Wearable sensor patch

Abstract

A wearable sensor patch including a plurality of microneedles, a hydrogel body, and a flexible sensor element with elastomeric electrodes, including methods of manufacturing such a wearable sensor patch. The sensor patch allows for minimally invasive collection of ISF from the skin, while simultaneously measuring biomarker levels in-situ. The sensor patch's stretchability accommodates swelling of the hydrogel body during ISF extraction and adapts to changes in skin deformation.

Description

WEARABLE SENSOR PATCH FIELD OF THE DISCLOSURE

[0001] The present disclosure pertains to wearable sensors, and in particular to wearable sensor patches with hydrogel microneedles for in-situ analysis of interstitial fluid. BACKGROUND

[0002] Interstitial fluid (ISF) is a biofluid that originates from trans-capillary blood exchange, and its potential for in-situ biomarker analysis has gained considerable interest. It shares at least 84% of key biomarkers with plasma, including glucose and lactate, demonstrating a strong correlation between the levels of these biomarkers in ISF and blood. The in-situ analysis of ISF enables the real-time tracking of metabolic changes within tissues, thereby facilitating early detection of diseases and paving the way for personalized medical interventions. Given its direct interaction with relevant cells and tissues, ISF allows the measurement of multiple biomarkers that are either unique to it or found in higher concentrations than in conventional blood or urine tests.

[0003] Wearable in-situ sensing platforms have emerged, offering potential for real-time biomarker monitoring by extracting ISF through either ionophoresis or sonophoresis. Ionophoresis methods use skin-mounted electrodes to generate an electrical charge, inducing the movement of charged particles within ISF and facilitating neutral particle transport, such as glucose, for sampling and measurement. Sonophoresis methods apply low-frequency ultrasound to increase skin porosity through cavitation bubble induction, followed by vacuum pressure application to extract ISF for analysis. The non-invasive nature of these approaches offers valuable advantages in patient comfort and ease of use, but they encounter challenges such as potential interference with biomarkers from sweat, along with possible skin irritation. Additionally, existing devices face limitations in predicting and compensating for measurement delays, affecting the accuracy and effectiveness of real-time monitoring.

[0004] Microneedles (MNs), which come in a variety of materials and structures, offer a way to extract ISF in a minimally invasive manner by piercing the skin. However,hollow MNs face challenges, such as fracturing, complex manufacturing, empirical / iterative design, limited flexibility and durability, or brittleness. Hydrogel-based MNs constructed from materials such as polyvinyl alcohol (PVA), chitosan, osmolytes, and methacrylated hyaluronic acid (MeHA) have demonstrated their efficacy in ISF extraction, attributed to their rapid fluid extraction rates, biosafety, and streamlined synthesis procedure. Yet, integrating these hydrogel MNs with sensing modules remains challenging because of mechanical mismatches or complications arising from hydrogel swelling upon absorbing ISF. As a result, ISF is often collected using MNs and subsequently extracted and then analyzed using specialized lab equipment, facing issues such as sampling variability, time lags, sample degradation, and contamination, ultimately diminishing measurement accuracy and convenience.

[0005] Applicants appreciate the need for in-situ ISF analysis to provide real-time data on multiple target biomarkers, thus improving both accuracy and user acceptance. The present disclosure addresses these and other needs. SUMMARY

[0006] The present disclosure provides a wearable sensor patch including a plurality of microneedles, a hydrogel body, and a flexible sensor element with elastomeric electrodes. Methods of making the sensor patch and its components are also disclosed. Advantageously, the sensor patch allows for minimally invasive collection of ISF from the skin, while simultaneously measuring biomarker levels in-situ. The sensor patch’s stretchability accommodates swelling of the hydrogel body during ISF extraction and adapts to changes in skin deformation.

[0007] In one aspect of the disclosure a wearable sensor patch for monitoring interstitial fluid is disclosed. The sensor patch includes a base layer, a hydrogel body, and a sensor element embedded within the hydrogel body. The sensor element includes a conductive layer embedded within a polydimethylsiloxane (PDMS) body. The conductive layer includes a working electrode including thermoplastic polyurethane (TPU) and a first conducting material, a serpentine-shaped counter electrode comprising thermoplastic polyurethane (TPU) and a second conducting material, and areference electrode comprising thermoplastic urethane (TPU) and a third conducting material.

[0008] In another aspect of the disclosure, a method of manufacturing a wearable sensor patch for monitoring of interstitial fluid is disclosed. The method includes printing a sensor element onto a substrate, encapsulating said sensor element in a hydrogel body, and curing the hydrogel body with a plurality of hydrogel microneedles.

[0009] In various aspects of the disclosure, the patch may be configured to recoverably deform or elongate in response to user movement or swelling from the uptake of interstitial fluid, without delaminating or breaking any one of the base layer, the hydrogel body, the sensor element, and a combination thereof.

[0010] In various aspects of the disclosure, the base layer may include hydrogel microneedles. The hydrogel microneedles may be integrated with the hydrogel body such that the hydrogel microneedles and the hydrogel body form one homogenous hydrogel structure. The hydrogel microneedles and hydrogel body may both include methacrylated hyaluronic acid (MeHA).

[0011] In various aspects of the disclosure, at least one of the first conducting material and the second conducting material may include silver flakes.

[0012] In various aspects of the disclosure, the third conducting material may include at least one of silver and silver chloride.

[0013] In various aspects of the disclosure, the hydrogel body may include a material selected from the group consisting of polyvinyl alcohol (PVA), chitosan, osmolytes, and methacrylated hyaluronic acid (MeHA).

[0014] In various aspects of the disclosure, the wearable sensor may include an enzyme layer which contacts the working electrode, counter electrode, and reference electrodes. The enzyme layer may be formed from a solution of oxidase enzyme, bovine serum albumin (BSA), and a crosslinking agent.

[0015] In various aspects of the disclosure, curing the hydrogel body with the plurality of hydrogel microneedles may cause the hydrogel body and the hydrogel microneedles to form a single homogenous hydrogel structure. The hydrogel microneedles and the body may include methacryalted hyaluronic acid (MeHA).

[0016] In various aspects of the disclosure, the hydrogel microneedles may be formed by pouring a liquid hydrogel solution into a mold. The method may further include placing the hydrogel body encapsulating said sensor element on top of the hydrogel solution before the liquid hydrogel solution is fully cured.

[0017] In various aspects of the disclosure, printing of the sensor element may include printing a first ink comprising a first conducting material and thermoplastic polyurethane (TPU) onto a substrate to form a working electrode, printing a second ink comprising a second conducting material and thermoplastic polyurethane (TPU) onto said substrate to form a counter electrode, and printing a third ink comprising a third conducting material and thermoplastic urethane (TPU) onto said substrate to form a reference electrode. At least one of the working electrode, the counter electrode, and the reference electrode may be configured to recoverably deform or elongate without breaking. At least one of the working electrode, the counter electrode, and the reference electrode may include a portion that is serpentine-shaped to facilitate recoverable deformation or elongation. The substrate may include polyvinyl alcohol (PVA). The hydrogel body may be formed by pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a polydimethylsiloxane (PDMS) mold. The hydrogel microneedles may be formed by pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a mold, extracting the hydrogel microneedles from the mold, and curing the hydrogel microneedles with ultraviolet radiation. BRIEF DESCRIPTION OF THE DRAWINGS

[0018] The above-mentioned and other features and advantages of this disclosure, and the manner of attaining them, will become more apparent and will be better understood by reference to the following description of embodiments of the disclosure taken in conjunction with the accompanying drawings, wherein:

[0019] FIG.1 is a schematic of a sensor patch applied to the skin and an exploded view revealing the unassembled components of the sensor patch in accordance with one embodiment of the present disclosure.

[0020] FIG.2 is a schematic of a sensor patch during interstitial fluid (ISF) extraction and corresponding images of hydrogel microneedles swelling during such interstitial fluid extraction.

[0021] FIG.3 is an image illustrating a fully constructed sensor patch without microneedles.

[0022] FIG.4 is an image of a sensor patch highlighting hydrogel microneedles at one working electrode.

[0023] FIG.5 is an image of a sensor patch highlighting hydrogel microneedles at a conductive trace.

[0024] FIG.6A is a flowchart illustrating a method for manufacturing a sensor patch.

[0025] FIG.6B is a schematic illustrating a method for manufacturing a sensor patch.

[0026] FIG.7 is a graph showing swelling ratio and estimated maximum principal strain of the sensor patch over time in a swelling test.

[0027] FIG.8 is a graph showing the relative resistance change of a sensor patch in response to strain.

[0028] FIG.9 illustrates relative resistance changes of a sensor patch following 1,000 cycles of stretching at 10%, 15%, and 20%.

[0029] FIG.10A illustrates FEA results of a sensor element under uniaxial or biaxial stretching.

[0030] FIG.10B illustrates optical images and FEA results of sensor electrodes under uniaxial stretching.

[0031] FIG.10C illustrates optical images and FEA results of sensor electrodes under biaxial stretching.

[0032] FIG.11A is a schematic illustration of hydrogel microneedles in accordance with one embodiment of the present disclosure.

[0033] FIG.11B is a photograph of a 15 x 15 array of bare hydrogel microneedles.

[0034] FIG.11C illustrates sequential images of hydrogel microneedles as they uptake fluid over time.

[0035] FIG.11D is a graph depicting fluid absorption over time for bare microneedles and microneedles integrated within a sensor patch.

[0036] FIG.12A illustrates FEA results of a sensor patch and the internal sensor arrays under swelling.

[0037] FIG.12B is a graph of stress-strain curves for a sensor patch and a sensor patch without hydrogel.

[0038] FIG.13A is an optical image showcasing porcine skin before and after the insertion of hydrogel microneedles.

[0039] FIG.13B is an image of porcine skin after insertion of hydrogel microneedles and incubation of methylene blue dye.

[0040] FIG.14 is a modeling schematic of the sensor patch with a circuit analogy, and the corresponding sensing delay.

[0041] FIG.15 illustrates simulation results of the signal delay in response to the MN aspect ratio (r = ln / s), the MN length (ln), the diffusivity (D = DB = Dez = Dn) and the thickness of the bulk hydrogel patch (xB).

[0042] FIG.16 illustrates the amperometric response of a glucose sensor patch without microneedles in bath solution.

[0043] FIG.17 illustrates a graph of a calibration curve of a glucose sensor patch without microneedles in bath solution.

[0044] FIG.18A is a graph of an amperometric response of a lactate sensor patch without microneedles in bath solution.

[0045] FIG.18B is graph of a calibration curve for a lactate sensor patch without microneedles in bath solution.

[0046] FIG.19A illustrates the amperometric response of a lactate sensor patch without microneedles in bath solution.

[0047] FIG.19B is a graph of a calibration curve of a lactate sensor patch without microneedles, including an extra BSA layer on top of the enzyme layer, in bath solution.

[0048] FIG.20A is a schematic of a glucose sensor patch in the Franz cell test.

[0049] FIG.20B shows amperometric response of a glucose sensor patch in Franz cell test.

[0050] FIG.20C is a calibration curve of a glucose sensor patch in the Franz cell test.

[0051] FIG.21 is a graph of the long-term amperometric response of a glucose sensor patch without microneedles under fixed glucose levels.

[0052] FIG.22A shows an amperometric response and quantification of a glucose sensor patch without microneedles against interference species in bath solution.

[0053] FIG.22B shows a calibration curve of a glucose or lactate sensor patch without microneedles in bath solution using a MM-inspired equation.

[0054] FIG.22C shows a calibration curve of a lactate sensor patch without microneedles in bath solution using a MM-inspired equation.

[0055] FIG.23A is an image of a sensor patch on mouse skin, and an enlarged image of one working electrode.

[0056] FIG.23B is an image of mouse skin and a sensor patch five minutes after removal of microneedles.

[0057] FIG.23C is a graph of the amperometric response of a glucose sensor patch and blood glucose level in a glucose sensing test.

[0058] FIG.23D is a graph of the a calibration curve for the glucose sensing test shown in FIG.23C.

[0059] FIG.23E is a graph of the amperometric response of the lactate sensor patch and blood lactate level in the lactate sensing test.

[0060] FIG.23F is a graph of a schematic of a model for the lactate sensing test, and corresponding calibration curve.

[0061] FIG.24 is a graph of background drift fitting of the amperometric response of a lactate sensor patch in a lactate sensing test.

[0062] Corresponding reference characters indicate corresponding parts throughout the figures. Although the drawings represent embodiments of the present disclosure, the drawings are not necessarily to scale, and certain features may be exaggerated or omitted in some of the drawings in order to better illustrate and explain the present disclosure.DETAILED DESCRIPTION I. Definitions

[0063] For purposes of the following detailed description, it is to be understood that the disclosure may assume various alternative variations and step sequences, except where expressly specified to the contrary. Moreover, other than in any operating examples, or where otherwise indicated, all numbers expressing, for example, quantities of ingredients used in the specification and claims are to be understood as being modified in all instances by the term “about”. Accordingly, unless indicated to the contrary, the numerical parameters set forth in the following specification and attached claims are approximations that may vary depending upon the desired properties to be obtained by the present disclosure. At the very least, and not as an attempt to limit the application of the doctrine of equivalents to the scope of the claims, each numerical parameter should at least be construed in light of the number of reported significant digits and by applying ordinary rounding techniques.

[0064] Notwithstanding that the numerical ranges and parameters setting forth the broad scope of the disclosure are approximations, the numerical values set forth in the specific examples are reported as precisely as possible. Any numerical value, however, inherently contains certain errors necessarily resulting from the standard variation found in their respective testing measurements.

[0065] Also, it should be understood that any numerical range recited herein is intended to include all sub-ranges subsumed therein. For example, a range of “1 to 10” is intended to include all sub-ranges between (and including) the recited minimum value of 1 and the recited maximum value of 10, that is, having a minimum value equal to or greater than 1 and a maximum value of equal to or less than 10.

[0066] In this application, the use of the singular includes the plural and plural encompasses singular, unless specifically stated otherwise. In addition, in this application, the use of “or” means “and / or” unless specifically stated otherwise, even though “and / or” may be explicitly used in certain instances. Further, in this application, the use of “a” or “an” means “at least one” unless specifically stated otherwise. II. Wearable Sensor Patches

[0067] The present disclosure provides a wearable sensor patch that incorporates hydrogel microneedles for rapid, minimally invasive monitoring of biomarkers in interstitial fluid. The sensor patch is stretchable, ensuring consistent signal stability by adjusting to skin deformation during the ISF extraction process.

[0068] In general, the wearable sensor patch includes (i) a base layer; (ii) a hydrogel body; and (iii) a sensor element.

[0069] The base layer of the wearable sensor patch is intended to contact a user’s skin when the sensor patch is applied. It also functions as the bottom structural layer of the patch, providing support for the sensor element and hydrogel body above. Accordingly, it is preferred that the base layer is sufficiently pliable for interfacing with both the skin and the sensor element. In some embodiments, the base layer includes a hydrogel. Suitable hydrogels include those derived from polyvinyl alcohol (PVA), polyacrylamide, polyethyelene glycol (PEG), hyaluronic acid, methacrylated hyaluronic acid (MeHA) chitosan, gelatin, agarose, osmolytes, alginate, and combinations of the foregoing.

[0070] In some embodiments, the base layer includes hydrogel microneedles which contact the skin and collect interstitial fluid (ISF) for analysis by the wearable sensor patch. These microneedles penetrate the skin and extract ISF that flows into the body of the wearable sensor patch and onto the sensor element. In preferred embodiments, the hydrogel microneedles may include methacrylate hyaluronic acid with a molecular weight of 250 kDa. FIG.1 shows a schematic representation of a wearable sensor patch applied to skin 2 (left) and an exploded view revealing the individual components of a sensor patch 100 (right) including a plurality of hydrogel microneedles 102 at the bottom.

[0071] As shown, when applied to skin 2, hydrogel microneedles 102 may extend through an epidermis layer 3 of skin 2 and into dermis layer 4 of skin 2 for uptake of ISF as described further herein. In some embodiments, the hydrogel microneedles of the base layer may have a fluid absorption rate of 7.4 ± 0.9 μL min-1(R2= 0.970).

[0072] FIG.2 shows a schematic of the sensor patch during ISF extraction. As shown therein, the hydrogel body 104 and sensor element 108 are capable of swelling in response to uptake of ISF. Optical images of microneedles before ISF uptake 128 are shown in comparison to the swollen microneedles 130 after ISF uptake.

[0073] Referring to FIGs.1 and 2, a hydrogel body 104 of the wearable sensor patch is cured to a base layer 106 below, including hydrogel microneedles 102, and contains a sensor element 108. Hydrogel body 104 may include any suitable hydrogel material such as those derived from polyvinyl alcohol (PVA), polyacrylamide, polyethyelene glycol (PEG), hyaluronic acid, methacrylated hyaluronic acid (MeHA) chitosan, gelatin, agarose, osmolytes, alginate, and combinations of the foregoing. In some embodiments, the curing of hydrogel body 104 to base layer 106 including hydrogel microneedles 102 causes hydrogel body 104 and hydrogel microneedles 102 to form one homogenous hydrogel structure. This uniform encapsulation of hydrogel improves structural integrity of sensor patch 100, allowing it to deform without delamination of its components.

[0074] Sensor element 108 is embedded within hydrogel body 104 and typically includes a conductive layer 110. In some embodiments, the conductive layer is embedded within an elastomeric body 112, such as one made from polydimethylsiloxane (PMDS). The sensor element is configured to detect and analyze certain biomarkers within ISF such as glucose, lactate, and alcohols. Details of sensor element 108 are provided in FIGs.1 and 3. As shown therein, sensor element 108 includes conductive layer 110 which is sandwiched within a polydimethylsiloxane (PMDS) body 112. Sensor element 108 including elastomeric body 112 and conductive layer 110 is capable of flexing and elongating in response to movement translated through wearable sensor patch 100 from a user’s skin 2 and from swelling due to uptake of ISF.

[0075] Conductive layer 110 of sensor element 108 is responsible for detecting and quantifying metabolites, electrolytes, and other clinically relevant targets in the ISF. FIG.3 provides a schematic of conductive layer 110 of sensor element 108 in accordance with one embodiment of the present disclosure. As shown therein the conductive layer includes:(i) a working electrode 114; (ii) a counter electrode 116; and (iii) a reference electrode 118.

[0076] Advantageously, each of the three electrodes may include an elastomeric material, allowing the electrodes to flex and adjust to skin movements or deformations during the ISF extraction and measuring process. The elastomeric material may be a thermoplastic elastomer such as styrene block copolymers or polyolefin based elastomers; a polybutadiene; a polyurethane elastomer such as thermoplastic polyurethane.

[0077] Each of the three electrodes may also include a conductive material such as copper, aluminum, silver, gold, brass, nickel, zinc, or iron. Salts of the foregoing metals, such as silver chloride, may also be used. The conductive metal may be finely divided or fragmented into a powder, pellets, shavings, or flakes.

[0078] As shown in FIG.3, the working electrode 114, counter electrode 116, and reference electrode 118 may be partially or wholly defined in a curving, serpentine shape. This shape, combined with the flexibility provided by the elastomeric material, allows the electrodes to recoverably deform or elongate without breaking. In some embodiments, the working and reference electrodes may adopt a circular shape to maximize surface area. FIG.4 is an image of a sensor patch showing a circular working electrode 114 and microneedles 150 and 152. FIG.5 is an image of a sensor patch showing a microneedle 154 on a curving reference electrode 118.

[0079] Together with the hydrogel body and the base layer, the entire sensor element may provide consistent signal stability by adjusting to deformations of a user’s skin, or deformations caused by uptake of ISF and subsequent swelling of the patch. The wearable sensor patch is configured to recoverably deform or elongate in response to user movement without delaminating or breaking any one of the base layer, the hydrogel body, the sensor element and a combination thereof. III. Properties of Wearable Sensor Patches

[0080] The wearable sensor patches of the present disclosure may have a number of properties which make them desirable for a variety of medical and wearable sensing applications.

[0081] In some embodiments, the wearable sensor patch may exhibit an in vitro glucose sensitivity of 0.020 ± 0.002 μA mM-1, 0.021 ± 0.002 μA mM-1, 0.022 ± 0.002 μA mM-1, 0.023 ± 0.002 μA mM-1, 0.024 ± 0.002 μA mM-1, 0.025 ± 0.002 μA mM-1, 0.026 ± 0.002 μA mM-1, 0.027 ± 0.002 μA mM-1, 0.028 ± 0.002 μA mM-1, 0.029 ± 0.002 μA mM-1, or 0.030 ± 0.002 μA mM-1, with a linear range of 0.1 to 3 mM.

[0082] In some embodiments, the wearable sensor patch may exhibit an in vitro lactate sensitivity of 0.0025 ± 0.0004 μA mM-1, 0.0026 ± 0.0004 μA mM-1, 0.0027 ± 0.0004 μA mM-1, 0.0028 ± 0.0004 μA mM-1, 0.0029 ± 0.0004 μA mM-1, 0.0030 ± 0.0004 μA mM-1, 0.0031 ± 0.0004 μA mM-1, 0.0032 ± 0.0004 μA mM-1, 0.0033 ± 0.0004 μA mM-1, 0.0034 ± 0.0004 μA mM-1, or 0.0035 ± 0.0004 μA mM-1, with a linear ranges of 0.1 to 12 mM.

[0083] In some embodiments, the wearable sensor patch may exhibit an in vivo glucose sensitivity of 0.015 ± 0.001 μA mM-1, 0.016 ± 0.001 μA mM-1, 0.017 ± 0.001 μA mM-1, 0.018 ± 0.001 μA mM-1, 0.019 ± 0.001 μA mM-1, 0.020 ± 0.001 μA mM-1, 0.021 ± 0.001 μA mM-1, 0.022 ± 0.001 μA mM-1, 0.023 ± 0.001 μA mM-1, 0.024 ± 0.001 μA mM-1, or 0.020 ± 0.001 μA mM-1with a detection range of 1 to 8 mM. IV. Methods of Making Wearable Sensor Patches

[0084] The present disclosure also provides methods for manufacturing wearable sensor patches, such as those described in sections II and III above.

[0085] Referring to FIG.6A, in general, method 500 includes the following steps: (i) printing a sensor element onto a substrate 502; (ii) encapsulating said sensor element in a hydrogel body 504; and (iii) curing the hydrogel body with a plurality of hydrogel microneedles 506.

[0086] The first step involves constructing a sensor element on top of a substrate by printing a series of inks. In some embodiments, the substrate is a hydrogel material such as one derived from polyvinyl alcohol (PVA), polyacrylamide, polyethyelene glycol(PEG), hyaluronic acid, methacrylated hyaluronic acid (MeHA) chitosan, gelatin, agarose, osmolytes, alginate, and combinations of the foregoing.

[0087] In some embodiments, the printing step includes: (i) printing a first ink including a first plurality of silver flakes and thermoplastic polyurethane (TPU) onto a substrate to form a working electrode; (ii) printing a second ink including a second plurality of silver flakes and thermoplastic polyurethane (TPU) onto said substrate to form a counter electrode; and (iii) printing a third ink including at least one of silver and silver chloride and thermoplastic polyurethane (TPU) onto said substrate to form a reference electrode.

[0088] The first, second, and third inks, once cured, may form separate electrodes in the sensor element. In addition to thermoplastic polyurethane (TPU) and a conductive material, each ink may further include a solvent such as tetrahydrofuran (THF) or dimethylformamide (DMF). The ingredients in each ink composition may be combined using a variety of techniques such as drum mixers, planetary centrifugal mixers, cyclone mixers. The first, second, and third inks may be different from each other and / or prepared as separate batches. In some embodiments, each ink may be the same and / or prepared in a single batch. In some embodiments, the first, second, and third inks correspond to a working electrode, a counter electrode, and a reference electrode. Additionally, each of first, second, and third inks may include conductive materials other than silver or silver chloride, including, but not limited to, copper, aluminum, gold, brass, nickel, zinc, iron, or salts thereof, for example.

[0089] In some embodiments, an additional fourth ink including polydimethylsiloxane (PMDS) may be printed on top of and around the electrodes to insulate them from each other and form a unified sensor element. The fourth ink may also include a curing agent.

[0090] The inks may be printed using an automated nozzle injection system. Preferably, while printing the resolution is at least 100 µm in line width, repeatability is about ± 3 µm, and the nozzle speed is about 2.5 mm s-1. After printing, the layers of ink may optionally be cured at a temperature of 50°C, 60°C, 70°C, 80°C, 90°C, 100°C, 110°C, or 120°C. The curing may last for 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, or 6 hours.

[0091] In some embodiments, the electrodes, once printed and cured, may optionally be electroplated with a conductive material such as gold. The electroplating may be done using any known method.

[0092] In some embodiments, an additional transducer layer may be deposited onto the gold-plated electrodes. The transducer layer helps facilitate the conversion of chemical changes into an electrical signal. The transducer layer may be deposited onto the gold-plated electrodes using cyclic voltammetry and a voltage of from –0.5 V to 0.9 V. Suitable materials for the transducer layer include conductive polymers, semiconducting materials, and dyes like Prussian blue, for example. Following the optional electroplating and deposition of the transducer layer, the sensor electrodes may be washed with deionized water and dried with compressed air.

[0093] In some embodiments, an additional enzyme layer may be deposited onto the electrodes such that it is in direct contact with the working electrode, counter electrode, and / or reference electrode. The enzymes in this layer may react and form a complex with biomarkers in the ISF. These reactions trigger an amperometric response which is sensed by the electrodes. The enzyme layer may be immobilized onto the sensor electrodes by drop casting a solution of oxidase enzyme and a crosslinking agent onto the electrodes. Suitable oxidase enzymes include lacate oxidase, glucose oxidase, xanthine oxidase, amine oxidase, alcohol oxidase, amino acid oxidase, sarcosine oxidase, and the like. A preferred crosslinking agent is glutaraldehyde. In some embodiments, the solution also includes a protein such as bovine serum albumin, for example. Following drop-casting the solutions many be left to cross link for at least 6 hours, at least 12 hours, at least 24 hours, at least 48 hours, or at least 72 hours.

[0094] After the printing step, the formed sensor element may be encapsulated in a hydrogel body in step (ii). In general, the hydrogel body may be formed by pouring a solution including water, hydrogel, and a photoinitiator into a mold. The solution may contain any suitable hydrogel material such as one derived from polyvinyl alcohol (PVA), polyacrylamide, polyethyelene glycol (PEG), hyaluronic acid, methacrylated hyaluronic acid (MeHA) chitosan, gelatin, agarose, osmolytes, alginate, and combinations of the foregoing. The photoinitiator may be 2-hydroxy-4’-(2- hydroxyethoxy)-2-methylpropiophenone. The solution may be centrifuged tohomogenously combine the ingredients. The concentration of the hydrogel material in the solution may be 20 mg / mL, 30 mg / mL, 40 mg / mL, 50 mg / mL, 60 mg / mL, 70 mg / mL, or 80 mg / mL. The concentration of the photoinitiator in the solution may be 0.1 mg / mL, 0.2 mg / mL, 0.3 mg / mL, 0.4 mg / mL, 0.5 mg / mL, 0.6 mg / mL, 0.7 mg / mL, 0.8 mg / mL, or 0.9 mg / mL.

[0095] In some embodiments, the mold includes polydimethylsiloxane (PDMS). In general, molds with a thin and flat shape are preferred to provide the correct geometry for the wearable sensor patch. Once the solution containing water, hydrogel, and the photoinitiator is poured into the mold, it may be left to dry until the hydrogel body partially cures or completely solidifies. Upon partial or complete curing, a sensor element may be transferred onto or into the hydrogel body and additional hydrogel solution may be added to cover the sensor. In other words, upon precuring (partial or full curing) of the hydrogel body, the sensor may be positioned on the hydrogel body, and a second layer of uncured hydrogel is poured on top of the sensor to encapsulate the sensor within hydrogel material so that, upon completion of curing, the sensor is encapsulated within a single hydrogel body. For example, in some embodiments, the sensor element may be added to the mold upon or within a first partially cured hydrogel layer such that, upon full curing, the sensor element is completely encapsulated within a single homogenous hydrogel body. In other embodiments, the sensor element may be placed upon a fully cured or partially cured hydrogel body, then covered with fresh liquid hydrogel. Upon curing, the sensor element becomes fully encapsulated within a single homogenous hydrogel body.

[0096] FIG.6B provides a schematic of a typical fabrication process. In step 3010, a hydrogel solution 3004 is poured into mold 3002 which contains voids for forming microneedles 3012. Once the hydrogel solution is partially or fully cured, the sensor element 3000 is added atop the hydrogel solution 3004. In step 3020, additional hydrogel solution is added and fully cured atop the sensor element 3000, forming a single homogenous hydrogel body 3006. Step 3030 shows the wearable sensor patch removed from the mold 3002 with the sensor element fully encapsulated in the hydrogel body.

[0097] In the last step (iii), the hydrogel body from step (ii) is cured to a plurality of hydrogel microneedles. In general, the hydrogel microneedles are formed by pouring a solution containing water, hydrogel, and a photoinitiator into a mold. The solution may contain any suitable hydrogel material such as one derived from polyvinyl alcohol (PVA), polyacrylamide, polyethyelene glycol (PEG), hyaluronic acid, methacrylated hyaluronic acid (MeHA) chitosan, gelatin, agarose, osmolytes, alginate, and combinations of the foregoing. The photoinitiator may be 2-hydroxy-4’-(2- hydroxyethoxy)-2-methylpropiophenone. The solution may be centrifuged to homogenously combine the ingredients. The concentration of the hydrogel material in the solution may be 20 mg / mL, 30 mg / mL, 40 mg / mL, 50 mg / mL, 60 mg / mL, 70 mg / mL, or 80 mg / mL. The concentration of the photoinitiator in the solution may be 0.1 mg / mL, 0.2 mg / mL, 0.3 mg / mL, 0.4 mg / mL, 0.5 mg / mL, 0.6 mg / mL, 0.7 mg / mL, 0.8 mg / mL, or 0.9 mg / mL.

[0098] The mold used for forming the hydrogel microneedles may include polydimethylsiloxane (PDMS). Any microneedle mold, such as those commercially available from Blueacre Technologies may be used. The shape of the microneedles is not particularly critical and will be determined by the type of mold used. In some embodiments the microneedles may be in the shape of a cone or a pyramid. In a preferred embodiment, the microneedles are pyramid shaped and each have a length of 600 µm, a base width of 300 µm, and a tip-to-tip distance of 500 µm. In some embodiments, the microneedles may have a length greater than 300 µm, a base width of less than 500 µm, and a tip-to-tip distance of greater than 50 µm. FIGs.4 and 5 provide optical microscope images of the sensor patch highlighting the hydrogel MNs on the surface. Following addition of the hydrogel solution to the microneedle mold, the molds may undergo centrifugation to eliminate bubbles and facilitate the filling of the solution into the tips of the microneedle cavities. The molds may then be left at room temperature until the hydrogel completely dries. After drying, the microneedles may be extracted from the molds and exposed to UV light of 365 nm for 5-20 minutes to induce crosslinking.

[0099] Once the sensor patch is assembled with a sensor element encapsulated within a continuous hydrogel body, the entire patch may then be crosslinked using 365UV radiation for 5-20 minutes to fully cure the hydrogel body to the plurality of hydrogel microneedles. After the patch is cured, it is ready to be applied to a user’s skin. EXAMPLES Example 1: Fabrication of the Sensor Patch

[0100] A water-soluble PVA solution (10 wt.t% of Mowiol 4-88 in deionized water) was spin cast onto a glass slide (Dow Corning) at 2,000 rpm for 30 seconds and was then annealed at 80°C for 2 hours. Concurrently, the PDMS ink for printing was prepared by mixing base solutions (Sylgard 184 & Dowsil SE 1700, Dow Corning, Inc.) and a curing agent in a weight ratio of 5:5:1. The Ag-TPU ink was created by first dissolving 1.48g TPU (Elastollan® C60AW) in a combination of 3.76 g tetrahydrofuran (THF, Sigma Aldrich, Inc.) and 4 g dimethylformamide (DMF, Sigma Aldrich, Inc.), and then blending it with 5.96 g Ag flakes (average particle size 2-5 μm, Inframat Advanced Materials, Inc.) using a planetary centrifugal mixer (Thinky, ARE-310). The freshly prepared inks and the Ag / AgCl ink (Cl-4025, Nagase America, Inc.) were directly printed layer-by-layer on the glass slide with a water-soluble PVA layer using an automated nozzle injection system (Nordson EFD) situated on a three-axis computer-controlled translation stage. The resolution was at least 100 µm in line width, repeatability was ± 3 µm, and the nozzle speed was 2.5 mm s-1. The printed layers were then cured at 80°C for 2 hours. Cu wires were connected to the extended pads of the electrodes using Ag paste and baked at 80°C for 1 hour. These junctions were subsequently insulated with PDMS. In the next step, two working electrodes were electroplated with Au (24K Pure Gold Bath Solution, Gold Plating Services, Inc.) using the automotive plating kit (Universal Plater – Chrome Edition, Gold Plating Services, Inc.). A Prussian blue transducer layer was deposited onto the Au-plated Ag-TPU working electrodes using cyclic voltammetry. The electrolyte solution contained 2.5 mM iron trichloride (FeCl3), 100 mM potassium chloride (KCl), 2.5 mM potassium ferricyanide (K3Fe(CN)6), and 100 mM hydrochloric acid (HCl). Cyclic voltammetry was conducted from –0.5 V to 0.9 V versus Ag / AgCl (3 M NaCl) and reversed 10 times at a rate of 50 mV s-1. Following deposition, the sensor electrodes were washed in deionized (DI) water and dried withcompressed air. Enzymes were immobilized onto the sensor electrodes by drop casting a solution of oxidase enzyme (either GOx or LOx), BSA, and the crosslinking agent, glutaraldehyde. The solution for lactate sensors included 0.8 U µL-1LOx, 0.8% BSA, and 0.1% glutaraldehyde. For glucose sensors, the solution contained 0.2 U µL-1GOx, 0.8% BSA, and 0.1% glutaraldehyde. Using a 10 µL micropipette, 0.5 µL of solution was placed onto an electrode and then immediately aspirated back off the electrode. This process was repeated three times for each electrode at one-minute intervals. Following drop-casting, the solutions were left to crosslink at room temperature for at least 1 day. Lastly, for lactate sensors, an additional 0.5 µL of the glutaraldehyde / BSA solution without LOx was drop-casted. The prepared sensor patch was then integrated with hydrogel MNs, as detailed in the next section. Example 2: Integration of Microneedles into the Sensor Patch

[0101] In this Example, microneedles were incorporated into the sensor patch from Example 1.

[0102] MeHA was synthesized by mixing 4 g of hyaluronic acid (250 kDa, Creative PEGWorks, Inc.) with 200 mL of DI water, and the solution was stirred at 4 °C overnight. This was followed by the dropwise addition of 133 mL DMF and 4 mL methacrylic anhydride (Sigma Aldrich, Inc.), with pH regulated using 1 M NaOH, while maintaining stirring at 4 °C overnight. Subsequently, NaCl was added to achieve a final concentration of 0.5 M. MeHA was precipitated out from the aqueous solution by adding 300 mL of ethanol. The precipitate was thrice washed in ethanol and then subjected to centrifugation for separation. Ultrapure water was utilized to dissolve the precipitate, and this solution was placed in dialysis membranes with a molecular weight cutoff of 12,000 kDa (Sigma Aldrich, Inc.) and dialyzed against ultrapure water for seven days, with water changes occurring twice daily. The product was then lyophilized for three days to remove all residual water, resulting in purified MeHA. This dried MeHA was stored at -20°C until needed. For the creation of hydrogel microneedles, the dried MeHA was dissolved in DI water at a concentration of 50 mg mL-1. The photoinitiator (2- Hydroxy-4’-(2-hydroxyethoxy)-2-methylpropiophenone) was dissolved in ethanol at a concentration of 100 mg mL-1. This photoinitiator solution was added to the hydrogelsolution to achieve a final photoinitiator concentration of 0.5 mg mL-1within the hydrogel solution. This solution was subsequently poured into PDMS microneedle molds (Blueacre Technologies, Inc.). The microneedle molds incorporated a 15 × 15 array of square pyramidal microneedles, each with a length of 600 µm, a base width of 300 µm, and a tip-to-tip distance of 500 µm. Following the addition of the hydrogel solution, the microneedle molds underwent centrifugation at 2,500 rpm for 10 minutes to eliminate bubbles and to facilitate the filling of the hydrogel solution into the tips of the microneedle cavities. The molds were then left at room temperature until the hydrogel was dry, i.e., partially cured. Upon drying, the microneedles were extracted from the molds and exposed to UV light of 365 nm for 10 minutes to induce crosslinking before application. For the assembly of the sensor patch, the sensor, initially fabricated on a PVA-coated glass slide, was submerged in deionized (DI) water overnight to facilitate its release from the glass slide. Simultaneously, to prepare the hydrogel substrate, a hydrogel solution consisting of 50 mg mL-1MeHA and 0.5 mg mL-1photoinitiator in DI water was created. This solution was then poured into a PDMS mold to form a thin, flat substrate and was left to dry thoroughly. Upon drying, the substrate was detached from the mold and the sensor was transferred onto it. The MN mold was subsequently prepared, into which the hydrogel solution was added and centrifuged. The sensor patch, placed on the flat hydrogel substrate, was then introduced into the MN mold and left to dry completely. Following drying, the integrated sensor patch, now combined with the hydrogel MNs, was extracted from the mold. Any excess hydrogel was carefully trimmed off, and the patch was crosslinked using 365 nm UV radiation for 10 minutes before application. Example 3: Mechanical Testing

[0103] Standard tensile tests were conducted to determine the elastic moduli of the sensor patch without MNs, hydrogel MNs, and the sensor patch including MNs). The specimens were loaded on the chuck of a tensile testing machine (ESM303, Mark-10) and then stretched at the elongation rate of 5% per minute. For the measurement of the relative resistance change (ΔR / R), one working electrode was selected as the representative, and the electrode and its extended pad were connected to a sourcemeter (Keithley 2400; Keysight, Inc.) via a conducting Cu wire during cyclic stretching up to 1,000 times at the applied strain of 10%, 15%, and 20% with the elongation rate of 100% per minute, respectively. ΔR is the difference between current resistance and initial resistance; and R is the initial resistance value.

[0104] Referring again to FIG.2, a schematic of the stretched sensor patch during ISF extraction is depicted, illustrating the swelling of both the hydrogel body and MNs upon ISF extraction. Referring to FIG.7, the experimental results of the swelling ratio is provided along line 600, wherein the results include the average swelling ratio of three sensor patches inserted into an agarose gel covered with Parafilm M, simulating the skin. The dimensions of the hydrogel MNs were measured every 5 minutes, revealing a saturation point at approximately 9% swelling. The swelling ratio was defined as ΔP / P, where ΔP represents the difference between the current and initial perimeters, and P is the initial perimeter value. Still referring to FIG.7, the estimated maximum principal strain is provided along line 602. As illustrated, less than 15% of maximum principal strain is estimated to be experienced by the conductive layer of the sensor patch without MN over time.

[0105] As shown in FIG.8, the sensor patch without MNs exhibited an insignificant relative resistance change (ΔR / R < 20%) when subjected to strains below 30%, where the demonstrated resistance change is shown along line 700 and the 30% strain threshold is shown along line 702. However, at higher strains, i.e., beyond threshold 702, the ΔR / R increased rapidly as the serpentine traces straighten. To further assess the stability of the sensor patch without MNs, cyclic tests were conducted for 1,000 cycles under 10%, 15%, and 20% strain, revealing that the ΔR / R remained below 10%, 20%, and 50% as shown in FIG.9, taking into account the sensor arrays saturated swelling ratio of approximately 10%. The FEA results shown in FIG.10A validate that the conductive layer experienced a maximum principal strain of 74% under uniaxial stretching up to 100% and 38% under biaxial stretching up to 40%, respectively. Additionally, under 30% uniaxial or biaxial stretching, the conductive layer would experience a maximum principal strain of 28% or 30%. This is greater than the maximum principal strain obtained from the swelling test, demonstrating the stable generation of electrical signals of the sensor patch during measurements. In thebiaxially stretched state, relatively greater strain concentration was observed compared to uniaxial stretching, leading to fractures occurring at 40% strain, consistent with the experimental results. The corresponding optical images and FEA results of the sensor patch without MNs under uniaxial and biaxial stretching are shown in FIG.10B and 10C, respectively. Example 4: Fluid Uptake and Swelling Tests

[0106] Fluid uptake was tested on the hydrogel MNs and the sensor patches. A skin model was made with 1.4% agarose (Fisher Scientific, Inc.) combined with 10 mM Rhodamine B (Sigma Aldrich, Inc.) and covered with a layer of Parafilm M. The MNs were pierced through the Parafilm M layer with a spring-loaded application tool (Micropoint, Inc.) and swelled for 2, 5, 10, or 15 minutes. Samples were fixed by a top holder (75 g) to prevent detachment. Samples were weighed before and after swelling to evaluate fluid uptake, and microscope images were taken to evaluate swelling and visualize Rhodamine B uptake into the MNs. Swelling ratio tests were performed on the sensor patches similarly, and the perimeter (P) of the hydrogel MN array was measured every 5 minutes until saturation. The swelling ratio was defined as ΔP / P, where ΔP is the difference between current perimeter and initial perimeter; and P is the initial perimeter value.

[0107] FIG.11A presents a schematic of the hydrogel MNs made from MeHA. Specifically, hyaluronic acid with a molecular weight of 250 kDa was selected for synthesizing MeHA, in order to produce a gel that was sufficiently pliable for effective interfacing with both the skin and the sensor, while still maintaining a high enough molecular weight to prevent accelerated degradation rates. FIG.11B depicts a 15 × 15 array of these MNs.

[0108] Referring to FIG.11C, the results of fluid absorption by the MNs is illustrated. To simulate the skin, an agarose gel infused with 10 mM Rhodamine B was used, overlaid with a layer of Parafilm M. The samples were initially weighed, then inserted into the agarose gel, and later removed after set time intervals for reweighing and imaging. The presence of the red hue from Rhodamine B, indicative of fluid flow, showcased the extraction and movement of fluid from the MN tips to the sensor patchwithin 10 minutes. The volume of the absorbed fluid was calculated based on the difference in weight before and after insertion, and the results of such is shown in FIG. 11D. As shown along line 1000, a fluid absorption rate of 7.4 ± 0.9 μL min-1(R2= 0.970) for bare MNs was observed, and a fluid absorption rate of 5.3 ± 0.2 μL min-1(R2= 0.998) for the MNs integrated with the sensor patch as shown along line 1002. The presence of the sensor patch caused a minor reduction in the fluid absorption rate of the MNs. Example 5: Mechanical Simulation

[0109] A commercial software package, Abaqus, was used for FEA to investigate the strain levels impacting the conductive layers of the sensor patch under conditions of stretching and swelling deformation. Three layers of the sensor patch on a PDMS film were modeled to enable analysis of uniaxial and biaxial stretching. The tie-constraint option was utilized to bond each interface between the layers and the PDMS film. The mechanical properties of PDMS were defined using the Neo-Hookean hyperelastic model, based on uniaxial tension test results and a Poisson's ratio of 0.495. Both Ag- TPU and hydrogel were treated as linear elastic materials with a modulus of 0.248 MPa and a Poisson's ratio of 0.4. For stretching, displacement boundary conditions were implemented on the side surfaces of the PDMS film, while out-of-plane displacement was constrained on the surface opposing the sensor patch. For the analysis of swelling- like deformation, hydrogel was selected as the host region using the embedded region constraint option, being considered a linear elastic material with a modulus of 16.7 MPa and a Poisson’s ratio of 0.46. To roughly estimate the strain concentration on the conductive layer resulting from the swelling of the hydrogel encasing the sensor patch, a thermal expansion simulation was performed, driven by conceptual similarity. The top and bottom surfaces of the hydrogel were restrained by rigid surfaces, leaving the square region designated for MN fabrication exposed. The maximum principal strains inflicted on the conductive layer were plotted, corresponding to the point when the swelling ratio of the FEA result matched the value observed experimentally.

[0110] FIG.12A provides the FEA results of the sensor patch and the internal sensor arrays under swelling, while FIG.12B maps the stress-strain curves of thesensor patch along line 1200 and the sensor patch without hydrogel along line 1202. Referring again to FIG.9, the relative resistance changes of the sensor patch following 1,000 cycles of stretching is illustrated at 10%, 15%, and 20%. Example 6: Ex vivo Porcine Skin Studies and Histology

[0111] Porcine skin was procured from the lower belly of a 2-year-old specimen. The skin was sanitized with ethanol and shaved to eliminate any excess hair. The hydrogel MN patches were then positioned on the excised skin and inserted using a spring-loaded application tool. Subsequent to this application, methylene blue (1 mg mL-1, Sigma Aldrich, Inc.) was applied to the skin and allowed to incubate for one minute. Afterwards, the skin was cleansed and imaged to view the insertion of the MNs. Following this, the skin samples were preserved in a formalin solution for three days, which was succeeded by histological processing that included staining with hematoxylin and eosin.

[0112] FIG.13A presents optical images of porcine skin, both before and after the insertion of the MNs. Despite the abrasions on the porcine skin incurred during preparation for the microneedle insertion study, evidence of the MNs penetrating the porcine skin is provided by the methylene blue marks observed in the pores post- insertion, as shown in FIG.13B. This is consistent with the designed length of the MNs exceeding the minimum required for skin penetration and ISF collection. Example 7: Simulation of Signal Delay in Response to Different Parameters

[0113] A unit cell was defined, which included two rectangular domains (the sensor patch and skin dermis) interconnected by a hollow parallelepiped with a square cross-section (representing an individual hydrogel MN). The sensing site was positioned at the upper end of the sensor patch. The simulation was conducted using COMSOL Multiphysics finite element software. Initially, the physical laws within each domain were defined, followed by the application of flux continuity across the interfaces. The diffusion of the analyte was governed by Fick’s second law, with corresponding diffusion coefficient Diin each domain. The enzyme layer was implicated in enzymatic conversion (from analyte G to product P), under the assumption of oxygen availability. Given thefollowing chemical reaction, the balanced Equation 1 in the enzyme layer was defined as: ^^^^ , ^^^^^E + G�f ^^^¾¾�rEG→cE + PEquation 1 where E, EG and P are the free enzyme, intermediate complex, and product molecules; kf, kr, and kc are the rate constants for the forward, reverse, and catalytic steps of the conversion process. Finally, the redox reaction on the electrode surface was described by the Butler-Volmer formalism.

[0114] In order to calculate the sensing delay through the sensor patch, a modeling schematic and a circuit analogy were employed, as shown in FIG.14. While various approaches exist for ISF extraction and analysis, passive methods stand out as they avoid extra power consumption and prevent skin inflammation. The model used emphasized passive methods that rely on diffusive transport for ISF absorption. An inherent time lag is present in these patches between biofluid extraction and sensor responses, stemming from the diffusion required by analyte molecules to reach the sensor patch's upper end. The delay in sensor response accounted for both the intrinsic transport time between blood and dermis and the components dependent on the design of the sensor. To model the sensor patch, it was initially regarded as a stack of multiple domains (MN for the hydrogel MN, B for the bulk hydrogel patch, ez for the enzyme layer), with each domain defined by the geometric and physical attributes of the system (diffusivity Diand thickness xifor each domain i, MN length ln, MN base and tip aperture sbase and stip, MN tip-to-tip distance sd, width of the enzyme layer L). The MNs were depicted with a blunted shape, rather than an idealized conical form, to emphasize the significance of the effective tip radius (stip) in modeling.

[0115] Secondly, as displayed in the circuit analogy of FIG.14, each domain was modeled with a diffusive resistor (Ri in s m-3), in line with the theory of effective media formulation, to approximate the transport of the analyte molecule across the domains. From a phenomenological perspective, the analyte absorption across the MN tip and base is comparable to ion uptake by bacteria, while the transport along the MN lengthand the bulk hydrogel patch resembles the spreading resistance of a point contact. Lastly, the resistance was multiplied by the corresponding volume of transported biofluid (Viin m3) to determine the delay time across each domain (ti= RiVi). The total response time (ttot) equated to the sum of the individual transient contributions necessary for the analyte (G) to diffuse across the individual domains until it reached the sensing site (electrode + enzyme): ^^^^tot = ^^^^MN + ^^^^B + ^^^^ez2where tMNand tBare the delays due to the MN and the bulk hydrogel patch, respectively, and tez is the effective time resulting from diffusion and reaction in the enzyme layer. In other words, any variation of the analyte concentration in the dermis, GR, requires a lag time ttotto be reflected in the upper end of the enzyme layer, wherethe sensing site is located, G�ez(t). It is numerically shown that the analyte responsewithin the sensing site requires an initial turn-on delay (tMN + tB) before an exponential-like behavior with time constant equal to tez reaches the steady state level G�ez,SS.

[0116] The individual contributions to Equation 2 were quantified in terms of geometric and physical modeling of the domain. Specifically: ^^^^MN = ^^^^tip^^^^tip,eff + ^^^^n^^^n^ + ^^^^base^^^^base,eff ≈ ^^^^n^^^n^tez = time resulting from enzyme reaction vs. molecule supply Equation 5

[0117] In Equation 3, Rtip∝ (Dbodystip)-1, Rbase∝ (DBsbase)-1and Rn∝ ln(Dnstip2)-2are the resistances of MN tip and base apertures, and MN length, respectively. Vtip,eff, Vbase,eff and Vn are the volume of transported biofluid corresponding to MN tip and base apertures, and to MN volume (truncated pyramid or cone, for example), respectively. In Equation 5, tezresults from the balance between enzymatic conversion of G molecules against its supply from the dermis. For traditional MN-based patches with enzyme located at the upper end of the patch, the molecule supply is the bottleneck for tez, delaying the overall response time.

[0118] The analytical results of the signal delay in response to different parameters are represented in FIG.15, including the MN aspect ratio (r = ln / s) shown in graph 1400, the MN length (ln) shown in graph 1402, the diffusivity (D = DB= Dez= Dn) shown in graph 1404, and the thickness of the bulk hydrogel patch (xB) shown in graph 1406. The theory, validated against COMSOL Multiphysics finite element software simulations, confirms that depending on the sensor design, the overall delay ranges from several minutes to hundreds of minutes. To justify the experimental characterization, it was assumed that sbase = 150 μm, stip = 35 μm, sd = 500 μm, ln = 600 μm, Dbody= 3 × 10-11m2s-1, Dn= DB= 7.3 × 10-11m2s-1and apply Equation 3. The estimated diffusion time across the MN was equal to tMN,th= 12.7 minutes, which is consistent with the experimentally measured delay of tMN,exp = 10 minutes. Example 8: Benchtop Evaluations in Bath Solution

[0119] Glucose or lactate sensor patches, without MNs, were fabricated on a glass slide and connected to the potentiostat. These were then immersed in a 1× PBS solution, stirred at 200 rpm. Following this, the concentration of the target biomarker was incrementally increased (for glucose: 0, 0.1, 0.5, 1, 2, 3, 5, 7, 10, 20 mM; for lactate: 0, 0.02, 0.05, 0.09, 0.2, 0.5, 0.9, 1.8, 5.3, 8.8, 12, 18 mM) by adding the concentrated solution (either 1 M glucose solution or 100 mM lactate solution). An amperometric method, employing onboard reference and counter electrodes, was used to generate the current signal by applying -0.1 V versus Ag / AgCl. The sensing performance was displayed using smoothed lines and calibration was determined basedon the raw results. The final 50 seconds of the stabilized data were recorded as means ± SD for the current response at each biomarker concentration.

[0120] The sensing performance of the sensor patch was initially evaluated and calibrated in a bath solution. For glucose sensing, the glucose sensor patch without MN was connected to the potentiostat and immersed into a 1× phosphate buffered saline (PBS) solution, stirring at 200 rpm. Subsequently, the glucose concentration was incrementally elevated from 0 up to 20 mM by adding a 1 M glucose solution. An amperometric method, with onboard reference and counter electrodes applying -0.1 V versus Ag / AgCl, generated the current signal. The glucose concentration was adjusted every 100 seconds while the current signal stabilized within 60 seconds, as illustrated in FIG.16. The calibration curve, demonstrating the current in response to the glucose concentration, was illustrated in FIG.17 and revealed a 0.024 ± 0.002 μA mM-1(R2= 0.966) sensitivity within the glucose concentration range of 0.1-3 mM.

[0121] The lactate sensing with the lactate sensor patch without MN was performed accordingly, with the only difference being the replacement of 0.2 U µL-1GOx with 0.8 U µL-1LOx. The amperometric result illustrated in FIG.18A and calibration curve illustrated in FIG.18B showed a high sensitivity of 0.68 ± 0.06 μA mM-1(R2= 0.950) within the lactate concentration range of 0-1.8 mM, but a saturation point at 1.8 mM was observed. This saturation level of 1.8 mM was insufficient for measuring the normal human lactate level range (1-13 mM).

[0122] An additional drop-casting process of the same glutaraldehyde / bovine serum albumin (BSA) solution, devoid of LOx, was utilized to create a diffusive barrier on top of the enzyme layer of the lactate sensor, leading to a boost in sensor linearity. The amperometric results of the revised lactate sensor are displayed in FIG.19A. The lactate concentration was adjusted after the signal had stabilized, and the current signal stabilized in 120 seconds. The calibration curve, depicting the current in response to the lactate concentration, was presented in FIG.19B. The lactate sensor had a sensitivity of 0.0030 ± 0.0004 μA mM-1(R2= 0.942) and operated within the lactate concentration range of 0.1-12 mM. The sensitivity declined and the current signal at low lactate levels (0-0.1 mM, outside of the normal human lactate range) due to the additional diffusive barrier in lactate transport and sensing.Example 9: Benchtop Evaluations in Franz Cell

[0123] In order to further examine the combined sensing performance and the MNs' fluid absorption ability, the glucose sensor patch with MNs was chosen as the representative for the Franz cell test. A schematic of the Franz cell system 1900 is represented in FIG.20A, including a sensor patch 1902, Parafilm M 1904, and glucose solution 1906. FIG.20B presents the amperometry results of the glucose sensor patch in the Franz cell 1900. The current signal stabilizes within 600 seconds, and the glucose concentration is adjusted once the signal stabilizes and presents a flat platform within 100 seconds. FIG.20C exhibits the calibration curve of the current response to the glucose concentration, with a sensitivity of 0.023 ± 0.002 μA mM-1(R2= 0.991) within the glucose concentration range of 0-3 mM. As expected, compared to the bath solution, the glucose sensor patch exhibits a longer stabilization time of up to 10 minutes in the Franz cell 1900 due to the signal delay caused by glucose extraction and transport.

[0124] To mimic skin insertion, the glucose sensor patch was fabricated and penetrated through one layer of Parafilm M with its hydrogel MNs, exposing it to a 1× PBS solution. To prevent potential detachment due to swelling, the glucose sensor patch was secured with a top holder. The glucose concentration in the PBS solution was then altered (0, 1, 2, 3, 5, 10, 20 mM) by injecting a concentrated glucose solution through the side sampling port. The current signal was generated by employing an amperometric method with onboard reference and counter electrodes applying -0.1 V versus Ag / AgCl. The sensing performance was showcased with smoothed lines and calibration was calculated based on raw results. The last 100 seconds of the stabilized data were recorded as means ± SD for the current response at each biomarker concentration.

[0125] The current signal of the glucose sensor patch without MN in bath solution under fixed glucose concentrations for up to 20 minutes (1,200 seconds) was also measured, as illustrated in FIG.21. Measured glucose concentrations included 0.1 mM, illustrated along line 2000; 0.5 mM, illustrated along line 2002; 1 mM, illustrated along line 2004; 2 mM, illustrated along line 2006; 3 mM, illustrated along line 2008; 5 mM,illustrated along line 2010; 7 mM, illustrated along line 2012; 10 mM, illustrated along line 2014; and 20 mM, illustrated along line 2016. Following stabilization during the initial 120 seconds, the current values between 200-300 seconds are averaged as a reference point, and the relative current changes remained below 15% over the 20 minutes at all glucose levels, demonstrating the current signal's stability.

[0126] The anti-interference performance of the glucose sensor patch without MN in bath solution was also tested. Amperometry was applied in the presence of common easily oxidized species, such as UA for 0.1 mM uric acid, DA for 0.1 mM dopamine, and LA for 1 mM lactic acid, sugar such as Fru for 1 mM fructose, and Suc for 1 mM sucrose, and electrolyte such as NaCl for 1 mM sodium chloride (NaCl), as depicted in FIG.22A within graph 2100. The absolute current change in response to different species is also illustrated in FIG.22A within graph 2102, showing no significant variation with the addition of interference agents and demonstrating good anti-interference capability.

[0127] Additionally, the results of the electrochemical characterization were interpreted by calibrating the sensitivity plots for GOx and LOx using a Michaelis- Menten (MM)-inspired equation, as depicted in FIG.22B and 22C respectively. To reduce the size effect of the electrodes, the current signal was transferred as the absolute value of the current density. Unlike traditional theory for enzymatic sensors, here, many of the assumptions may not hold. In fact, while the traditional theory assumes an initial uniform and homogeneous distribution of molecules, here the analyte molecules (glucose G or lactate L) were introduced subsequently at increased concentrations. To account for this, a reframed version of the MM formalism's validity for the electrochemical setup was used: ^^ = ^^^^^^^^[^^^^^^]where q is the electron charge, NAis the Avogadro number, tezis the theoretical deposited thickness of the enzyme layer, [i] is the analyte concentration and Kiis theMM constant. The empirical parameter αiaccounts for deviations from the MM equation's ideal formalism, incorporating (a) the fraction of enzymatic active sites contributing to analyte conversion, (b) the catalytic rate constant efficiency, and (c) the capacitor partition between generated product fluxes. The latter process suppresses the efficiency of the overall reaction further; after being generated by the enzyme-analyte complex, only a limited fraction of product molecules approaches the sensing site (electrode) to produce an amperometric response, while the remaining amount is lost in the beaker solution. Equation 6 was used to calibrate the experimental data. The theory accurately reproduces the experiments, resulting in KG = 2 mM and KL = 3 mM, which is consistent with the expected range discussed in the literature. By using the MM equation for calibration rather than simple linear fitting, the sensor patch's effective monitoring range for target biomarkers (glucose or lactate) could be extended. Example 10: Long-term Stability and Anti-interference Tests

[0128] For the long-term stability test, a glucose sensor patch without MNs was fabricated on the glass slide and connected to the potentiostat. It was then submerged in a 1× PBS solution, stirred at 200 rpm. Following this, the glucose concentration was adjusted to specific values (0, 0.1, 0.5, 1, 2, 3, 5, 7, 10, 20 mM) by adding a 1 M glucose solution. An amperometric method, using onboard reference and counter electrodes applying -0.1 V versus Ag / AgCl, generated the current signal for more than 1,200 seconds. The sensing performance was represented using smoothed lines. The stabilized current data at 200-300 seconds were averaged as the reference current response Iref, while the relative current change was calculated by: ^^^^^^^^^^^^^^^^ = (^^^^^^^^ − ^^^^ref) / ^^^^refEquation 7 where Itis the current response at certain time t, and Irefis the reference current response by averaging the stabilized current data at 200-300 seconds. For the anti-interference test, a glucose sensor patch without MNs was fabricated on the glass slide and connected to the potentiostat. It was then submerged in a 1× PBS solution, stirred at 200 rpm.Subsequently, interference species, including 0.1 mM uric acid, 0.1 mM dopamine, 1 mM lactic acid, 1 mM fructose, 1 mM sucrose, and 1 mM NaCl, were added at 10-minute intervals, with 1 mM glucose added at the end. An amperometric method, using onboard reference and counter electrodes applying -0.1 V versus Ag / AgCl, generated the current signal. The sensing performance was represented using smoothed lines, with calculations based on raw results. The absolute value of the final 100 seconds of the stabilized data were recorded as mean ± SD for the absolute current response to each species. Example 11: In Vivo Evaluations on Mice

[0129] All animal experimental procedures were reviewed and approved by the Purdue Animal Care and Use Committee (PACUC) under protocol number 2001001998. Male NU / J athymic nude mice were procured from the Jackson Laboratory (Bar Harbor, ME, USA), with each mouse receiving one sensor patch at each experimental time point. All animals were maintained in a 12-hour light / dark cycle, with humidity and temperature kept constant. Animals were anesthetized using isoflurane vapor (5.0% for induction, 1.5-3.0% for maintenance) mixed with oxygen (flow rate: 500 mL min-1for induction, 30-70 mL min-1for maintenance), regulated by an automated system (SomnoSuite, Kent Scientific Corporation, Inc.). After initial induction, the animal was positioned in left lateral recumbency for anesthesia maintenance, allowing the sensor patch to be attached to the right lateral side of the dorsal skin while the abdomen remained accessible for intraperitoneal injections. The animal's temperature was kept at 35.9-37.5ºC using a heat pad. The sensor patch was attached to the right lateral dorsal skin and was ensured to puncture the skin using a spring- loaded application tool and gentle pinching with the index finger and thumb. It was then secured in place with Tegaderm tape (3M Tegaderm transparent film dressing). An incision was made to the tail tip using sterile surgical scissors (501739, World Precision Instruments, Inc.) and the tail was gently massaged to draw blood from the tail vein to measure blood glucose or lactate levels using a commercially available glucometer (Accu-Chek Guide Me, Roche Diabetes Care, Inc.) or a lactate meter (Lactate Plus, Nova Biomedical, Inc.). Simultaneously, the sensor patch was connected to the potentiostat, and an amperometric method with onboard reference and counterelectrodes applying -0.1 V versus Ag / AgCl was used for measurement. For the glucose measurement experiment, 2 U kg-1(~80 g kg-1) insulin in 200 µL PBS and 1 g kg-1glucose in 200 µL PBS were injected intraperitoneally. For the lactate measurement experiment, 1 g kg-1lactate in 200 µL PBS subcutaneously and PBS (as control) was injected subcutaneously directly underneath the site of the inserted sensory microneedle patch. Blood glucose or lactate levels were measured three times approximately every 10 minutes and averaged to serve as the reference value. The sensing performance was presented with smoothed lines, and calculations were performed based on raw results. For blood glucose levels measured at one specific point in time, the absolute values of current data for the subsequent 100 seconds were recorded as means ±SD for the current response. The set of data points was fitted with different signal delays to achieve the best fitting results.

[0130] The in vivo performance of the sensor patch on mice was assessed, comparing it to reference glucose and lactate levels measured from the tail vein using commercial meters and test strips. The measurements were taken 3 times at 10-minute intervals and then averaged to serve as the reference value for comparison to the absolute current signals obtained from the sensor patch by amperometric measurement.

[0131] FIG.23A displays the image of the sensor patch 2200 on the mouse skin 2, including a plurality of electrodes 2202, immediately upon the insertion of hydrogel MNs and a zoomed-in image 4 of one working electrode 2202 on mouse skin 2. FIG. 23B presents an image of the same area after the sensor patch has been removed following in vivo testing, revealing an array of swollen MNs without signs of detachment, fracture, or degradation. This observation implies that the stability of the system remains uncompromised, despite potential alterations to the hydrogel MNs. The mouse skin was imaged at 5 minutes after MNs removal and confirmed effective insertion by observing an array of pores without any bleeding.

[0132] FIG.23C shows the absolute value of the current signal 2204 and blood glucose level 2206 over time during glucose sensing test. Due to hyperglycemia in the mouse with a glucose level of 216 mg dL-1, insulin was injected intraperitoneally first at 600 seconds. Subsequently, glucose was injected intraperitoneally at 4,000 secondsand 7,000 seconds. FIG.23D displays a signal delay of 10 minutes that is obtained through fitting the absolute value of the current signal and the blood glucose level. The sensor patch can measure the glucose level in the range of 20-150 mg dL-1(1.11-8.32 mM), with a sensitivity of 0.00113 ± 0.00008 μA (mg dL-1)-1(or 0.020 ± 0.001 μA mM-1, R2= 0.917) within the linear range of 20-71 mg dL-1(1.11-3.94 mM). Table 1 below contrasts the glucose sensing performance of the device disclosed herein with that of other existing sensors. Table 1 Comparison of glucose sensing performance Linear range Sensitivity Test type Ref PBA-CNT-MNa0-16.7 mM 1.357 nA mM-1In vivo: rats 1 Au-Si-MNAb1-9 mM 0.1622 μA mM-1cm-2In vivo: mice 2 * Au / solid MNc3-24 mM 0.108 μA mM-1In vivo: mice 3 Au / Pt-black / NFd1-20 mM 4.38 μA mM-1cm-2 *In vivo: rats 4 FreeStyle Libre™ e1-30 mM 1.1-1.7 nA mM-1 **In vivo: human 5 1.11-3.94 mM 0.020 μA mM-1or In vivo: mice This work 1.8 μA mM-1cm-2aMicroneedle coated with composites of Phenylboronic acid-containing polymer and carbon nanotubes.bGold coated silicon microneedle array.cGold coated solid microneedle array.dPorous Platinum-black modified gold microneedles packed with Nafion.eCommercially available FreeStyle Libre™flash continuous glucose monitoring systems. * The sensitivity was calculated from in vitro test. ** The sensitivity was estimated from the FIG.1 of Hoss, U., & Budiman, E. S. (2017). Factory-calibrated continuous glucose sensors: The science behind the Technology. Diabetes Technology & Therapeutics, 19(S2), S-44-S-50. https: / / doi.org / 10.1089 / dia.2017.0025.

[0133] A strong correlation was evident between the glucose levels detected by the inventive sensor signals and those measured by commercial meters. Furthermore,the signal delay observed in the experiment aligns with both the theoretical prediction and the fluid uptake test.

[0134] FIG.23E illustrates the absolute value of the current signal 2208 and blood lactate level 2210 over time during lactate sensing test. Lactate was injected subcutaneously twice at 2,000 seconds and 8,000 seconds, while only the same amount of PBS solution was injected at 6,000 seconds as a control. The current signal displays a distinctive peak-response pattern that does not correspond to the blood lactate levels. The change in current response for both lactate injections was similar (0.028 ± 0.003 μA for the first lactate injection, and 0.027 ± 0.002 μA for the second lactate injection), and significantly higher than the change observed for the PBS injection (0.013 ± 0.002 μA), indicating the effective response of the sensor patch to lactate. Furthermore, the signal delay, (i.e., the time from lactate injection to the peak value of the current signal) was less than 5 minutes since the sensor patch is closer to the subcutaneously injected lactate.

[0135] A physics-based model for the amperometric response to subcutaneously injected lactate was created. In the model, the situation was represented as a stochastic process with a threshold, as illustrated in FIG.23F. The lactate is injected at a location (x0) distant to the sensing site (S, applied sensor patch). On the opposite side of the sensing unit, an absorption site (xb, bloodstream site) is responsible for absorbing the injected molecules. A flow velocity (u) is applied to account for the body's response to counteract the lactate injection and enhance its dissipation into the bloodstream. The equation used for the time-dependent lactate concentration at the sensing site (Ls) by combining balance equations and molecule conservation is as follows: ∂^^^^ ∂2^^^^ ∂^^^^ =− ^^^^

[0136] By applying the boundary conditions (L(x0, 0) = L0 δ(x – x0) and L(xb, t) = 0), the time-dependent concentration on the sensor surface in a 3D geometry was quantified:(^^^^ −^^^^^^^^)2 ^^^^ ^^^^ 0 − 0S(^^^^) ∝ ^^^^0e6^^^^body^^^^

[0137] Therefore, an empirical factor β was introduced, accounting for the relationship between analyte concentration and amperometric response (Ls to J conversion, and J for absolute value of current density): (^^^^ −^^^^^ )2 1− 0^^^^^^^th ∝ ^^^^S(^^^^) = ^^^^^^^^A^^^^e6^^^^body^^^^

[0138] The parameter β indicates the efficiency of the sensor patch in detecting the particular lactate analyte and is thus linked to the sensor's sensitivity and selectivity.

[0139] FIG.23F demonstrates the consistency between the experimental results and theoretical predictions. To calibrate the response outlined in Equation 10, the current signal curve 2208 depicted in FIG.23E was used. Additionally, the background noise stemming from drift-affected data by fitting it to an exponential relationship was removed, as shown in FIG.24. Assuming a 2 mm distance (x0 = 2 mm), the data using parameters outlined in Table S2 below was fitted, based on the noise-free response as illustrated by curve 2212 of FIG.23F. Table 2 Fitting parameters for amperometric response of the lactate sensor patch β (mol m-1) Dbody(m2s-1) u (m s-1) Lactate (L1, L2) 4.5 × 10-56 × 10-103 × 10-6PBS (S1) 2.5 × 10-52 × 10-91 × 10-6

[0140] Despite the intricacy of the in-vivo measurement, the theoretical prediction corresponded well with the experimental results.

[0141] While the embodiments of this disclosure have been shown and described as having preferred designs, the described embodiments may be modified within the spirit and scope of this disclosure. This application is therefore intended to cover any variations, uses or adaptations of the embodiments using general principles. Further, this application is intended to cover such departures from the present disclosure as come within known or customary practice in the art to which this disclosure pertains.

[0142] Various aspects are described in this disclosure, which include, but are not limited to, the following aspects:

[0143] (1) A wearable sensor patch for monitoring interstitial fluid, the sensor patch comprising: a base layer; a hydrogel body; and a sensor element comprising: a conductive layer embedded within a polydimethylsiloxane (PDMS) body, the conductive layer comprising: a working electrode comprising thermoplastic polyurethane (TPU) and a first conducting material; a serpentine-shaped counter electrode comprising thermoplastic polyurethane (TPU) and a second conducting material; and a reference electrode comprising thermoplastic polyurethane (TPU) and a third conducting material; wherein the sensor element is embedded within the hydrogel body.

[0144] (2) The wearable sensor patch of aspect 1, wherein the patch is configured to recoverably deform or elongate in response to user movement or swelling from the uptake of interstitial fluid, without delaminating or breaking any one of the base layer, the hydrogel body, the sensor element and a combination thereof.

[0145] (3) The wearable sensor patch of aspect 1 or aspect 2, wherein the base layer comprises hydrogel microneedles.

[0146] (4) The wearable sensor patch of aspect 3, wherein the hydrogel microneedles are integrated with the hydrogel body such that the hydrogel microneedles and the hydrogel body form one homogenous hydrogel structure.

[0147] (5) The wearable sensor patch of aspect 4, wherein the hydrogel microneedles and hydrogel body both comprise methacrylated hyaluronic acid (MeHA).

[0148] (6) The wearable sensor patch of any one of aspects 1 to 5, wherein at least one of the first conducting material and the second conducting material comprises silver flakes.

[0149] (7) The wearable sensor patch of any one of aspects 1 to 6, wherein the third conducting material comprises at least one of silver and silver chloride.

[0150] (8) The wearable sensor patch of any one of aspects 1 to 7, wherein the hydrogel body comprises a material selected from the group consisting of polyvinyl alcohol (PVA), chitosan, osmolytes, and methacrylated hyaluronic acid (MeHA).

[0151] (9) The wearable sensor patch of any one of aspects 1 to 8, further comprising an enzyme layer which contacts the working electrode, counter electrode, and reference electrodes.

[0152] (10) The wearable sensor patch of any one of aspects 1 to 9, wherein the enzyme layer is formed from a solution of oxidase enzyme, bovine serum albumin (BSA), and a crosslinking agent.

[0153] (11) A method of manufacturing a wearable sensor patch for monitoring of interstitial fluid, the method comprising: printing a sensor element onto a substrate; encapsulating said sensor element in a hydrogel body; and curing the hydrogel body with a plurality of hydrogel microneedles.

[0154] (12) The method of aspect 11, wherein curing the hydrogel body with the plurality of hydrogel microneedles causes the hydrogel body and the hydrogel microneedles to form a single homogenous hydrogel structure.

[0155] (13) The method of aspect 12, wherein the hydrogel microneedles and the body comprise methacrylated hyaluronic acid (MeHA)

[0156] (14) The method of any one of aspects 11 to 13, wherein the hydrogel microneedles are formed by pouring a liquid hydrogel solution into a mold, the method further comprising: before the liquid hydrogel solution is fully cured, placing the hydrogel body encapsulating said sensor element on top of the hydrogel solution.

[0157] (15) The method of any one of aspects 11 to 14, wherein the printing of the sensor element comprises: printing a first ink comprising a first conducting material and thermoplastic polyurethane (TPU) onto a substrate to form a working electrode; printing a second ink comprising a second conducting material and thermoplasticpolyurethane (TPU) onto said substrate to form a counter electrode; and printing a third ink comprising a third conducting material and thermoplastic polyurethane (TPU) onto said substrate to form a reference electrode.

[0158] (16) The method of aspect 15, wherein at least one of the working electrode, the counter electrode, and the reference electrode are configured to recoverably deform or elongate without breaking.

[0159] (17) The method of aspect 15 or aspect 16, wherein at least one of the working electrode, the counter electrode, and the reference electrode comprise a portion that is serpentine-shaped to facilitate recoverable deformation or elongation.

[0160] (18) The method of any one of aspects 15 to 17, wherein the substrate comprises polyvinyl alcohol (PVA).

[0161] (19) The method of any one of aspects 15 to 18, wherein the hydrogel body is formed by pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a polydimethylsiloxane (PDMS) mold.

[0162] (20) The method of any one of aspects 15 to 19, wherein the hydrogel microneedles are formed by:pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a mold; extracting the hydrogel microneedles from the mold and curing them with ultraviolet radiation.

[0163] (21) An electrode comprising thermoplastic polyurethane (TPU) embedded with a silver material, wherein the electrode is defined in a serpentine shape.

[0164] (22) The electrode of aspect 21, wherein the silver comprises silver flakes.

[0165] (23) The electrode of any one of aspects 21 or 22, wherein the electrode is configured to recoverably deform or elongate.

Claims

CLAIMS 1. A wearable sensor patch for monitoring interstitial fluid, the sensor patch comprising: a base layer; a hydrogel body; and a sensor element comprising: a conductive layer embedded within a polydimethylsiloxane (PDMS) body, the conductive layer comprising: a working electrode comprising thermoplastic polyurethane (TPU) and a first conducting material; a serpentine-shaped counter electrode comprising thermoplastic polyurethane (TPU) and a second conducting material; and a reference electrode comprising thermoplastic polyurethane (TPU) and a third conducting material; wherein the sensor element is embedded within the hydrogel body.

2. The wearable sensor patch of claim 1, wherein the patch is configured to recoverably deform or elongate in response to user movement or swelling from uptake of interstitial fluid, without delaminating or breaking any one of the base layer, the hydrogel body, the sensor element and a combination thereof.

3. The wearable sensor patch of claim 1 or claim 2, wherein the base layer comprises hydrogel microneedles.

4. The wearable sensor patch of claim 3, wherein the hydrogel microneedles are integrated with the hydrogel body such that the hydrogel microneedles and the hydrogel body form one homogenous hydrogel structure.

5. The wearable sensor patch of claim 4, wherein the hydrogel microneedles and hydrogel body both comprise methacrylated hyaluronic acid (MeHA).

6. The wearable sensor patch of any one of claims 1 to 5, wherein at least one of the first conducting material and the second conducting material comprises silver flakes.

7. The wearable sensor patch of any one of claims 1 to 6, wherein the third conducting material comprises at least one of silver and silver chloride.

8. The wearable sensor patch of any one of claims 1 to 7, wherein the hydrogel body comprises a material selected from the group consisting of polyvinyl alcohol (PVA), chitosan, osmolytes, and methacrylated hyaluronic acid (MeHA).

9. The wearable sensor of any one of claims 1 to 8, further comprising an enzyme layer which contacts the working electrode, counter electrode, and reference electrodes.

10. The wearable sensor patch of claim 9, wherein the enzyme layer is formed from a solution of oxidase enzyme, bovine serum albumin (BSA), and a crosslinking agent.

11. A method of manufacturing a wearable sensor patch for monitoring of interstitial fluid, the method comprising: printing a sensor element onto a substrate; encapsulating said sensor element in a hydrogel body; and curing the hydrogel body with a plurality of hydrogel microneedles.

12. The method of claim 11, wherein curing the hydrogel body with the plurality of hydrogel microneedles causes the hydrogel body and the hydrogel microneedles to form a single homogenous hydrogel structure.

13. The method of claim 12, wherein the hydrogel microneedles and the body comprise methacrylated hyaluronic acid (MeHA)14. The method of any one of claims 11 to 13, wherein the hydrogel microneedles are formed by pouring a liquid hydrogel solution into a mold, the method further comprising: before the liquid hydrogel solution is fully cured, placing the hydrogel body encapsulating said sensor element on top of the hydrogel solution.

15. The method of any one of claims 11 to 14, wherein the printing of the sensor element comprises: printing a first ink comprising a first conducting material and thermoplastic polyurethane (TPU) onto a substrate to form a working electrode; printing a second ink comprising a second conducting material and thermoplastic polyurethane (TPU) onto said substrate to form a counter electrode; and printing a third ink comprising a third conducting material and thermoplastic polyurethane (TPU) onto said substrate to form a reference electrode.

16. The method of claim 15, wherein at least one of the working electrode, the counter electrode, and the reference electrode are configured to recoverably deform or elongate without breaking.

17. The method of claim 15 or claim 16, wherein at least one of the working electrode, the counter electrode, and the reference electrode comprise a portion that is serpentine-shaped to facilitate recoverable deformation or elongation.

18. The method of any one of claims 15 to 17, wherein the substrate comprises polyvinyl alcohol (PVA).

19. The method of any one of claims 15 to 18, wherein the hydrogel body is formed by pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a polydimethylsiloxane (PDMS) mold.

20. The method of any one of claims 15 to 19, wherein the hydrogel microneedles are formed by: pouring a solution comprising methacrylated hyaluronic acid (MeHA) and a photoinitiator into a mold; extracting the hydrogel microneedles from the mold; and curing the hydrogel microneedles with ultraviolet radiation.

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