versus area density of solid phase cholesterol. Figure 6C As shown in , the Pearson correlation coefficient r for both linear regions is 0.99. Figure 6C The inset shows the sensing interface at the solvation-diffusion layer and solid cholesterol. The area density was calculated by dividing the contact area at the sensing interface by the solid cholesterol. 2 and 47.1 nmol / cm 2 Between and at 47.1 nmol / cm 2 and 684.1 nmol / cm 2 Two linear regions were observed between the two samples, with a LOD of 2.5 nmol / cm 2 .
[0057] Reference now Fig.6D , showed selectivity characterization using chronoamperometry, where interfering molecules (i.e., urea, UA, glucose, dopamine (DA), AA, lactate) were tested on a solid-phase cholesterol sensor. The device showed excellent selectivity for other solid-phase epidermal analytes (e.g., urea, UA, glucose, ascorbyl dopamine, and lactate).
[0058] The responses of solid-phase lactate and cholesterol can be recorded via a printed circuit board (PCB) and wirelessly transmitted to a smartphone via Bluetooth. Although only cholesterol and lactate were demonstrated, other epidermal biomarkers such as glucose can be detected using the same device structure and sensing mechanism.
[0059] Example 3
[0060] The SPE sensor was then tested using ex vivo human samples. The samples were prepared by dropping 5 μl of human sweat and then allowing it to dry at ambient temperature. For the ex vivo validation of the SPE sensor, the sensor accuracy was quantified by comparing the SPE sensor with a commercial assay kit.
[0061] Reference now Fig. 7A and 7B , a comparison of lactate predicted by a solid phase electrochemical sensor and lactate measured by a commercial colorimetric assay kit is provided. The Pearson correlation coefficient r is 0.90. Figure 7B The dashed lines in the graph represent confidence intervals (95%).
[0062] Reference now Figure 7C and 7D , cholesterol predicted by the solid phase electrochemical sensor and cholesterol measured by a commercial colorimetric assay kit were compared. The Pearson correlation coefficient r was 0.82. Fig.7D The dashed lines in the graph represent confidence intervals (95%).
[0063] For lactic acid (such as from Fig. 7A and 7B for cholesterol (as can be seen from Figure 7C and 7D As can be seen in the figure, a positive correlation was observed between the biomarker amounts determined by the commercial kit and the biomarker amounts determined by the solid phase sensor, with Pearson correlation coefficients of 0.90 and 0.82 for lactate and cholesterol, respectively. In vitro testing showed that the sensor has good reliability for quantification of human biological samples in the solid phase.
[0064] As is apparent from the above discussion, the present invention provides epidermal biosensors that allow for continuous detection of epidermal solid analytes on human skin. In addition to measuring liquid analytes in specific biological fluids (e.g., sweat, urine, tissue fluid, etc.), the epidermal biosensors of the present invention are also capable of measuring solid analytes produced by dry sweat or insensible sweating, which is clinically relevant in predicting and monitoring chronic diseases. The present invention provides a sensing method for measuring solid biomarkers present on the surface of human skin, thereby eliminating the need for sweat induction or physical exercise. Advantageously, using the present invention, practical and actionable health signals can be easily measured simply by placing a stretchable wearable patch on the skin, thereby allowing remote health monitoring, preventive medicine, and telemedicine. The epidermal biosensor of the present invention is non-invasive, low-cost, and can be worn at home. Compliance can be improved by eliminating the need for exercise. Advantageously, the epidermal biosensor of the present invention is capable of providing real-time, in-situ, and continuous detection of solid epidermal analytes without the need for invasive (e.g., blood collection) or additional medical procedures (e.g., sweat induction and urine collection). The epidermal biosensor of the present invention also provides a universal platform for detecting hydrophilic and hydrophobic biomarkers. Further advantageously, the skin-integrated, stretchable, wearable biochemical sensor of the present invention does not require complex equipment such as mass spectrometry or / and liquid chromatography.
[0065] Although the preferred embodiments of the present invention have been described, it is apparent that the present invention is not limited to the described embodiments. For those skilled in the art, many modifications, changes, variations, substitutions and equivalents will be apparent without departing from the scope of the invention described in the claims.
[0066] Furthermore, unless the context clearly requires otherwise, throughout the specification and claims, the words "comprise," "comprising," etc. are to be interpreted in an inclusive rather than an exclusive or exhaustive sense; that is, in the sense of "including but not limited to."
Claims
1. An epidermal biosensor comprising: a diffusion layer operable to solubilize a solid phase epidermal analyte; an enzyme bioreceptor operable to oxidize a dissolved epidermal analyte from the diffusion layer; a transducer having an interface with the diffusion layer; a processor configured to process electrochemical data from the transducer; and A substrate is provided to which the enzyme bioreceptor and the transducer are attached.
2. The epidermal biosensor of claim 1, wherein the diffusion layer comprises a hydrophilic polymer matrix.
3. The epidermal biosensor of claim 1 or 2, wherein the diffusion layer comprises a hydrogel material.
4. The epidermal biosensor of claim 3, wherein the hydrogel material consists of between about 0.1% weight / volume (% w / v) and about 4% w / v agarose hydrogel.
5. The epidermal biosensor of any one of the preceding claims, wherein the diffusion layer further comprises at least one selected from glycerol and gelatin.
6. The epidermal biosensor of claim 3, wherein the hydrogel material consists of between about 0.1% w / v and about 10% w / v gelatin hydrogel.
7. The epidermal biosensor of any of the preceding claims, wherein the diffusion layer further comprises a surfactant.
8. The epidermal biosensor of claim 7, wherein the surfactant comprises between about 0.1% w / v and about 10% w / v 2-[4-(2,4,4-trimethylpentan-2-yl)phenoxy]ethanol.
9. The epidermal biosensor of any of the preceding claims, wherein the diffusion layer further comprises between about 0.1% w / v and about 10% w / v ethanol.
10. The epidermal biosensor of any of the preceding claims, wherein the diffusion layer has a thickness between about 300 micrometers (μm) and about 1.5 millimeters (mm).
11. The epidermal biosensor of any preceding claim, wherein the transducer comprises a screen-printed electrode.
12. The epidermal biosensor of claim 11, wherein the electrode comprises graphite and poly(3,4-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS).
13. The epidermal biosensor of claim 12, wherein the electrode further comprises at least one selected from the group consisting of iron (II, III) hexacyanoferrate (II, III), aqueous polyurethane, dimethyl sulfoxide (DMSO), and (3-glycidyloxypropyl)trimethoxysilane (GPTMS).
14. The epidermal biosensor of any of the preceding claims, further comprising a plurality of interconnects electronically connecting the transducer to the processor.
15. The epidermal biosensor of any of the preceding claims, wherein the substrate comprises styrene-ethylene-butylene-styrene (SEBS).
16. The epidermal biosensor of any preceding claim, further comprising an adhesive layer encapsulating a portion of the substrate.