Electrochemical hydrogel microneedle array biosensors

Abstract

This disclosure provides a wearable electrochemical biosensor for detecting at least one target analyte in interstitial fluid of a subject, comprising, a microneedle array comprising a highly- crosslinked hydrogel polymer; at least one electrochemical chip comprising at least one transducer surface configured for generating a signal in the presence of the at least one target analyte; and a low crosslinked hydrogel polymer, wherein the low crosslinked hydrogel polymer is configured to couple the microneedle array with the at least one transducer surface of the electrochemical chip. Methods of making and use thereof are also disclosed.

Description

ELECTROCHEMICAL HYDROGEL MICRONEEDLE ARRAY BIOSENSORS CROSS-REFERENCE TO RELATED APPLICATION

[0001] This application claims the benefit of United States Provisional PatentApplication No. 63 / 645,309 filed May 10, 2024, herein incorporated by reference in its entirety. SEQUENCE LISTING

[0002] A computer readable form of the sequence listing“P96459305PCT00_SequenceListing” (5,110 bytes) was created on May 7, 2025, is filedherewith by electronic submission and is incorporated by reference herein.FIELD

[0003] The present disclosure relates to biosensors, and in particular, electrochemicalbiosensors, methods of making and uses thereof for target analyte detection. BACKGROUND

[0004] Transdermal biosensing has the potential to revolutionize healthcare byenabling real-time and continuous monitoring of clinically-important biomarkers in vivo and inside dermal interstitial fluid (ISF).[1–3]Commercially-available continuous glucose monitors (CGMs) are examples of such systems. CGMs eliminate the need for patients to conduct frequent blood sampling and manual measurement of their blood glucose level; instead CGMs enable continuous ISF glucose measurement and blood glucose prediction every 5 minutes using a wearable patch that features a solid metallic needle for electrochemical analysis.[4,5]The long metallic microneedles (5 mm to 13 mm) used in existing CGMs cause user discomfort, hindering the adoption of CGMs by end users.[6,7]There is a need for a new class of transdermal bioanalytical devices that are comfortable to wear, inherently more stable than enzymatic devices, and applicable to a wide range of analytes and multiplexed panels, beyond just glucose or other enzyme-linked targets.[5,8]

[0005] Microneedle (MN) based biosensors are a class of wearable devices thatintegrate minimally invasive ISF extraction using short and painless microscale needles (<1.5 mm in length) with transdermal bioanalysis.[9]Many of the existing MN biosensors use solid and inflexible MNs made from conductive polymer (e.g., poly(methyl methacrylate)) or metals (e.g., stainless steel and gold) as electrodes for bioanalysis inside ISF.[1,10–12]Such solid MNsare not compatible with mechanically soft skin, lack biocompatibility, and pose a potential risk of needle breakage within the skin.

[0013] Solid MNs are typically surface functionalized with enzymes or aptamers, which are prone to detachment during the skin insertion process, resulting in a significant reduction in measurement accuracy and device lifetime.[1,10,12]Additionally, given the incompatibility of solid and inflexible MNs with soft and flexible skin, they experience physical displacement during in vivo measurements, causing severe sensor and measurement instability.

[0014] Such instability has been alleviated by magnetic placement of MN sensors in animal models

[0014] ; however, this technique necessitates the insertion of a metallic plate under the skin, making it inapplicable to human patients. Despite the existence of solid MN biosensors, there is a need for a new class of skin-compatible yet mechanically-robust and analytically-stable MN technology for continuous monitoring in vivo.

[0006] The background herein is included solely to explain the context of thedisclosure. This is not to be taken as an admission that any of the material referred to was published, known, or part of the common general knowledge as of the priority date. SUMMARY

[0007] The present disclosure includes an integrated system fabricated by combiningbiocompatible hydrogel microneedle (HMN) arrays for ISF extraction with an electrochemical aptamer-based biosensor for in situ monitoring of ISF analytes. The use of aptamers enables continuous monitoring of a wide range of analytes, beyond what is possible with enzymatic monitoring. This wearable aptamer-based biosensor enables minimally invasive multiplexed monitoring of biomarkers, such as glucose and lactate in ISF, in real time. With effective skin penetration and integration into functionalized electrodes, it offers high sensitivity for clinical detection of these biomarkers. Herein, validation experiments using live healthy mice and Type 1 Diabetes rat models demonstrate strong correlation between the ISF measurements collected from the wearable aptamer-based biosensor of the disclosure and those obtained from gold- standard techniques for blood glucose and lactate, for each analyte alone and in combination. Demonstration of stable signals and extended lactate monitoring in mouse and rat models indicates the wearable aptamer-based biosensor’s potential for personalized health monitoringand treatment approaches, such as that for diabetes care. The wearable aptamer-based biosensordisclosed herein effectively addresses the limitations inherent in enzymatic detection methods aswell as solid MN biosensors and addresses the need for reliable and multiplexed bioanalytical monitoring in vivo.

[0008] Accordingly, in an aspect of the present disclosure, provided herein is anelectrochemical biosensor for detecting at least one target analyte in interstitial fluid of a subject, comprising: a) a microneedle array comprising a highly-crosslinked hydrogel polymer;b) at least one electrochemical chip comprising at least one transducer surfaceconfigured for generating a signal in the presence of the at least one target analyte; and c) a low crosslinked hydrogel polymer,wherein the low crosslinked hydrogel polymer is configured to couple the microneedle array with the at least one transducer surface of the at least one electrochemical chip.

[0009] In some embodiments, the biosensor further comprises one or more probesfunctionalized on the at least one transducer surface that bind to the at least one target analyte, the one or more probes configured to generate a measurable signal in the presence of the at least one target analyte.

[0010] In some embodiments, the transducer surface is configured for reacting with theat least one target analyte to generate a measurable signal in the presence of the at least one target analyte.

[0011] In some embodiments, the biosensor further comprises (i) a first probe that bindsto a first target analyte functionalized on a first transducer surface and (ii) the first probe and a second probe that binds to a second target analyte functionalized on a second transducer surface, each transducer surface configured to generate measurable signals, wherein a difference between the measurable signal on the first transducer surface and the measurable signal on the second transducer surface is indicative of the presence of the second target analyte.

[0012] In some embodiments, the highly crosslinked hydrogel polymer has acrosslinking density that is at least 10% greater than that of the low-crosslinked hydrogel polymer.

[0013] In some embodiments, the highly crosslinked hydrogel polymer has acrosslinking density that is at least 20% greater than that of the low-crosslinked hydrogel polymer.

[0014] In some embodiments, the highly crosslinked hydrogel polymer has acrosslinking density that is at least 30% greater than that of the low-crosslinked hydrogel polymer.

[0015] In some embodiments, the highly crosslinked hydrogel polymer has acrosslinking density that is at least 40% greater than that of the low-crosslinked hydrogel polymer.

[0016] In some embodiments, the highly-crosslinked and the low crosslinked hydrogelpolymer, each independently, comprise gelatin, hyaluronic acid, alginate, chitosan, collagen, or combinations thereof.

[0017] In some embodiments, the highly-crosslinked and the low crosslinked hydrogelpolymer, each independently, comprise methacrylated gelatin, methacrylated hyaluronic acid, methacrylatedalginate, methacrylated chitosan, methacrylated collagen, dopamine- functionalized hyaluronic acid or combinations thereof.

[0018] In some embodiments, the highly-crosslinked and the low crosslinked hydrogelpolymer, each independently, comprise methacrylated hyaluronic acid.

[0019] In some embodiments, the highly-crosslinked and the low crosslinked hydrogelpolymer, each independently, further comprise a crosslinking agent.

[0020] In some embodiments, the crosslinking agent comprises N,Nʹ-methylenebisacrylamide (MBA), PEG–maleimide, HA–thiol, enzymes, Carbodiimide-based (EDC / NHS), or combinations thereof.

[0021] In some embodiments, the highly-crosslinked hydrogel polymer comprises aporosity of about 25% to about 50%.

[0022] In some embodiments, the low crosslinked hydrogel polymer comprises aporosity of above about 50% up to about 75%.

[0023] In some embodiments, the at least one electrochemical chip comprises one ormore working electrodes.

[0024] In some embodiments, the at least one electrochemical chip further comprises acounter electrode.

[0025] In some embodiments, the at least one electrochemical chip further comprises areference electrode.

[0026] In some embodiments, the counter electrode is a reference electrode.

[0027] In some embodiments, the at least one transducer surface comprises metals,metal alloys, metal oxides, superconductors, semi-conductors, carbon-based materials, conductive polymers, or combinations thereof.

[0028] In some embodiments, the at least one transducer surface comprises metals,metal alloys, carbon-based materials, or combinations thereof.

[0029] In some embodiments, the transducer surface comprises gold.

[0030] In some embodiments, the one or more probes is functionalized on the at leastone transducer surface via chemical bonding, an intermediate linker, physical adsorption, embedded in inks, embedded in polymers, or a combination thereof.

[0031] In some embodiments, the chemical bonding occurs via thiol chemistry, goldchemistry, or a combination thereof.

[0032] In some embodiments, the one or more probes is a nucleic acid probe, anantibody probe, or a combination thereof.

[0033] In some embodiments, the one or more nucleic acid probes comprises a single-stranded DNA, single-stranded RNA, double-stranded DNA, double stranded RNA, DNAzyme, an aptamer, or a combination thereof.

[0034] In some embodiments, the one or more nucleic acid probes comprises anaptamer.

[0035] In some embodiments, the one or more nucleic acid probes comprises adetection moiety.

[0036] In some embodiments, the detection moiety comprises an electroactive species,a redox active species, or a combination thereof.

[0037] In some embodiments, the redox active species comprises methylene blue.

[0038] In some embodiments, the one or more antibody probes comprises an anti-insulin antibody, an anti-troponin antibody, an anti-bovine serum albumin antibody, or a combination thereof.

[0039] In some embodiments, the one or more antibody probes comprises a detectionmoiety.

[0040] In some embodiments, the detection moiety comprises an electroactive species,a redox active species, or a combination thereof.

[0041] In some embodiments, the redox active species comprises methylene blue.

[0042] In some embodiments, the signal is a current, a potential or an impedance.

[0043] In some embodiments, the at least one target analyte is a biomolecule.

[0044] In some embodiments, the at least one target analyte is glucose, lactate, or acombination thereof.

[0045] In some embodiments, the microneedles are arranged in a wearable patchconfigured for transdermal sensing.

[0046] In some embodiments, the biosensor disclosed herein is for use in healthmonitoring, screening, diagnostics, or a combination thereof.

[0047] Also provided is a method for transdermal biosensing of at least one targetanalyte in interstitial fluid of a subject, comprising: a) applying a transdermal patch comprising the biosensor disclosed herein onto thesubject; and b) detecting the signal from the at least one electrochemical chip of the biosensor,whereby the at least one transducer surface generates an electrochemical signal in the presence of the at least one target analyte, wherein the electrochemical signal is indicative of the concentration of the at least one target analyte in the subject.

[0048] In some embodiments, the signal is monitored continuously over time.

[0049] In some embodiments, the signal comprises a current, a potential or animpedance.

[0050] In some embodiments, the signal comprises current.

[0051] In some embodiments, the signal is measured using square wave voltammetry.

[0052] In some embodiments, the signal is used to calculate a kinetic differentialmeasurement (KDM).

[0053] Also provided herein is a method for fabricating the electrochemical biosensordisclosed herein, the method comprising: a) providing a microneedle array comprising a highly-crosslinked hydrogelpolymer, the microneedle array prepared by drying the highly-crosslinked hydrogel polymer in a mold; b) providing at least one electrochemical chip comprising at least one transducersurface; and c) coupling the microneedle array with the at least one transducer surface of theelectrochemical chip using a low crosslinked hydrogel polymer.

[0054] In some embodiments, the method further comprises adjusting pH of the highly-crosslinked hydrogel polymer before drying in the mold.

[0055] In some embodiments, the method further comprises exposing the highly-crosslinked hydrogel polymer to ultraviolet (UV) light after drying in the mold.

[0056] In some embodiments, the method further comprises crosslinking the highly-crosslinked and low crosslinked hydrogel polymers after coupling the microneedle array with the at least one transducer surface of the at least one electrochemical chip.

[0057] In some embodiments, the method further comprises calibrating the biosensorby measuring a signal upon increasing and decreasing target analyte concentrations.

[0058] In some embodiments, a kinetic differential measurement (KDM) is calculatedto generate a calibration curve.

[0059] Other features and advantages of the present disclosure will become apparentfrom the following detailed description. It should be understood, however, that the detailed description and the specific examples, while indicating embodiments of the disclosure, are given by way of illustration only and the scope of the claims should not be limited by these embodiments, but should be given the broadest interpretation consistent with the description as a whole.BRIEF DESCRIPTION OF THE DRAWINGS

[0060] Certain embodiments of the disclosure will now be described in greater detailwith reference to the attached drawings in which:

[0061] FIG. 1 shows fabrication of the wearable aptamer-based biosensor for in vivoanalysis in exemplary embodiments of the disclosure: (a) schematic illustrating the in vivo sensing experimental set-up, with a zoomed-in cross-sectional view of the wearable aptamer- based biosensor; (b) schematic showing the fabrication process for the wearable aptamer-based biosensor; (c) magnified view of the components of the wearable aptamer-based biosensor, with a schematic showing the mechanism of signal transduction before and after target capture – upon target binding, the redox aptamer changes its conformation, increasing the electron transfer rate between the redox reporter and the gold surface; the SWV peak current increases (“signal on” behavior) in response to target analyte binding.

[0062] FIG. 2 shows an 1H NMR spectrum of MeHA in an exemplary embodiment ofthe disclosure: MeHA was characterized with 300MHz1H NMR with a 10 ms-time scale to determine the degree of methacrylate modification.

[0063] FIG. 3 shows fabrication and characterization of a wearable aptamer-basedbiosensor in exemplary embodiments of the disclosure: (a) schematic showing the different steps of wearable aptamer-based biosensor fabrication: (i) bio-functionalized electrodes were created by depositing 2 ^M methylene blue-tagged aptamer (glucose / lactate) onto the working electrodes, (ii) surface passivation was achieved by depositing 100 mercaptohexanol (MCH) for 10 minutes, (iii) attachment of HMN on the electrodes were done by adding LC MeHA film on the electrodes; EIS measurements were performed at each of the steps described above for the (b) glucose and (c) lactate wearable aptamer-based biosensors; cyclic voltammetry (CV) scans were performed after each step for the (d) glucose and (e) lactate wearable aptamer-based biosensors – all measurements were performed in 2 mM [Fe(CN)6]3- / 4-in 0.05 M KCl and 0.1 M Phosphate buffer solution (PBS).

[0064] FIG. 4 shows kinetic differential measurement (KDM) characterizationperformed by calculating the electrochemical signal change measured at different frequencies using square wave voltammetry in a potential range of 0 to -0.6 V with a scan rate of 0.1 V / s, before and after introducing the target analyte, in exemplary embodiments of the disclosure:(a) glucose and (b) lactate aptamer probes yielded two distinct behaviors of a ‘signal-on’ response at higher frequencies and a ‘signal-off’ response at lower frequencies (error bars show the standard deviation from the mean obtained using three (n=3) separate sensors).

[0065] FIG. 5 shows physical characterization of HMN patches and the integratedwearable aptamer-based biosensor in exemplary embodiments of the disclosure: (a) SEM images of HC+LC-MeHA HMN array (scale bar = 1 mm) and a single needle (inset, scale bar = 150 ^^m); (b) SEM images of the swelled (i) HC-MeHA hydrogel film, and (ii) LC-MeHA hydrogel film, scale bars for both images represent 100 ^^m; (c) swelling ratio of HC+LC- MeHA HMNs after applying on the porcine skin for different time durations (data expressed as mean ± standard deviation for three HMN patches (n=3) per condition); (d) mechanical compression testing of HC-MeHA HMN and HC+LC-MeHA HMN array (one HMN (n=1) was tested per condition); (e) images showing the wearable aptamer-based biosensor device upon integration (top, scale bar = 1 cm) and a zoomed-in view of the attached HMN patch (bottom, scale bar = 4 mm); (f) schematic of the pulling experimental set-up – a HC-MeHA thin film was attached to the electrode through a layer of LC-MeHA and a lateral pulling force was applied to the HC-MeHA thin film through an inelastic tape; (g) pulling force versus displacement of thin film for samples fabricated using 50 ^^L or 100 ^^L of LC-MeHA (one HMN (n=1) was tested per condition).

[0066] FIG. 6 shows ex vivo validation of the wearable aptamer-based biosensor inexemplary embodiments of the disclosure: electrochemical signals were obtained at two different frequencies upon introducing artificial interstitial fluid (aISF) spiked with different concentrations of (a) glucose (5 Hz and 80 Hz) and (b) lactate (15 Hz and 200 Hz), and the KDM was calculated to derive the upward and downward calibration curves; sensor response was calculated after inserting the fabricated sensors on porcine ear skins infused with different concentrations of (c) glucose and (d) lactate - the inset images show the experimental set-up; the specificity of the integrated wearable aptamer-based biosensor for (e) glucose and (f) L- lactate was tested against common interferences found in ISF by calculating the electrochemical signal gain (ΔI / I) measured at 60 Hz, before and after adding the target analyte – electrochemical signal changes were calculated before and after target incubation (significance was shown through asterisks corresponding to ****P < 0.0001; experiments were performed at n = 3 and data was expressed as mean ± standard deviation); all SWV measurements were performed in aISF solution, against an Ag / AgCl reference electrode andAu counter electrode in a potential range of 0 to -0.5 V and a scan rate of 0.1 V s-1; error bars show the standard deviation from the mean obtained using three (n=3) separate sensors.

[0067] FIG. 7 shows a log-linear regression model was applied to the ex vivocalibration curves of the wearable aptamer-based biosensor shown in FIG. 6 in exemplary embodiments of the disclosure: (a) linear regression equations of ΔI (%) = 39.657 ln(C) – 25.293 and ΔI (%) = 18.644 ln(C) + 30.316 with coefficient of determination of 0.9965 and 0.9830 were calculated for increasing and decreasing glucose concentrations; (b) linear regression equations of ΔI (%) = 23.854 ln(C) + 20.963 and ΔI (%) = 41.835 ln(C) – 24.258 with coefficient of determination of 0.9907 and 0.9652 were calculated for increasing and decreasing lactate concentrations – all SWV measurements were performed against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.5 V and a scan rate of 0.1 V s-1(error bars show the standard deviation from the mean obtained using three (n=3) separate sensors).

[0068] FIG. 8 shows the kinetic profiles of glucose and lactate aptamer-basedbiosensor under increasing and decreasing target concentrations in an exemplary embodiment of the disclosure: initially, 2.5 mM of glucose and 1 mM of lactate were introduced to the respective aptamer-based biosensor, followed by a 20-minute stabilization time, the concentrations were increased to 10 mM for glucose and 5 mM for lactate, and the electrochemical signal was recorded at signal-on and signal-off frequencies every 6 minutes for 30 minutes; the solution was then diluted back to the initial concentrations, and scanning was continued for an additional duration – all SWV measurements were performed in aISF solution, against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.5 V and a scan rate of 0.1 V s-1(error bars represent the standard deviation from the mean obtained using three (n=3) separate sensors).

[0069] FIG. 9 shows the kinetic profiles of the lactate aptamer-based biosensor on theporcine skin model in an exemplary embodiment of the disclosure: lactate aptamer-based biosensors were applied on porcine skin containing 5 or 20 mM of lactate and the electrochemical signal was measured after 5, 10, 15, 20, 30, and 40 minutes of incubation – all SWV measurements were performed against an Ag / AgCl reference electrode and an Aucounter electrode in a potential range of 0.0 to -0.5 V at a scan rate of 0.1 V s-1(error bars show the standard deviation from the mean obtained using three (n=3) separate sensors).

[0070] FIG. 10 shows quality control assessment for the fabricated wearable aptamer-based biosensor using four devices (S1-S4) in exemplary embodiments of the disclosure: (a) impedance measurements using electrochemical impedance spectroscopy (EIS) following aptamer functionalization and MN attachment; (b) ΔZ / Z, representing the changes in impedance after MN attachment compared to the impedance following the biofunctionalization of the chip with aptamer, calculated at 4 different frequencies including 10, 31.6, 100, and 1000 Hz to identify faulty attachment of MNs to electrodes – faulty attachment was defined as incomplete attachment of the HMN patches to the electrodes or imperfect insertion of the HMN patches to the skin; (c) ΔZ / Z measured at 31.6 Hz to determine the cut-off ratio for quality control of 27 glucose and 22 lactate sensors, comprising 14 glucose and 11 lactate sensors with faulty MN attachment; (d) ΔZ / Z measured at 31.6 Hz under the cut-off ratio of ΔZ / Z=500 for glucose and lactate sensors to eliminate faulty sensors having a ΔZ / Z>500 before performing analytical measurements – EIS measurement was performed in 2 mM [Fe(CN)6]3- / 4-in 0.05 M KCl and 0.1 M PBS solution at open circuit potential.

[0071] FIG. 11 shows in vivo glucose and lactate detection in a mouse model usingtwo wearable aptamer-based biosensor sensors applied to the dorsal skin of a mouse to track the level of glucose (using three mice) or lactate (using three mice) in exemplary embodiments of the disclosure: (a) image of a mouse under anesthesia with 3% isoflurane (the two lactate sensors were applied to the shaved back skin of the mouse); (b) image of a swelled HMN after skin penetration (scale bar = 3 mm); (c) histological section of mouse skin after wearable aptamer-based biosensor penetration – the dash-lined area indicates the needle cavity generated by needle penetration (scale bar = 100 µm); (d) to (f) glucose levels at various time points using average measurements from two different sensors (n=2) after each mouse received an injection of 3 g kg-1of glucose (i.p.) after baseline measurement, followed by a 0.75 U kg-1injection of human recombinant insulin (i.p.) at t = 60 minutes to reduce the blood glucose level; (g) to (i) L-lactate level average measurements using two applied sensors (n=2) after intraperitoneal injection with 0.6 g kg-1of sodium L-lactate was performed for each mouse to increase the blood L-lactate level – blood samples were collected from the tail three times (n=3) and one time (n=1) at each time point to measure glucose and L-lactate levels using a glucometer anda colorimetric lactate kit, respectively – data obtained from glucometer was expressed as mean ± standard deviation; all SWV measurements were done against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.5 V and a scan rate of 0.1 V s-1.

[0072] FIG. 12 shows in vivo characterization and multiplexed monitoring using thewearable aptamer-based biosensor in exemplary embodiments of the disclosure: (a) images of a wearable aptamer-based biosensor before and after skin penetration (scale bar = 4 mm); (b) magnified images showing the recovery process of the micropore trace on the rat skin acquired after 0 min, 5 min, 10 min, 15 min, 20 min, and 30 min of patch removal; (c) histological section of the rat skin after wearable aptamer-based biosensor penetration – dash-lines indicate the needle cavity generated by needle penetration (scale bar = 100 ^^m); (d) image illustrating the in vivo sensing experimental set-up; (e) long-term stability test of the wearable aptamer- based biosensor performed on rats throughout 3-days of device application – bar graphs represent lactate levels measured using three wearable aptamer-based biosensor devices, and grey circles indicate the blood L-lactate levels measured by benchtop L-lactate kit using collected blood samples; (f) stability of the wearable aptamer-based biosensor under multiple electrochemical scans – each electrode was continuously scanned 100 times using SWV (data is expressed as mean ± standard deviation of three devices per condition); (g) to (i) multiplexed measurements performed by two glucose and L-lactate wearable aptamer-based biosensor in three different diabetic rats – data was collected once rat blood glucose level reached specific target ranges (T1: 15 – 20 mM, T2: 15 – 10 mM, T3: 5 – 10 mM and T4: < 5 mM); the measurements collected from two glucose / L-lactate wearable aptamer-based biosensors were averaged; all SWV measurements were done against an Ag / AgCl reference electrode and an Au counter electrode in a potential range of 0 to -0.5 V and a scan rate of 0.1 V s-1.

[0073] FIG. 13 shows in vivo multiplexing validation with distinct data points fromtwo individual glucose and lactate wearable aptamer-based biosensors in three different diabetic rats in exemplary embodiments of the disclosure – data was collected once the rat blood glucose level reached specific target ranges (T1: 15 – 20 mM, T2: 15 – 10 mM, T3: 5 – 10 mM and T4: < 5 mM).

[0074] FIG. 14 shows characterization of the wearable biosensor with aptamers andantibodies. a) Schematic representation of biosensor using two electrodes and both antibodiesand aptamers; b) Current measured using the cardiac troponin I (cTnI) WA at 80 Hz and normalized to 20 Hz (I80 / I20), with the hydrogel microneedle exposed to a solution containing 4 mM lactate and various cTnI concentrations. The 0 ng / L indicates lactate only; c) cTnI and Lactate WA response in mouse model of coronary artery disease; d) Plasma cTnI level and differential signal response from WA of individual tested mouse.

[0075] FIG. 15A shows a schematic of the individual components of the biosensor.

[0076] FIG. 15B shows a schematic of the biosensor with a low crosslinked hydrogelpolymer applied on a transducer surface of an electrochemical chip.

[0077] FIG. 15C shows a schematic of the assembled electrochemical biosensor.DETAILED DESCRIPTION I. Definitions

[0078] Unless otherwise indicated, the definitions and embodiments described in thisand other sections are intended to be applicable to all embodiments and aspects of the present disclosure herein described for which they are suitable as would be understood by a person skilled in the art. It is also to be understood that the terminology used herein is for the purpose of describing particular aspects only and is not intended to be limiting.

[0079] The term "sample" or "test sample" as used herein may refer to any material inwhich the presence or amount of a target analyte is unknown and can be determined in an assay. The sample may be from any source, for example, any biological (e.g. human or animal samples, including clinical samples), environmental (e.g. water, soil or air) or natural (e.g. plants) source, or from any manufactured or synthetic source (e.g. food or drinks). The sample may be comprised or is suspected of comprising one or more analytes. The sample may be a "biological sample" comprising cellular and non-cellular material, including, but not limited to, tissue samples, urine, blood, serum, interstitial fluids, other bodily fluids and / or secretions.

[0080] The term “subject” as used herein includes all members of the animal kingdomincluding mammals such as a mouse, a rat, a dog, a cattle, a sheep, a horse, and a human. Thus, the methods and uses of the present disclosure are applicable to both human therapy and veterinary applications.

[0081] The term “target”, “analyte” or “target analyte” as used herein may refer to anyagent, including, but not limited to, a small inorganic molecule, small organic molecule, metal ion,biomolecule, toxin, biopolymer (such as a nucleic acid, carbohydrate, lipid, peptide, protein), cell, tissue, microorganism and virus, for which one would like to sense or detect. The analyte may be either isolated from a natural source or is synthetic. The analyte may be a single compound or a class of compounds, such as a class of compounds that share structural or functional features. The term analyte also includes combinations (e.g. mixtures) of compounds or agents such as, but not limited, to combinatorial libraries and samples from an organism or a natural environment.

[0082] The term “nucleic acid” as used herein refers to a biopolymer comprisingmonomers of nucleotides, such as deoxyribonucleic acid (DNA), ribonucleic acid (RNA) and other polynucleotides of modified nucleotides and / or nucleotide derivatives, and may be either double stranded (ds) or single stranded (ss). Nucleic acids may be modified to contain modified nucleotides comprising one or more modified bases (e.g. unusual bases such as inosine, and functional modifications to the bases such as amino), modified backbones (e.g. peptide nucleic acid, PNA) and / or other chemically, enzymatically, or metabolically modified forms.

[0083] The term “catalytic nucleic acid”, “catalytic DNA”, “deoxyribozyme”, “DNAenzyme” or “DNAzyme” as used herein refers to a nucleic acid molecule or oligonucleotide sequence that can catalyze or initiate a reaction, optionally in response to specifically recognizing and binding to a target analyte. DNAzymes may be single-stranded DNA, and may include RNA, modified nucleotides and / or nucleotide derivatives.

[0084] The term "aptamer" as used herein refers to a short, chemically synthesized nucleicacid molecule or oligonucleotide sequence which can be generated by in vitro selection to fold into specific three-dimensional (3D) structures that bind to a specific analyte with dissociation constants, for example, in the pico- to nano-molar range. Aptamers may be single-stranded DNA, and may include RNA, modified nucleotides and / or nucleotide derivatives. Aptamers may also be naturally occurring RNA aptamers termed “riboswitches”.

[0085] The term “hybridizes”, “hybridized” or “hybridization” as used herein refers to thesequence specific non-covalent binding interaction with a complementary, or partially complementary, nucleic acid sequence. When, for example, the 5ʹ-end region of an aptamer hybridizes to the 3ʹ-end region, it can form a duplex DNA element.

[0086] The term “biosensor” as used herein refers to a device that incorporates a biologicalentity as a molecular recognition element and is capable of producing a measurable signal uponbinding of a target analyte to the molecular recognition element. The biosensor can also be part of a larger device.

[0087] The term “functionalizing” or “functionalized on” as used herein refers to variouscommon approaches for functionalizing a material, which can be classified as mechanical, physical, chemical and biological. Any suitable form of coupling may be utilized (e.g. coating, binding, etc.). The functionalized material, for example, an aptamer or a blocking species, is also immobilized.

[0088] The term “room temperature” as used herein refers to a temperature in the range ofabout 20 °C and about 25 °C.

[0089] The term “working electrode” as used herein refers to an electrode in anelectrochemical system on which the reaction of interest is occurring. The working electrode can be used in conjunction with a counter electrode in a two-electrode system, and further a reference electrode in a three-electrode system. The counter electrode, also called the auxiliary electrode, is an electrode used in, for example, a three-electrode electrochemical cell for voltammetric analysis or other reactions in which an electric current is expected to flow. The counter electrode can also be part of a two-electrode system. The counter electrode is distinct from the reference electrode, which establishes the electrical potential against which other potentials can be measured, and the working electrode, at which the cell reaction takes place. As such, a reference electrode has a stable and well-known electrode potential. Depending on whether the reaction on the electrode is a reduction or an oxidation, the working electrode is called cathodic or anodic, respectively. Working electrodes can, for example, comprise materials ranging from inert metals such as gold, silver or platinum, to inert carbon such as glassy carbon, boron doped diamond or pyrolytic carbon, and mercury drop and film electrodes.

[0090] The term “highly-crosslinked” as used herein refers to a hydrogel polymer whichis both i) non-dissolvable; and ii) porous enough to swell by absorbing interstitial fluid. A highly-crosslinked hydrogel polymer may also be referred to as a first hydrogel polymer.

[0091] The term “low crosslinked” as used herein refers to a hydrogel polymer whichis both i) non-dissolvable; and ii) adhesive enough to enable a microneedle array to attach to an electrochemical chip. A low crosslinked hydrogel polymer may also be referred to as a second hydrogel polymer.

[0092] In understanding the scope of the present disclosure, the term “comprising” and itsderivatives, as used herein, are intended to be open ended terms that specify the presence of the stated features, elements, components, groups, integers, and / or steps, but do not exclude the presence of other unstated features, elements, components, groups, integers and / or steps. The foregoing also applies to words having similar meanings such as the terms, “including”, “having” and their derivatives. The term “consisting” and its derivatives, as used herein, are intended to be closed terms that specify the presence of the stated features, elements, components, groups, integers, and / or steps, but exclude the presence of other unstated features, elements, components, groups, integers and / or steps. The term “consisting essentially of”, as used herein, is intended to specify the presence of the stated features, elements, components, groups, integers, and / or steps as well as those that do not materially affect the basic and novel characteristic(s) of features, elements, components, groups, integers, and / or steps.

[0093] Terms of degree such as “substantially”, “about” and “approximately” as usedherein mean a reasonable amount of deviation of the modified term such that the end result is not significantly changed. These terms of degree should be construed as including a deviation of at least ^5% of the modified term if this deviation would not negate the meaning of the word it modifies. In addition, all ranges given herein include the end of the ranges and also any intermediate range points, whether explicitly stated or not.

[0094] As used in this disclosure, the singular forms “a”, “an” and “the” include pluralreferences unless the content clearly dictates otherwise.

[0095] In embodiments comprising an “additional” or “second” component, the secondcomponent as used herein is chemically different from the other components or first component. A “third” component is different from the other, first, and second components, and further enumerated or “additional” components are similarly different.

[0096] The term “and / or” as used herein means that the listed items are present, or used,individually or in combination. In effect, this term means that “at least one of” or “one or more” of the listed items is used or present.

[0097] The abbreviation, “e.g.” is derived from the Latin exempli gratia and is usedherein to indicate a non-limiting example. Thus, the abbreviation “e.g.” is synonymous with the term “for example.” The word “or” is intended to include “and” unless the context clearly indicates otherwise.

[0098] It will be understood that any component defined herein as being included maybe explicitly excluded by way of proviso or negative limitation, such as any specific compounds or method steps, whether implicitly or explicitly defined herein.

[0099] Although methods and materials similar or equivalent to those described hereincan be used in the practice or testing of this disclosure, suitable methods and materials are described below. II. Compositions and Methods of the Disclosure

[0100] The present disclosure provides a wearable electrochemical biosensor fordetecting a target analyte, such as an aptamer-based biosensor, that integrates hydrogel microneedle (HMN) arrays, such as methacrylated hyaluronic acid (MeHA) HMN arrays, with an electrochemical aptamer-based biosensor for real-time, continuous, and multiplexed monitoring of biomarkers, such as glucose and L-lactate, which can predict blood levels of these analytes.

[0101] The use of the presently disclosed HMNs addresses challenges associated withsolid MNs[6,13,25,26], offering a biocompatible, flexible yet mechanically robust medium for ISF extraction and sensing. Compared to prior studies demonstrating measurements lasting only a few hours[1,11], the wearable aptamer-based biosensor tracks biomarkers for three days in live rats. In exemplary embodiments of the disclosure, experiments were conducted to assess the performance of the wearable aptamer-based biosensor in transdermal glucose and lactate sensing by conducting experiments on both healthy mice injected with glucose or lactate, as well as type I diabetic rats. The results from both animal models indicate that the wearable aptamer-based biosensor is capable of monitoring the increasing and decreasing concentrations of glucose and lactate in vivo and in real-time.

[0102] Accordingly, herein provided is an electrochemical biosensor for detecting atleast one target analyte in interstitial fluid of a subject, comprising: a) a microneedle array comprising a highly-crosslinked hydrogel polymer;b) at least one electrochemical chip comprising at least one transducer surfaceconfigured for generating a signal in the presence of the at least one target analyte; c) a low crosslinked hydrogel polymer,wherein the low crosslinked hydrogel polymer is configured to couple the microneedle array with the at least one transducer surface of the at least one electrochemical chip.

[0103] With reference to FIG. 15A, the various elements of an exampleelectrochemical sensor for detecting at least one target analyte is shown. The electrochemical sensor 100 includes a microneedle array 105, a low crosslinked hydrogel polymer 110, and an electrochemical chip 115. The microneedle array 105 can be made from a highly-crosslinked hydrogel polymer. The highly-crosslinked hydrogel polymer being a hydrogel polymer which is both i) non-dissolvable; and ii) porous enough to swell by absorbing interstitial fluid. The low crosslinked hydrogel polymer 110 can be made from a low crosslinked hydrogel polymer which is both i) non-dissolvable; and ii) adhesive enough to enable the microneedle array 105 to attach to the electrochemical chip 115. The electrochemical chip 115 can have a transducer surface 120. The transducer surface 120 of the electrochemical chip can comprise one or more working electrodes. The low crosslinked hydrogel polymer 110 can be applied to the transducer surface 120, as shown in FIG. 15B. The low crosslinked hydrogel polymer 110 can allow for the coupling of the microneedle array 105 to the electromechanical chip 115. An exemplary assembled electrochemical biosensor is shown in FIG.15C.

[0104] In some embodiments, the highly crosslinked hydrogel polymer has acrosslinking density that is at least 10%, 20%, 30%, 40%, 50%, or 60% greater than that of the low-crosslinked hydrogel polymer. In some embodiments, the highly-crosslinked and the low crosslinked hydrogel polymer, each independently, comprise gelatin, hyaluronic acid, alginate, chitosan, collagen, or combinations thereof. In some embodiments, the highly-crosslinked and the low crosslinked hydrogel polymer, each independently, comprise methacrylated gelatin, methacrylated hyaluronic acid, methacrylatedalginate, methacrylated chitosan, methacrylated collagen, dopamine-functionalized hyaluronic acid or combinations thereof. In some embodiments, the highly-crosslinked and low crosslinked hydrogel polymers, each independently, comprise methacrylated hyaluronic acid. The highly-crosslinked hydrogel polymer can be any hydrogel polymer that is both i) non-dissolvable; and ii) sufficiently porous to allow swelling with interstitial fluid. The low crosslinked hydrogel polymer can be any hydrogel polymer that is both: i) non-dissolvable and ii) adhesive enough to enable a microneedle array to attach securely to the electrochemical chip.

[0105] In some embodiments, the highly-crosslinked hydrogel polymer comprises aporosity of about 25% to about 50%, about 30% to about 45%, about 35% to about 40% orabout 37%. In some embodiments, the highly-crosslinked hydrogel polymer comprises about 40 mg / mL to about 60 mg / mL, about 45 mg / mL to about 55 mg / mL, or about 50 mg / mL polymer. In some embodiments, the low crosslinked hydrogel polymer comprises a porosity of about 50% to about 75%. In some embodiments, the low crosslinked hydrogel polymer comprises a porosity of about 50% to about 75%, about 55% to about 70%, about 60% to about 67% or about 65%. In some embodiments, the second hydrogel polymer comprises about 25 mg / mL to about 40 mg / mL, about 30 mg / mL to about 35 mg / mL, or about 33 mg / mL polymer.

[0106] In some embodiments, the highly-crosslinked and the low crosslinked hydrogelpolymer, each independently, further comprise a crosslinking agent. The highly-crosslinked and low crosslinked hydrogel polymer may be crosslinked by a change in a physical and / or chemical property of the polymer, such as a change in pH. In some embodiments, the crosslinking agent comprises N,Nʹ-methylenebisacrylamide (MBA), PEG–maleimide, HA– thiol, enzymes, Carbodiimide-based (EDC / NHS), or combinations thereof.

[0107] In some embodiments, the highly-crosslinked and low crosslinked hydrogelpolymer, each independently, further comprise a photo initiator. In some embodiments, the photo initiator is Irgacure 2959 (2-Hydroxy-4’-(2-hydroxyethcxy)-2-methylpropiophenone). In some embodiments, the highly-crosslinked and low crosslinked hydrogel polymer comprise the same polymer. In some embodiments, the highly-crosslinked and low crosslinked hydrogel polymer comprise a different polymer.

[0108] The biosensor described herein may comprise one or more probesfunctionalized on the at least one transducer surface. In some embodiments, the electrochemical biosensor further comprises one or more probes functionalized on the at least one transducer surface that bind to the at least one target analyte, the one or more probes configured to generate a measurable signal in the presence of the at least one target analyte.

[0109] For targets like glucose, it is possible to directly detect their oxidation orreduction on electrode surfaces. These electrodes can be used on their own to detect the target or they can be combined with a separate transducer to do differential measurement. The other transducer can have antibodies or aptamers on its surface and target binding would reduce the access of the target (e.g. glucose) to the electrode surface, which would reduce the electrochemical signal. Comparison of the signals from the two transducers would allow a differential measurement system for a wide range of analytes. Accordingly, in someembodiments, the transducer surface is configured for reacting with the at least one target analyte to generate a measurable signal in the presence of the at least one target.

[0110] In some embodiments, the electrochemical biosensor further comprises (i) afirst probe that binds to a first target analyte functionalized on a first transducer surface and (ii) the first probe and a second probe that binds to a second target analyte functionalized on a second transducer surface, each transducer surface configured to generate measurable signals, wherein a difference between the measurable signal on the first transducer surface and the measurable signal on the second transducer surface is indicative of the presence of the second target analyte.

[0111] The electrochemical biosensor described herein uses at least oneelectrochemical chip comprising at least one transducer surface. In some embodiments, the at least one electrochemical chip comprises one or more working electrodes. In some embodiments, the one or more working electrodes comprises the transducer surface of the at least one electrochemical chip. In some embodiments, the at least one electrochemical chip further comprises a counter electrode. In some embodiments, the at least one electrochemical chip further comprises a reference electrode and / or the counter electrode is a reference electrode. In some embodiments, the transducer surface comprises a conductive material, semi- conductive material, or a combination thereof. In some embodiments, the transducer surface comprises metals, metal alloys, metal oxides, superconductors, semi-conductors, carbon-based materials, conductive polymers, or combinations thereof. In some embodiments, the transducer surface comprises metals, metal alloys, carbon-based materials, or combinations thereof. Examples include, but are not limited to, gold, platinum, palladium, carbon-based materials such as glassy carbon, graphite, graphene, or carbon nanotubes, nickel oxide, bismuth oxide, indium tin oxide, and titanium dioxide. In some embodiments, the transducer surface comprises gold. In some embodiments, the transducer surface further comprises a surface blocker agent. In some embodiments, the surface blocker agent comprises mercaptohexanol.

[0112] In some embodiments, the one or more probes are functionalized on thetransducer surface via chemical bonding, an intermediate linker, physical adsorption, embedded in inks, embedded in polymers, or a combination thereof. In some embodiments, the one or more probes are directly linked to the transducer surface. In some embodiments, the linker is a bifunctional linker comprising complementary reactive functional groups. Examplesof reactive functional groups include, but are not limited to, thiol, amine, epoxy, carboxylic acid, azide, and alkyne. In some embodiments, the linker comprises a thiol. In some embodiments, the one or more probes is functionalized on the transducer surface via chemical bonding. In some embodiments, the chemical bonding occurs via thiol and / or gold chemistry. In some embodiments, the one or more probes are indirectly linked to the working electrode. In some embodiments, the linker is a biotin, streptavidin, cystamine, or glutaraldehyde. In some embodiments, the one or more probes are linked to the electrode via biotin-streptavidin interactions.

[0113] In some embodiments, the electrochemical biosensor comprises one or moreprobes functionalized on a transducer surface. In some embodiments, the one or more probes is a nucleic acid probe, an antibody probe, or a combination thereof. In some embodiments, the one or more nucleic acid probes comprises a single-stranded DNA, single-stranded RNA, double-stranded DNA, double stranded RNA, DNAzyme and / or aptamer. In some embodiments, the nucleic acid probes comprise an aptamer. In some embodiments, the one of more nucleic acid probes comprises a detection moiety. In some embodiments, the one or more antibody probes comprises an anti-insulin antibody, an anti-troponin antibody, an anti-bovine serum antibody, or a combination thereof. In some embodiments, the one or more antibody probes comprises a detection moiety. In some embodiments, the detection moiety comprises an electroactive species, a redox active species, or combinations thereof. In some embodiments, the detection moiety comprises a redox active species. Examples of redox species include, but are not limited to, ruthenium haxaamine chloride, 3,7-Bis-[(2-Ammoniumethyl) (methyl)amino]phenothiazin-5-ium trifluoroacetate; 3,7-Bis-(piperazin-4-ium-1- yl)phenothiazin-5-ium trifluoroacetate; 3,7-Bis-[(2-ammoniumethyl)(methyl)amino] phenothiazin-5-ium chloride; and 3,7-Bis-(piperazin-4-ium-1-yl)phenothiazin-5-ium chloride, methylene blue, methylene blue succinimide, methylene blue maleimide, Atto MB2 maleimide (Sigma Aldrich) and other methylene blue derivatives, ferrocene and Fe+2 and / or Fe+3 ions. In some embodiments, the redox active species comprises methylene blue.

[0114] The biosensor described herein can be used for detecting a signal in the presenceof the at least one target analyte. In some embodiments, the signal comprises current, potential or impedance. In some embodiments, the target analyte is a biomolecule. In some embodiments, the target analyte comprises a small molecule, peptide or protein, nucleic acid,monosaccharide or oligosaccharide, or any combination thereof. In some embodiments, the target analyte is glucose and / or lactate.

[0115] The biosensor described herein comprises a microneedle array. In someembodiments, the microneedles are arranged in a wearable patch configured for transdermal sensing. In some embodiments, the biosensor is for use in health monitoring, screening and / or diagnostics. In some embodiments, the biosensor detects hypoglycemia, hyperglycemia and / or hyperlactatemia in a subject. In some embodiments, the biosensor is for use in diabetes monitoring and management.

[0116] Also provided is a transdermal patch comprising the biosensor disclosed herein.In some embodiments, the transdermal patch is for use in health monitoring, screening and / or diagnostics. In some embodiments, the transdermal patch detects hypoglycemia, hyperglycemia and / or hyperlactatemia in a subject. In some embodiments, the biosensor is for use in diabetes monitoring and management.

[0117] Also provided is a method for transdermal biosensing of at least one targetanalyte in interstitial fluid of a subject, the method comprising: a) applying a transdermal patch comprising the biosensor disclosed herein onto a subject;and b) detecting the signal from the at least one electrochemical chip of the biosensor, wherebythe at least one transducer surface generates an electrochemical signal in the presence of the at least one target analyte, wherein the electrochemical signal is indicative of the concentration of the at least one target analyte in the subject.

[0118] In some embodiments, the method further comprises swelling the transdermalpatch after step a). In some embodiments, swelling occurs for about 5 minutes to about 15 minutes, about 7 minutes to about 12 minutes, or about 10 minutes. In some embodiments, the signal is monitored continuously over time. In some embodiments, the signal is measured by amperometry, voltammetry, photoelectrochemistry, electrochemiluminescence, potentiometry or impedance. In some embodiments, the signal comprises current, potential or impedance. In some embodiments, the signal comprises current. In some embodiments, the signal is measuredusing square wave voltammetry. In some embodiments, the signal is used to calculate a kinetic differential measurement (KDM).

[0119] Also provided is a method for fabricating the wearable electrochemicalbiosensor disclosed herein, the method comprising: a) providing a microneedle array comprising a highly-crosslinked hydrogel polymer, themicroneedle array prepared by drying the highly-crosslinked hydrogel polymer in a mold; b) providing at least one electrochemical chip comprising at least one transducer surface;and c) coupling the microneedle array with the at least one transducer surface of theelectrochemical chip using a low crosslinked hydrogen polymer.

[0120] In some embodiments, crosslinking the polymer in step a) comprises adjustingpH before drying in the mold or exposing to ultraviolet (UV) light after drying in the mold. In some embodiments, microneedle array fabrication includes adjusting the pH of the highly- crosslinked hydrogel polymer to induce crosslinking and then drying in a mold. In some embodiments, fabrication of microneedle array comprises drying the polymer in the mold followed by demolding and exposure to UV light to induce crosslinking. In some embodiments, the highly-crosslinked hydrogel polymer is dried in the mold for about 6 hours to about 16 hours, about 8 to about 12 hours, or about 10 hours. In some embodiments, the highly- crosslinked hydrogel polymer is dried in the mold overnight. In some embodiments, the highly- crosslinked hydrogel polymer is exposed to UV light for about 20 minutes to about 60 minutes, about 30 minutes to about 50 minutes, or about 40 minutes.

[0121] In some embodiments, the electrochemical chip is functionalized with a surfaceblocker agent prior to step b). In some embodiments, the transducer surface is functionalized with one or more probes prior to being functionalized with the surface blocker agent. In some embodiments, the surface blocker agent comprises mercaptohexanol. In some embodiments, the method further comprises further crosslinking the first and second hydrogel polymer after coupling the microneedle array with the transducer surface of the electrochemical chip. In some embodiments, further crosslinking comprises exposure to UV light. In some embodiments, the first and second hydrogel polymer are exposed to UV light for about 5 minutes to about 15minutes, about 7 minutes to about 12 minutes, or about 10 minutes. In some embodiments, the method further comprises calibrating the biosensor by measuring a signal upon increasing and decreasing target analyte concentrations. In some embodiments, a KDM is calculated to generate a calibration curve.

[0122] In some embodiments, the method further comprises testing for faultybiosensors. In some embodiments, testing for faulty biosensors comprises measuring the electrochemical impedance change observed after coupling the microneedle array with the transducer surface of the electrochemical chip with the electrochemical chip functionalized with one or more biorecognition elements i.e. probes. In some embodiments, a |ΔZ / Z|>500 were determined as faulty biosensors. In some embodiments, currents of -30 nA at 80 Hz for measuring glucose and -100 nA at 200 Hz for measuring lactate were determined as cut-off points for in vivo tests. In some embodiments, a root mean square (RMS) of noise value of 0.01 and signal-to-noise ratio of 3 for signal-on conditions were also determined as cut-off points for in vivo tests. EXAMPLES

[0123] There is a need for MN-based biosensors that can monitor a wide array ofclinically-important target analytes, ranging from small molecules to large proteins. MN-based biosensing can analyze target molecules that undergo enzymatic conversion[15-17]or have redox activity.

[0018] A recent study developed an electrochemical sensor by integrating MeHA hydrogel microneedle (HMN) patches with conductive thermoplastic electrodes modified with glucose and lactate oxidase for real time monitoring of glucose and lactate, respectively.

[0017] This work demonstrated the ability of HMN in ISF extraction and delivery to an integrated electrode system. Recently, there is also increasing efforts invested toward coupling MN technology with aptamer-based electrochemical sensing that is independent of molecular reactivity or enzyme coupling of the target analyte.[11,19,21]In aptamer-based electrochemical biosensors, aptamer probes undergo a conformational change upon target binding, which alters the electron transfer efficiency between a redox reporter coupled to the aptamer and the working electrode, resulting in a detectable electrochemical signal change.[22,24]To date, aptamer-based electrochemical biosensors for ISF monitoring have only been applied to solid MNs, inheriting the challenges of solid-state systems with the aptamers functionalized on the surface of the solid MNs, such as the aptamer-assisted microneedle-based device described in WO2021 / 119546. Due to thesedifficulties, continuous, in vivo monitoring by aptamer-based MNs has been restricted to just a few hours and has been demonstrated solely in the context of therapeutic drug monitoring.

[0011]

[0124] The following non-limiting examples are illustrative of the present disclosure:

[0125] Results and Discussion

[0126] Wearable Aptamer-Based Biosensor Fabrication. The integrated wearableaptamer-based biosensor combines HMN patches for ISF extraction with electrodes functionalized with redox aptamers for target analysis. Upon skin insertion, the HMN patch undergoes swelling, facilitating the diffusion of target analytes to aptamer-functionalized electrodes (FIG. 1a). In this configuration, aptamers are not directly placed in ISF in order to eliminate probe damage encountered in in vivo systems; instead, HMN patches are attached tothree-electrode electrochemical chips via hybrid integration – i.e. the electrochemical chip isintegrated with other off-chip materials. For example, the attachment of HMNs to the electrodes may be enabled using a thin layer of highly swellable hydrogel, acting as a flexible and biocompatible glue.

[0127] A multi-step process was employed to fabricate the wearable aptamer-basedbiosensor device (FIG. 1b). To fabricate HMN patches, MeHA was synthesized based on previously published protocols.[6]Briefly, hyaluronic acid (HA) was chemically modified with methacrylated anhydride overnight. The resulting MeHA was then purified through precipitation and dialysis before being lyophilized.[6]The degree of methacrylation was confirmed using1H- NMR (FIG. 2). The synthesized MeHA polymer was utilized to prepare highly crosslinked MeHA (HC-MeHA) and low crosslinked MeHA (LC-MeHA) solutions, comprising of 50 mg mL-1and 33 mg mL-1of MeHA, respectively, and crosslinking agents. The HC-MeHA solution was cast into the patterned molds for fabricating HMN patches. Following an overnight drying process, the patches were demolded and crosslinked under UV irradiation for 40 minutes. Next, 100 ^^L of LC-MeHA was applied on the functionalized electrodes to create a thin adhesion layer that covers the working, counter, and reference electrodes. The LC-MeHA functions as a glue to attach the HMN patches on the solid electrodes.[27,28]Subsequently, the crosslinked HMN patches were trimmed, placed on the LC-MeHA, and air dried for improving adhesion. At the final step, the integrated device was further crosslinked under UV for 10 minutes. The utilization of LC- MeHA allows strong attachment of the electrodes to the HMN patches (FIG.1c), thus enhancing the integrated device robustness during the following animal experiments.

[0128] To convert the integrated patch / electrode device into the wearable aptamer-basedbiosensor a well-characterized aptamer – lactate

[0029] (SEQ ID NO:1; 5'- / 5ThioMC6-D / GAC GAC GAG TAG CGC GTA TGA ATG CTT TTC TAT GGA G / iAmMC6 / TC GTC-3'; where / 5ThioMC6-D / represents a thiol modification at the 5’ guanosine, for example, through a six- carbon linker, and / iAmMC6 / represents a six-carbon linker connected to thymidine at position 38 for conjugation to methylene blue) or glucose

[0030] (SEQ ID NO: 2; 5'- / 5thioMC6-D / ACG ACC GTG TGT GTA / iAmMC6 / TT TCT ATA CAG TGT CCA TTG TCG T-3'; where / 5ThioMC6- D / represents a thiol modification at the 5’ adenosine, for example, through a six-carbon linker, and / iAmMC6 / represents a six-carbon linker connected to thymidine at position 16 for conjugation to methylene blue) – was anchored onto gold working electrodes through gold-thiol chemistry (FIG. 1c) and its attachment was verified using electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV) (FIG. 3). The aptamer was tagged with methylene blue as a redox reporter for electrochemical signal transduction. Upon target binding, the redox aptamer undergoes a conformational change which brings the methylene blue close to the gold surface, resulting in an increased electron transfer rate compared to the case where the target is absent.[14,21]Square wave voltammetry (SWV) was used to measure the signal generated from the wearable aptamer-based biosensor. The difference in the charge transfer kinetics between the bound and unbound states of the redox aptamer was utilized to calculate the kinetic differential measurement (KDM).

[0031] This method, previously developed for aptamer-based biosensors, effectively reduces sensor-to-sensor variability, corrects signal drift, and enhances the signal-to-noise ratio.[11,14]The signal-on and -off frequencies were determined using Lovric’s formalism

[0032] by calculating the electrochemical signal gain across frequencies ranging from 5 to 300 Hz (FIG. 4). Glucose and lactate sensors exhibited signal-on and signal-off responses at higher (80 Hz for glucose and 200 Hz for lactate aptamer) and lower frequencies (5 Hz for glucose and 15 Hz for lactate aptamer), respectively.

[0129] Physical Characterization of the Wearable Aptamer-Based Biosensor.Following the fabrication of HMN patches, their physical properties were first assessed. To ensure effective skin penetration, it is crucial that HMN patches have sharp and intact needles. The morphology of HMN patches was evaluated using scanning electron microscopy (SEM), showing the intact and sharp needle tips (FIG.5a). Next, the porosity of the HC-MeHA and LC- MeHA hydrogel was investigated by calculating the average pore size from the corresponding swelled hydrogels imaged using SEM (FIG.5b). A porosity of 37% and 65% was calculated forHC-MeHA and LC-MeHA, respectively. HC-MeHA, forms a dense polymer network within HMN, thus maintaining its structure after water absorption. In contrast, LC-MeHA, functioning as an adhesion layer between the HMN patch and the electrode, forms a less dense polymer network. The degree of crosslinking mainly depends on the polymer content, duration of UV curing, and the concentration of crosslinking agents.

[0033] To investigate the duration required for HMN patches to collect enough ISF, the swelling capability of HMNs was assessed using porcine skin. HMN patches were fabricated using HC-MeHA with an additional layer of LC-MeHA on their back (HC+LC-MeHA HMN) and were applied on a piece of porcine skin for different durations and their corresponding swelling ratio was measured. It was observed that the HMN patches reach their maximum swelling of 119.45 % within 6 minutes of porcine skin application (FIG. 5c), indicating that the HMN facilitates rapid and effective ISF sample collection under real skin environments.

[0130] The mechanical strength of HMN patches was next evaluated through acompression test, assessing their ability to penetrate the skin. HC+LC-MeHA HMN patches were examined to investigate the effect of adding LC-MeHA on the mechanical property of the HMN patches. The displacement of the needle tips following the application of a compressive force was measured and plotted (FIG. 5d). This data illustrates that both HC-MeHA HMN and HC+LC-MeHA HMN can withstand a force exceeding 0.3-0.4 N per needle, which is necessary for penetrating human skin, without experiencing significant needle breakage.

[0034] Importantly, the inclusion of the LC-MeHA layer does not impact the mechanical strength of the HMN patches.

[0131] After characterizing the physical properties of HMN patches, the HMN patcheswere integrated with the electrodes functionalized with the redox aptamer using LC-MeHA as the adhesion layer (FIG. 5e). To assess the attachment of HMN patches to the electrodes, a pulling experiment was performed. As shown in FIG. 5f, a HC-MeHA thin film, mimicking HMN patches, was attached to the functionalized electrodes through a layer of LC-MeHA. An inelastic tape was then attached and a lateral pulling force was applied, pulling the thin film away from the electrodes and stimulating the force applied on the HMN patches during body movement in practical applications. To study the role of LC-MeHA as an adhesion layer, 100 ^^L or 50 ^^L of LC-MeHA were used. FIG.5g shows the force applied versus displacement of HC- MeHA. Both samples showed a linear increase in displacement, then reached the peak value once the HC-MeHA-electrode adhesive joint failed. The HC-MeHA layer attached via 100 ^^L of LC-MeHA was detached after applying a force of 2.2 N, while HC-MeHA attached via 50 ^^L of LC- MeHA requires only 1.1 N for removal. These results demonstrate that LC-MeHA facilitates effective HMN-to-electrode attachment.

[0132] Wearable Aptamer-Based Biosensor Sensing and Optimization. Prior to theimplementation of the wearable aptamer-based biosensor, the electrochemical characterization of the device fabrication was performed using CV and electrochemical impedance spectroscopy at open circuit potential. The charge transfer resistance of the electrodes increased by 12366% and 3073% after biofunctionalization with the glucose and lactate aptamer probes, respectively. Additionally, the resistance further increased by 22% and 70% after the addition of mercaptohexanol as a surface-blocking agent for the glucose and lactate sensors, respectively. Finally, upon the attachment of HMN, the resistance increased by 23% and 25% as anticipated (FIG. 3). A CV measurement was also recorded after each fabrication step of the wearable aptamer-based biosensor. A decrease in electrochemical current was observed after every step of fabrication compared to the preceding step (FIG.3). This is due to the steric hindrance imposed by each layer added during fabrication, reducing the access of the electrolyte to the electrode surface.

[0035] The observed trend in the electrochemical current is in line with the charge transfer resistance results calculated from electrochemical impedance spectroscopy and confirms the successful addition of each functional layer to the device, further validating the successful fabrication of the wearable aptamer-based biosensor.

[0133] Next, the glucose and lactate aptamer-based biosensors were evaluated ex vivousing artificial ISF (aISF) spiked with different concentrations of the target analytes. The validation experiment was conducted iteratively, either starting with a blank aISF sample and subsequently increasing the concentration (upward calibration) or initiating with the highest concentration of the target analyte in aISF followed by dilution (downward calibration), (FIG. 6a and FIG.6b). The glucose aptamer-based biosensor was exposed to concentrations ranging from 0 mM to 50 mM, covering the normal physiological range,

[0036] as well as exceeding hypoglycemia

[0036] and hyperglycemia levels

[0036] (FIG.6a). A lactate range of 0 mM to 20 mM was chosen to include the physiological range, as well as hyperlactatemia (FIG. 6b). After a 20- minute incubation period, the electrochemical signals were measured. During upward calibration, the aptamer-based biosensors generate an electrochemical signal that monotonicallyincreases with increasing target analyte concentration (FIG. 6a and FIG. 6b), aligning with theaptamer sensor operation mechanism: increased target concentration leads to a largerelectrochemical current due to the more frequent interaction of the methylene blue with the electrode. The resulting KDM was then calculated to obtain the calibration curve for the glucose and lactate aptamer-based biosensors. A log-linear fit to the glucose calibration curve yields a limit-of-detection of 2.4 mM for detecting glucose that covers the hypoglycemia range (FIG. 7a). Also, a limit-of-detection of 1.04 mM was achieved by fitting a log-linear line to the lactate calibration curve (FIG. 7b). A different behavior was observed during downward calibration, resulting in a higher limit-of-detection of 7.44 and 4.43 mM for glucose and lactate, respectively (FIG.8a and FIG.8b).

[0032] [33,34]This divergence can be caused by the differences in dissociation rate constant (koff) and association rate constant (kon) for target-aptamer binding, while maintaining a constant koff / konratio (equilibrium rate constants (Kd)).

[0037] Previous experiments performed using surface plasmon resonance have demonstrated that aptamer-target dissociation is slower than aptamer-target association.

[0038] To investigate the differences between association and dissociation rates of the glucose and lactate aptamers used in the wearable aptamer-based biosensor, the glucose and lactate concentrations were increased from 2.5 mM and 1 mM (after a 20 minute stabilization period) to 10 mM and 5 mM and decreased their concentrations back to 2.5 mM and 1 mM, respectively, while measuring the electrochemical signals at every 6 minutes (FIG.8). These results confirmed previous findings showing slower dissociation compared to association in aptamer-based systems, explaining the differences in upward and downward slopes when measurements were performed at fixed time points. Upon the successful validation of the glucose and lactate aptamer-based biosensors using spiked aISF solutions, the fully integrated wearable aptamer-based biosensor was tested using ex vivo porcine skin. The wearable aptamer-based biosensor was applied on the porcine ear skin loaded with various concentrations of lactate or glucose (FIG. 5c and FIG. 5d, inset). Upon application, HMN patches underwent swelling, facilitating the transfer of the target analytes to the electrode surface. A 10-minute period was allocated for the HMN patches to reach maximum swelling, as determined in the swelling experiment (FIG. 4c), followed by a 20-minute target incubation. After 30 minutes, the electrochemical signal measurement was recorded using SWV, the averaged KDM values were calculated across a range of target analyte concentrations, and the calibration curves were obtained (FIG. 5c and FIG. 5d). Notably, an increase in electrochemical current was observed by increasing the target concentrations, which is in line with prior experiments and confirms the successful penetration of HMN patches and detection of the target analytes. An ex vivo kinetic experiment was conducted to determine the timeresolution of the wearable aptamer-based biosensor. The lactate aptamer-based biosensors were applied on porcine skin containing 5 mM or 20 mM of lactate, followed by electrochemical measurements after 5, 10, 15, 20, 30 and 40 minutes of incubation (FIG.9). Although the HMN reached maximum swelling within 6 minutes (FIG.5c), the KDM value became stable after a 20-minute incubation period, indicating that the sensor’s time resolution could be dependent on the kinetics of target diffusion and binding.

[0134] To further investigate the integrity of the developed platform, the glucoseaptamer-based biosensor was challenged by introducing a panel of common interfering agents found within ISF, including lactate, ascorbic acid, glycine, L-alanine, magnesium chloride, hydroxybutyrate, sucrose, uric acid, bovine serum albumin (BSA), and galactose, all spiked in aISF (FIG. 6e). The concentrations of all the compounds were 10 mM in aISF except for magnesium chloride, which was 5 mM. The concentration of glucose was within the physiological range, however, the concentration of the other interferents was significantly higher than the normal levels. Among all the tested interferents, including sugars and amino acids, none demonstrated interference with detecting glucose, showing exceptional specificity for the glucose aptamer-based biosensor. Similarly, the selectivity of the lactate aptamer-based biosensor was examined by introducing the same concentration of interfering compounds found in ISF. The calculated signal gains showed that despite the significantly high concentration of non-target molecules, the lactate aptamer-based biosensor exhibited excellent selectivity toward lactate (FIG.6f).

[0135] In vivo Validation of Wearable Aptamer-Based Biosensor Using a MouseModel. Having validated the performance of the wearable aptamer-based biosensor ex vivo, the integrated device performance was then assessed using animal models. Before the in vivo assessment, a quality control procedure, based on single-frequency electrochemical impedance measurement, was established to identify faulty sensors (FIG. 10). Based on this procedure, sensors demonstrating |ΔZ / Z|>500 after being inserted into skin, where ΔZ / Z represents the electrochemical impedance change observed after attaching the HMN for the wearable aptamer- based biosensor compared to functionalized sensors, were considered faulty and were excluded. Additionally, the currents of -30 nA at 80 Hz for the glucose device and -100 nA at 200 Hz for the lactate device were determined as cut-off points for in vivo tests. A root mean square (RMS) of noise value of 0.01 and signal-to-noise ratio of 3 for signal-on conditions were also determined as the cut-off for the in vivo experiments. Faults in sensors could be caused by incompleteattachment of the HMN patches to the electrodes or imperfect insertion of the HMN patches to the skin.

[0136] Wearable aptamer-based biosensor sensors were employed for continuousglucose and L-lactate measurements in healthy wild-type mice by applying the integrated device on their dorsal skin (FIG.11a). The penetration of HMN patches was first evaluated by applying the wearable aptamer-based biosensor for 5 minutes and collecting the skin samples after device removal. Successful penetration of the HMN patches was confirmed by observing a swelled HMN patch without needle rupture (FIG. 11b) and performing hematoxylin and eosin stain (H&E) staining on the collected skin (FIG.11c). H&E staining images showed a needle cavity with an approximate depth of 78 µm. To increase and decrease blood glucose levels, each mouse had an intraperitoneal injection of a 3 g kg-1dosage of D-glucose, followed by a 0.75 U kg-1injection of insulin after a 1-hour interval. The average results of glucose measurement for three different mice, each using two wearable aptamer-based biosensors are shown in FIG.11d-f. The ex vivo calibration curve (FIG. 6a) was used to determine the glucose concentration measured by the wearable aptamer-based biosensor. Simultaneously, mouse blood samples collected from the tail were used to measure blood glucose using a handheld glucometer. An average difference of 2.77, 1.75, and 1.68 mM was calculated between wearable aptamer-based biosensor measurements and the glucometer readings for the first, second, and third tested animals, respectively. This confirms the reliability of the wearable aptamer-based biosensor for detecting fluctuating glucose concentrations.

[0137] To elevate the L-lactate level, each mouse received an intraperitoneal injection ofa 0.6 g kg-1dosage of L-lactate and the wearable aptamer-based biosensor was employed to track the fluctuation in L-lactate levels. The ex vivo lactate calibration curve (FIG. 6b) was used to obtain L-lactate concentrations in vivo. Similar to glucose experiments, mouse blood samples were collected from the tail, and serum L-lactate were analyzed using a colorimetric kit. An average difference of 2.10, 2.09, and 3.03 mM were calculated between the wearable aptamer- based biosensor measurements, and the results obtained from the colorimetric kit for the first, second, and third mouse, respectively (FIG. 11g-i), affirming the capability of the wearable aptamer-based biosensor for the measurement of L-lactate levels.

[0138] In vivo Validation of Wearable Aptamer-Based Biosensor for MultiplexedMeasurement. To assess the potential applicability of the biosensing assay across various animalmodels, the wearable aptamer-based biosensor was further characterized and tested using a diabetic rat model. First, the effectiveness of the device for rat skin penetration as well as the skin recovery upon HMN application were studied. FIG. 12a shows the wearable aptamer-based biosensor before and after skin penetration, highlighting the presence of swelling and intact needles upon skin penetration. In addition, the attachment between HMN and the electrodes remains firmly in place. Subsequently, the needle’s trace on the skin of the rat after HMN penetration was monitored over time (FIG. 12b). Immediately upon the removal of the HMN patch, a noticeable imprint of the needles was observed, providing clear evidence of effective skin penetration. However, the needle imprint disappeared after 20 minutes, indicating skin recovery. The H&E staining of skin also demonstrates a needle cavity with a depth of 92.5 ^^m within the rat skin (FIG.12c). These results indicate that the wearable aptamer-based biosensor is capable of effectively penetrating rat skin, and upon removal, the skin exhibits a rapid recovery.

[0139] To show the applicability of the wearable aptamer-based biosensor sensor for invivo measurement in rats, the devices were applied on the rat dorsal skin and connected to a potentiostat as shown in FIG.12d. The electrochemical current response was measured by SWV after a one-hour incubation period. To assess long-term sensing stability, lactate aptamer-based biosensors were applied on the rat dorsal skin and left in place for 3 days. The lactate level was monitored daily using both the wearable aptamer-based biosensors, measuring the ISF lactate, and the colorimetric L-lactate assay kit, assessing the blood lactate. Results shown in FIG.12e indicate the lactate levels measured by the wearable aptamer-based biosensor and conventional blood lactate assay are well-aligned and the wearable aptamer-based biosensor devices produce stable and reliable electrochemical signals for a lifespan of at least 3 days. During the experiment, the wearable aptamer-based biosensor sensors were securely fixed on the rat body using Tegaderm and kinesiology tapes, and a protective cone was placed over the rat to prevent the rat from removing the electrodes. The wearable aptamer-based biosensor monitored the stress- induced lactate fluctuation on a diabetic rat. As shown in FIG.12e, the lactate level increased to hyperlactatemia level at day 1, since the rat experienced constant stress induced by the experimental set-up. The rat was then gradually acclimated to the experimental condition, resulting in a decrease in lactate level on day 2, eventually returning to the normal range by day 3. It should be noted that the experiment was stopped after three days due to rat health concerns. Previously reported aptamer-based MN biosensors, where the solid needle surface was functionalized with aptamer probes, could only monitor target concentrations for up to two hoursin vivo.[14,39]In these platforms, the sensor performance gradually deteriorated as the aptamer probes came into direct contact with the complex ISF medium.

[0014] The wearable aptamer-based biosensor has demonstrated improved performance for two possible reasons: 1) the redox aptamer probes are not in direct contact with ISF, and 2) the hydrogel network serves as a protective layer, decreasing the access of proteins and cells to the sensor surface, and thus reducing sensor fouling.

[0040] The reproducibility and stability of the wearable aptamer-based biosensor was further validated using multiple SWV scans. Upon device application, each electrode was scanned 100 times continuously and the KDM was calculated and plotted for each round of scans (FIG. 12f). The average KDM remains stable after 100 consecutive scans, and the variance between different replicates was minimal.

[0140] The rat model, with a larger surface skin area in comparison to the mouse model,enables the attachment of several wearable aptamer-based biosensor devices. Consequently, simultaneous, multiplexed measurements of glucose and L-lactate were conducted in diabetic rats by employing two devices per target analyte. Specifically, continuous monitoring of lactate along with glucose enables accurate evaluation of a metabolic by product of exercise that correlates with exercise intensity in individuals with diabetes and facilitates precise predictions of the glycemic response and the necessary insulin dosage.

[0041] The diabetic rats were fasted for 4 hours prior to the experiment and then given human recombinant insulin to reduce their blood glucose levels. Two lactate and two glucose wearable aptamer-based biosensors were applied on the shaved back of the rat and continuously monitored the level of glucose and lactate in the animals. The data was collected at 5 different time points (FIG.12i-e). The individual data points obtained from the two individual wearable aptamer-based biosensors were included in FIG. 13. In parallel, the data obtained from glucose and lactate devices were cross-validated with the handheld glucometer and colorimetric lactate kit, respectively. The decrease in glucose levels following insulin injection was monitored using the glucose wearable aptamer-based biosensor, with an average difference of 2.45, 1.82, 1.58 mM compared to blood measurements for the first, second, and third rat, respectively. In contrast, the lactate level remained consistent over time, and the measurements obtained from the wearable aptamer-based biosensor deviated by only 0.17, 0.49, 0.32 mM compared to lactate blood measurements for the first, second, and third rats, respectively. The wearable aptamer-based biosensor was compared and summarized with recently published enzymatic and aptamer based electrochemical sensor in Table 1. Theseresults indicate that the wearable aptamer-based biosensor is capable of tracking multiple targets continuously, adapting to decreasing and stable trends. Table 1. Comparison of recently published enzymatic and aptamer based electrochemical sensor with the wearable aptamer-based biosensor disclosed herein Ref Type of MN Sensing mechanism Target Comments Enzymatic Dexcom G6 Metallic needle electrochemical Glucose 13 mm long metallic needle is sensing required Abbott Enzymatic FreeStyle Metallic needle electrochemical Glucose 6 mm long metallic needle is Libre sensing required Enzymatic Devices co MeHA HMN electrochemical Glucos mprise a bulk MeHA

[17] e and hydroge ng lact l body, requiring 20 sensi ate minutes for sample collection. Poly (methyl Enzymatic Glucose, lactat Enzyme coated on the surface [1] methacrylate) electrochemical e and of solid MN might be damaged solid MN sensing alcoholduring skin insertion process.

[43] Gold-plated Aptamer based The in vivo study only electroche vancomycin, metallic MN mical validated the sensor sensing doxorubicin, andthrombin a single target. Gold-plated Aptamer based An invasive magnet

[0011] polymeric solid electrochemical Tobramycin implantation is required to affix MN sensing the sensor on the skin. Wearable Biocompatible and swellable aptamer- Aptamer based HMN is integrated to external based MeHA HMN electrochemical Glucose and electrodes modified wit biosensor ensing l h s actate aptamer, reducing the protentional damage to aptamer.

[00141] Wearable Antibody-Aptamer-Based Biosensor Fabrication. In addition todepositing aptamers on the electrode, a sensor that uses a combination of aptamers and antibodies for signal generation was developed. In this approach, the sensor has two chips or electrodes, one chip / electrode measures lactate concentration used lactate aptamers to establish a baseline measurement; whereas, another chip measured the lactate signal using an electrode modified with antibodies in addition to lactate aptamers. Capture of the target, in this case cardiac troponin I (cTnI) by the antibodies (anti cTnI antibodies) decreased the concentration of lactate at theelectrode vicinity due to steric hindrance, reducing the measured lactate signal (FIG. 14a). Bycontinuously calculating the difference in the signal measured from the two electrodes (differential signaling), we extrapolated the cTnI concentration as demonstrated by (FIG.14b). This approach was also evaluated in animals by using two chips, one that measured lactate only and another one that measured cardiac troponin by combining antibodies and aptamers on the electrode surface. (FIG.14a-d)

[0142] Experimental Section

[0143] Materials and Reagents. All oligonucleotides were purchased from Biosearchand Integrated DNA Technologies (IDT) and purified by standard 10% denaturing (8 M urea) polyacrylamide gel electrophoresis (dPAGE) before use. The water used in the experiments was purified with a Milli-Q Synthesis A10 water-purification system. Phosphate buffer solution (1.0M, pH 7.4), sodium dihydrogen phosphate, anhydrous calcium chloride, magnesium sulfate, sodium hydroxide, sodium L-lactate, (±)-sodium 3-hydroxybutyrate, sucrose, L-alanine, glycine, galactose, D-(+)- glucose, uric acid, ascorbic acid, magnesium chloride, potassium chloride, 6- mercapto-1-hexanol (MCH, 99%), tris(2-carboxyethyl)phosphine hydrochloride (TCEP), bovine serum albumin (BSA), potassium hexacyanoferrate(ii) trihydrate, and 4-(2- hydroxyethyl)-1- piperazineethanesulfonic acid (HEPES), Irgacure 2959 (2-Hydroxy-4’-(2-hydroxyethcxy)-2- methylpropiophenone, photo initiator, PI), acrylamide (MBA), deuterium oxide, methacrylate anhydride (MA), human recombinant insulin were purchased from Sigma-Aldrich. Sulfuric acid (98%) and 2-propanol (99.5%) were purchased from Caledon Laboratories. Hydrochloric acid (HCl; 37% w / w) was purchased from LabChem. The pharma-grade sodium hyaluronic acid (HA, MW 300 kDa) was purchased from Bloomage Co., Ltd (China). Colorimetric L-lactate assay kits (ab65331) were purchased from Abcam. Insulin aspart solution for injection was purchased from NovoRapid. The cloth restrainer for the mice was purchased from Lomir Biomedical. Microhematocrit heparinized capillary tubes were purchased from Fisher brand.

[0144] Methacrylated Hyaluronic Acid (MeHA) Synthesis. MeHA was synthesizedfollowing on the modified protocol established before.[6]Briefly, 2 mL methacrylate anhydride (MA) was added to 100 mL of 20 mg mL-1HA solution (Bloomage Co., Ltd, China) under 4 °C. Meanwhile, 4.2 mL of 5 N NaOH was added gradually to adjust the solution to about pH 8-9. After overnight reaction, MeHA was precipitated by acetone and washed three times with ethanol. The precipitated MeHA was redissolved in Millipore water and then dialyzed for 3 daysto eliminate impurities. The purified MeHA was freeze-dried for 3 days then stored in a desiccator for future use. Eventually, 5 mg of resulting MeHA was dissolved in 1 mL of deuterium oxide (D2O) and then tested using 300 MHz1H NMR with 10 ms time scale. The degree of methacrylate modification was determined by integration of methacrylate proton signals at 6.1, and 5.7 ppm.

[0145] Hydrogel Microneedle (HMN) Fabrication. HMN was prepared using highcrosslink degree methacrylated hyaluronic acid (HC-MeHA) solution.50 mg of MeHA, 2 mg of PI, and 2 mg of MBA were dissolved in 1 ml DNase free water. A 0.3 mL HC-MeHA solution was then casted on a polydimethylsiloxane (PDMS) mold, degassed for 1.5 minutes, then demolded after overnight drying process. The resulting high crosslinking degree HMN (HC- HMN) patches were crosslinked under Ultraviolet (UV) in 365 nm wavelength for 40 minutes.

[0146] Wearable Aptamer-Based Biosensor Fabrication. The wearable aptamer-based biosensors were prepared in a series of steps. Initially, the chips were incubated with isopropanol and then rinsed with water. Subsequently, electrochemical cleaning was executed through cyclic voltammetry in 0.1 M sulfuric acid, involving a potential range from 0 to 1.5 V, a scan rate of 0.1 V s-1, and a total of 10 cycles. After this thorough cleaning process, the working electrode was carefully functionalized with glucose or lactate aptamers, serving as the biorecognition elements. A 2 µM thiol-terminated glucose / lactate aptamers solution was reduced with 200 µM TCEP, and this reduction process was completed for 2 hours in a dark environment at room temperature. The reduced aptamer solution was then deposited onto the working electrodes and left to incubate overnight, also at room temperature. Following the aptamer deposition, a 100 mM MCH solution was drop-deposited for backfilling the surface of the working electrode for 10 minutes in the dark at room temperature. Subsequently, the chips were washed with a buffer solution and then dried with a gentle flow of nitrogen gas. The next step involved the attachment of HMN arrays onto the functionalized chips. For this purpose, a low crosslinking degree MeHA solution (LC-MeHA solution) comprising 33.3 mg mL-1of MeHA, 0.67 mg mL-1mg of PI, and 0.67 mg mL-1of MBA was used to act as a glue to attach the HMN arrays onto the functionalized chips.100 µL LC-MeHA solution covering the counter, reference, and working electrodes was dried at room temperature for 10 minutes. Subsequently, the crosslinked HC HMN array was trimmed and affixed onto electrodes through the partially dried LC-MeHA. The entire device was further air-dried for 1hr, followed by 10 minutes UV crosslinking.

[0147] Electrochemical Characterizations of the Wearable Aptamer-Basedbiosensors. Different electrochemical techniques including square wave voltammetry (SWV), cyclic voltammetry (CV), and electrochemical impedance spectroscopy were carried out using PalmSense and EmStat4S potentiostats. All the SWV measurements were performed at room temperature in artificial ISF (2.5 mM CaCl2, 10 mM HEPES, 3.5 mM KCl, 0.7 mM MgSO4, 123 mM NaCl, 1.5 mM NaH2PO4, 7.4 mM saccharose, pH 7.5) or SELEX buffer (500 mM NaCl, 10 mM MgCl2, 10mM KCl, 50 mM HEPES, pH 7.5) over a potential range of 0 to −0.6 V at frequencies of 5 and 80 Hz for glucose sensors, and 15 and 200 Hz for lactate sensors. The stepwise development of the wearable aptamer-based biosensor was thoroughly characterized. This included the immobilization of a thiol-terminated aptamer probe and an MCH surface blocker agent on the working electrode and also HMN attachment onto the functionalized chips. This characterization was done using a CV scan from -0.2 V to 0.5 V at a scan rate of 0.1 V s-1and EIS measurements from 0.1 to 10000 Hz in an electrolyte solution containing 2 mM [Fe(CN)6]3- / 4-in 0.05 M KCl and 0.1 M PBS solution. Kinetic Differential Measurement (KDM) values were determined by calculating the difference between the normalized peak currents obtained at a signal-on and a signal-off frequency, then dividing by the average of the signal-on and signal-off signals. For determining on and off frequencies, the electrochemical signal gain was calculated at different frequencies, in a potential range of 0 to -0.6 V with a scan rate of 0.1 V s-1(FIG.3).

[0148] Scanning Electron Microscopy (SEM) Imaging. To investigate the porosity ofHC and LC MeHA hydrogel, crosslinked HC-MeHA and LC-MeHA thin films were prepared and swelled in water for 5 minutes. The swelled films were then frozen in liquid nitrogen and freeze-dried for 48 hours to fix the pore size. The dried MeHA thin films were then coated with a 2 nm thick layer of gold and imaged using a Hitachi SU5000 Field Emission Scanning Electron Microscope (FESEM). Similarly, the HC+LC MeHA HMN array was gold-coated and subsequently imaged using the Hitachi SU5000 FESEM to examine its morphology.

[0149] Swelling Studies. HMN patches for the swelling experiment were preparedfollowing the protocol mentioned above. However, an additional layer of 100 ^^L of LC-MeHA solution was added on the half-dried HC-MeHA HMN to mimic the actual morphology of the HMN integrated electrodes. The resulting HC+LC MeHA HMNs were trimmed and weighted to determine the dry mass (W0) of HMN. Next, HMNs were applied on the porcine ear skin which had been pre-equilibrated in 1X PBS solution overnight. The HMN was affixed on the porcineskin using Tegaderm tape (3M, USA) and allowed to swell for 1, 2, 4, 6, 8, and 10 minutes. The wet mass (Wt) of swelled HMNs were measured and recorded. The swelling ratio of HMNs was calculated using the formula:

[0150] Swelling ratio = (^^^^−^^0^^0 ) × 100%

[0151] MechanicalPatches. The mechanical property of theHC+LC HMNs was tested using Instron 5548 micro tester loaded with a 500 N compression cell. The HMN patch with its tips facing upwards was placed on the lower platen molding a 1.5 mm distance from the upper platen. During the experiment, the upper platen descended at a constant speed of 0.5 mm min-1, applying a vertical compression force to the tip of the HMN until it reached 70 N. The compression force applied on the HMN (N) and displacement of the HMN tips (mm) was recorded and plotted to obtain the mechanical strength of the HMN array.

[0152] Pulling Experiment. The pulling experiments were conducted to characterize theattachment strength between the HMN patch and sensing chip. In this study a HC-MeHA thin film to mimic the HC-MeHA HMN array and was attached to the functionalized electrodes through a thin layer of LC-MeHA. The MeHA thin films attached to the electrodes with 100 ^^L of LC-MeHA were designed as “100 ^^L LC-MeHA”, whereas those with 50 ^^L LC-MeHA were termed “50 ^^L LC-MeHA”. Subsequently, an inelastic tape was affixed to the thin film, with the surrounding area covered by a layer of DMSO to ensure the isolation from the chip surface. A lateral pulling force was applied on the inelastic tape to remove the thin film from the electrodes using Universal Mechanical Tester (UMT, Bruker), with a rate of 0.1 mm s-1. The pulling force (N) versus thin film displacement (mm) were recorded and plotted.

[0153] Ex Vivo Characterization of Wearable Aptamer-Based Biosensors. Toevaluate the performance of the aptamer-based biosensors for glucose or lactate detection, the target analyte in an artificial ISF solution was introduced to the sensor for 20-minute incubation. To calculate the limit-of-detection, the kinetic differential measurement (KDM) versus the log of concentration were plotted for both glucose and lactate detection experiments. The limit of detection was calculated after determining the limit of blank (LOB) which is equal to the y-value of the fitted regression line. LOB is defined as LOB = KDMblank + 1.96× σblank where σblank is the blank standard deviation and 1.96 was the score for calculating the limit of detection within a 95% confidence interval. The limit-of-detection is calculated as limit-of-detection = 10(^^^^^^-^^) / ^^where, ^^ is the x-intercept of the regression line and m is the slope of the regression line of thelimit-of-detection curve, which signifies the sensitivity of detection. The selectivity of the biosensors was verified by calculating the electrochemical signal gain measured at 60 Hz, in a potential range of 0 to -0.6 V with a scan rate of 0.1 V s-1, before and after adding the target analyte as well as other common interfering biomolecules found in artificial ISF. The final ex vivo characterization was performed using porcine skin model. The porcine ear skin was rinsed with DI water, cut into 1 cm by 1 cm square, then pre-equilibrated in SELEX buffer with various concentrations of glucose (0, 2, 5, 10, 15, 20 mM) or lactate (0, 0.5, 1.5, 2, 2.5, 5, 10, 20 mM) overnight. Next, glucose and lactate wearable aptamer-based biosensors were applied and fixed on the porcine skin introduced with the corresponding targets for 20 minutes. Three wearable aptamer-based biosensors were tested per concentration. Subsequently, SWV measurements were performed at two different frequencies for glucose (5 Hz and 80 Hz) and lactate (15 Hz and 200 Hz) sensing. All SWV measurements were performed against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.6 V with a scan rate of 0.1 V s-1. The methylene blue reduction peak magnitudes were measured for each device and the KDM was calculated.

[0154] Skin Penetration Test and Histology. To investigate the skin penetrationefficiency, HMN integrated electrodes were applied on the shaved dorsal skin for 20 minutes. The back skin with the micropores array located in the center was collected immediately after the rat and mouse were euthanized. The collected skins were washed with 0.9% NaCl solution, fixed with neutral buffered 10% formalin for 24 hours, then stored in 70% ethanol at 4 °C. Fixed skin samples were cryopreserved in 15% sucrose / PBS solution at 4 °C overnight. Next, samples were mounted on cork using OCT (TissueTek) and sectioned into 10 ^^m thick slides at -20 °C. Skin slides were stained by hematoxylin and eosin to visualize the histological morphology. Slides were dipped in Harris-modified hematoxylin (Sigma, HHS32) for 30 s, washed in distilled water, then counterstained with 1% Eosin Y (Sigma, E4009) for 2 minutes. Subsequently, slides were dehydrated in 75% and 95% ethanol, and cleared in Xylene. Images were captured using a Cytation-5 multimode imager (Agilent) at 20 magnification and processed using the Gen5 software (Agilent). To visualize the healing process, HMN integrated electrodes were applied on the shaved rat dorsal skin for 20 minutes. The micropore array generated by HMN was imaged 0, 10, 15 and 20 minutes upon removal of the HMN.

[0155] Quality Control Protocols. The process started by measuring electrochemicalimpedance spectroscopy, and the impedance change for the HMN attachment step relative to theaptamer immobilization step was plotted. Four distinct frequencies (1000, 100, and 31.6 and 10 Hz), ensuring they remained outside the Warburg region based on their Nyquist plot were selected (FIG. 10a). Considering both standard deviation and the highest impedance change between faulty HMN attachment and acceptable attachment, 31.6 Hz was chosen for the single- frequency electrochemical impedance measurement (FIG. 10b). A total of 27 glucose and 22 lactate wearable aptamer-based biosensors were fabricated, containing 14 glucose and 11 lactate devices with faulty HMN attachment. The cut-off ratio of (|ΔZ / Z=500%|) for this step of the quality control, was determined by comparing the calculated impedance changes of all the devices (FIG.10c and 10d). Additionally, the background signal in the square wave voltammetry was significantly lower for the devices with imperfect attachment or with bad needle penetration, especially at higher frequency scans. By comparing all the electrochemical scans, the currents of 30 nA at 80 Hz for the glucose devices and 100 nA at 200 Hz for the lactate devices were determined as cut-off points. Also, the signal-to-noise ratio and the root mean square noise (RMS) were obtained for the scans of each device using an open-source MATLAB code. the sensors with scans showing high RMS and low signal-to-noise ratio were excluded. An RMS ratio of 0.01 and signal-to-noise ratio of 3 were determined as the cut-off for this step of the quality control.

[0156] In Vivo Experiments in Healthy Mice. All animal procedures were approved bythe Animal Research Ethics Board of McMaster University (Investigator: L.S. AUP# 23-45). Male wildtype C57Bl / 6J at 10-15 weeks of age were fasted for 12 hours overnight. The next morning, mice were then anesthetized using isoflurane, and placed in flexible cloth restrainer with wearable aptamer-based biosensors affixed on the shaved mouse dorsal skin with Tegaderm tape. Once fully awake, blood glucose was determined using a hand-held glucometer from their tail before and after an intraperitoneal injection of 3 g kg-1D-glucose. Afterwards, mice received an insulin injection of 0.75 U kg-1, followed by subsequent blood glucose measurements. For the L-lactate sensing experiment, the initial steps involving anesthesia and setup were the same. Blood was drawn from tail pricks using heparinized capillary tubes before and after an intraperitoneal injection of 0.6g kg-1of sodium L-lactate in random fed male wildtype C57Bl / 6J mice at 10-15 weeks of age. Blood was then centrifuged at 10000 g for 10 minutes at 4 °C to separate and collect plasma, and subsequently stored at -80 degrees until needed. Plasma L- lactate levels were quantified using a commercial L-lactate assay kit as per manufacturer's instructions.

[0157] In Vivo Long-Term Stability Test. To validate the long-term sensing stability,four lactate aptamer-based biosensors were applied on the rat dorsal skin and left in place for 3 days. Rat lactate level was monitored daily using both lactate HMN-EAB patches and a benchtop lactate assay kit with fresh blood samples. During the experiment, lactate aptamer-based biosensors were securely fixed on the rat body using Tegaderm tapes and kinesiology tapes. Meanwhile, a protective cone was placed over the rat to prevent the rat from chewing the electrodes. SWV measurements were performed at 15 Hz and 200 Hz for lactate sensing on each day. All SWV measurements were performed against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.6 V with a scan rate of 0.1 V / s. The KDM was calculated using the equation shown above, then interpolated into the ex vivo lactate calibration curve to obtain the lactate level. The lactate concentration obtained from lactate aptamer-based biosensors and bench top lactate assay kit were compared.

[0158] In Vivo Multiplexing Experiments in Diabetic Rats. Animal studies wereperformed in accordance with the Guidelines for the Care and Use of Laboratory Animals and the Animal Welfare Act Regulations; all protocols were approved by the University of Waterloo Institutional Animal Care and Use Committee (Investigator: M.P. and the ethnical approval # 43809). Male Sprague Dawley rats (Charles River, 100-150 g) were used to establish the type 1 diabetic rat model by intraperitoneally injecting 75 mg kg-1of streptozotocin (STZ). Rats with blood sugar exceeding 20 mM were selected for future experiments. Rats were fasted for 4 hours prior to the experiment and then anesthetized under isoflurane during the procedure. Two lactate and two glucose wearable aptamer-based biosensor sensors were affixed on the shaved rat dorsal skin with Tegaderm tape (3M, USA) and then incubated for 1 hour. Meanwhile, 4 U kg-1of insulin was injected into the rats subcutaneously followed by continuous blood glucose measurement by glucometer (OneTouch® Ultra®, LifeScan, Inc., USA) every 5 minutes. The electrochemical signal of glucose and lactate sensor was collected simultaneously when rat blood glucose level reached certain ranges (T1: 15 – 20 mM, T2: 15 – 10 mM, T3: 5 – 10 mM and T4: 5 mM). SWV measurements were performed at two different frequencies for glucose (5 Hz and 80 Hz) and lactate (15 Hz and 200 Hz) sensing. All SWV measurements were performed against an Ag / AgCl reference electrode and Au counter electrode in a potential range of 0 to -0.5 V with a scan rate of 0.1 V / s. Meanwhile, blood samples were collected from the tail vein using microhematocrit capillary tubes (ThermoFisher, USA) at three different time points. Rat bloodserum was collected from the blood samples by a 10-minute centrifugation at 2000 g and at 4 °C and tested for lactate concentration using the benchtop L-lactate assay kit (Abcam, ab65344).

[0159] Statistical analysis. All statistical analyses were conducted using GraphPadPrism 9. The statistical difference between groups in specificity test was analyzed using ordinary one-way ANOVA with Dunnett’s multiple comparison test. The significance of statistical difference was calculated with 99% confidence interval (alpha threshold = 0.1) and shown in GP style (0.1234 (ns), 0.0332 (*), 0.0021 (**), 0.0002 (***), p <0.0001 (****)) in graphs. Each experiment was done using three different sensors (N=3). All data is expressed as mean ± standard deviation.

[0160] For upward and downward glucose and lactate measurements, the peak currentsfrom square wave voltammetry (SWV) curves were used to calculate the kinetic differential measurement (KDM) and building the calibration curves. Additionally, the mean of all the electrochemical signals (i.e. peak current from SWV) for each concentration (I) was normalized based on the blank signal (Iblank) using the formula I normalized = I – I blank. The blank signal was obtained upon incubation of the biosensor with unspiked artificial ISF samples. The sigmoidal 4PL nonlinear regression model was used to interpolate the concentration of glucose and lactate during in vivo measurements from the calibration curves.

[0161] To calculate the limit-of-detection, the KDM values versus the log ofconcentration were plotted for glucose and lactate detection experiments. The limit-of- detection was calculated after determining the limit-of-blank = KDMblank+ 1.96× σblank, where σblank is the standard deviation of the blank and 1.96 is the score for calculating the limit-of- detection within a 95% confidence interval. The limit-of-detection = 10(^^^^^^-^^) / ^^is then calculated, where ^^ is the x-intercept of the regression line and m is theline of the limit-of-detection curve, which signifies the sensitivity of detection.

[0042]

[0162] Conclusion

[0163] An aptamer-based electrochemical sensor, integrated with hydrogelmicroneedles, for the real-time and in vivo analysis of biomarkers in ISF is disclosed herein. This device combines in situ biomarker measurement using ultrasensitive electrochemical readout with robust mechanical strength of hydrogel microneedles for effective skin penetration and ISF sampling. The wearable aptamer-based biosensor demonstrates high sensitivity and specificityfor detecting clinically relevant concentrations of both glucose and lactate in real skin environments and in vivo. The performance of the wearable aptamer-based biosensor was studied using both live mouse and rat models. The wearable aptamer-based biosensor showed stable electrochemical signals, enabling accurate and continuous lactate monitoring, which correlated with blood lactate for at least three days in an awake diabetic rat. The in vivo results obtained from healthy mice and diabetic rats demonstrated the efficacy of the wearable aptamer-based biosensor in monitoring fluctuations in glucose and lactate levels across two distinct animal models. In particular, the utilization of the wearable aptamer-based biosensor for simultaneous monitoring of glucose and lactate highlights its potential for tailoring personalized treatment approaches for the management of diabetes. The wearable aptamer-based biosensor serves as a universal sensing platform, enabling the tracking of a diverse range of clinically important biomarkers using different aptamers or other probes, which cannot be achieved enzymatically. This demonstrates the potential of applying the wearable biosensor to a broad range of targets for various continuous monitoring technologies.

[0164] While the present disclosure has been described with reference to examples, it isto be understood that the scope of the claims should not be limited by the embodiments set forth in the examples, but should be given the broadest interpretation consistent with the description as a whole.

[0165] All publications, patents and patent applications are herein incorporated byreference in their entirety to the same extent as if each individual publication, patent or patent application was specifically and individually indicated to be incorporated by reference in its entirety. Where a term in the present disclosure is found to be defined differently in a document incorporated herein by reference, the definition provided herein is to serve as the definition for the term.REFERENCES [1] F. Tehrani, H. Teymourian, B. Wuerstle, J. Kavner, R. Patel, A. Furmidge, R. Aghavali, H. Hosseini-Toudeshki, C. Brown, F. Zhang, K. Mahato, Z. Li, A. Barfidokht, L. Yin, P. Warren, N. Huang, Z. Patel, P. P. Mercier, J. Wang, Nat Biomed Eng 2022, 6, 1214.[2] P. C. Pandey, S. Shukla, S. A. Skoog, R. D. Boehm, R. J. Narayan, Sensors (Switzerland) 2019,19, 571.[3] N. Kashaninejad, A. Munaz, H. Moghadas, S. Yadav, M. Umer, Chemosensors 2021, 9, 83.[4] P. GhavamiNejad, A. GhavamiNejad, H. Zheng, K. Dhingra, M. Samarikhalaj, M. Poudineh, Adv Healthc Mater 2023, 12, DOI 10.1002 / adhm.202202362.[5] I. Lee, D. Probst, D. Klonoff, K. Sode, Biosens Bioelectron 2021, 181, 113054.[6] H. Zheng, A. GhavamiNejad, P. GhavamiNejad, M. Samarikhalaj, A. Giacca, M. Poudineh, ACS Sens 2022, 7, 2387.[7] M. N. Chien, Y. J. Chen, C. H. Bai, J. T. Huang, Micromachines (Basel) 2022, 13, DOI10.3390 / mi13030478.[8] H. Teymourian, A. Barfidokht, J. Wang, Chem Soc Rev 2020, 49, 7671.[9] M. Sharma, Applications of Targeted Nano Drugs and Delivery Systems: Nanoscience and Nanotechnology in Drug Delivery 2019, 499.

[0010] M. Dervisevic, N. H. Voelcker, ACS Mater Lett 2023, 5, 1851.

[0011] Y. Wu, F. Tehrani, H. Teymourian, J. Mack, A. Shaver, M. Reynoso, J. Kavner, N. Huang, A. Furmidge, A. Duvvuri, Y. Nie, L. M. Laffel, F. J. Doyle III, M.-E. Patti, E. Dassau, J. Wang, N. Arroyo-Currás, Anal Chem 2022, 94, 8335.

[0012] K. M. Saifullah, Z. Faraji Rad, Adv Mater Interfaces 2023, 10, DOI 10.1002 / admi.202201763.

[0013] F. Keyvani, H. Zheng, M. R. Kaysir, D. F. Mantaila, P. Ghavami Nejad, F. A. Rahman, J. Quadrilatero, D. Ban, M. Poudineh, Angewandte Chemie - International Edition 2023, 62, DOI10.1002 / anie.202301624.

[0014] M. Reynoso, A.-Y. Chang, Y. Wu, R. Murray, S. Suresh, Y. Dugas, J. Wang, N. Arroyo- Currás, Biosens Bioelectron 2023, 244, 956.

[0015] K. Y. Goud, C. Moonla, R. K. Mishra, C. Yu, R. Narayan, I. Litvan, J. Wang, ACS Sens 2019,4, 2196.

[0016] H. Teymourian, C. Moonla, F. Tehrani, E. Vargas, R. Aghavali, A. Barfidokht, T. Tangkuaram, P. P. Mercier, E. Dassau, J. Wang, Anal Chem 2020, 92, 2291.

[0017] Y. Dai, J. Nolan, E. Madsen, M. Fratus, J. Lee, J. Zhang, J. Lim, S. Hong, M. A. Alam, J. C. Linnes, H. Lee, C. H. Lee, ACS Appl Mater Interfaces 2023, 15, 56773.

[0018] R. K. Mishra, K. Y. Goud, Z. Li, C. Moonla, M. A. Mohamed, F. Tehrani, H. Teymourian, J. Wang, J Am Chem Soc 2020, 142, 5991.

[0019] M. Friedel, B. Werbovetz, A. Drexelius, Z. Watkins, A. Bali, K. W. Plaxco, J. Heikenfeld, Lab Chip 2023, 23, 3289.

[0020] D. Chan, J. C. Chien, E. Axpe, L. Blankemeier, S. W. Baker, S. Swaminathan, V. A. Piunova, D. Y. Zubarev, C. L. Maikawa, A. K. Grosskopf, J. L. Mann, H. T. Soh, E. A. Appel, Advanced Materials 2022, 34, DOI 10.1002 / adma.202109764.

[0021] S. Li, J. Dai, M. Zhu, N. Arroyo-Currás, H. Li, Y. Wang, Q. Wang, X. Lou, T. E. Kippin, S. Wang, K. W. Plaxco, H. Li, F. Xia, ACS Nano 2023, 17, 18525.

[0022] P. Dauphin-Ducharme, K. W. Plaxco, Anal Chem 2016, 88, 11654.

[0023] H. Li, P. Dauphin-Ducharme, G. Ortega, K. W. Plaxco, J Am Chem Soc 2017, 139, 11207.

[0024] R. J. White, K. W. Plaxco, Anal Chem 2010, 82, 73.

[0025] H. Chang, M. Zheng, X. Yu, A. Than, R. Z. Seeni, R. Kang, J. Tian, D. P. Khanh, L. Liu, P. Chen, C. Xu, Advanced Materials 2017, 29, 1.

[0026] M. T. Poldervaart, B. Goversen, M. De Ruijter, A. Abbadessa, F. P. W. Melchels, F. C. Öner, W. J. A. Dhert, T. Vermonden, J. Alblas, PLoS One 2017, 12, DOI10.1371 / JOURNAL.PONE.0177628.

[0027] S. H. Park, Y. J. Kim, E. Kostal, V. Matylitskaya, S. Partel, W. H. Ryu, Biosens Bioelectron 2023, 220, 114912.

[0028] D. D. Zhu, Y. R. Tan, L. W. Zheng, J. Z. Lao, J. Y. Liu, J. Yu, P. Chen, ACS Appl Mater Interfaces 2023, 15, DOI10.1021 / ACSAMI.3C00573 / SUPPL_FILE / AM3C00573_SI_001.PDF.

[0029] P. J. J. Huang, J. Liu, Angewandte Chemie - International Edition 2023, 62, DOI10.1002 / ANIE.202212879.

[0030] N. Nakatsuka, K.-A. Yang, J. M. Abendroth, K. M. Cheung, X. Xu, H. Yang, C. Zhao, B. Zhu, Y. S. Rim, Y. Yang, P. S. Weiss, M. N. Stojanović, A. M. Andrews, Science (1979) 2018, 362,319.

[0031] B. S. Ferguson, D. A. Hoggarth, D. Maliniak, K. Ploense, R. J. White, N. Woodward, K. Hsieh, A. J. Bonham, M. Eisenstein, T. E. Kippin, K. W. Plaxco, H. T. Soh, Sci Transl Med 2013, 5.

[0032] A. Idili, N. Netzahualc´, N. Arroyo-Currás, C. Currás, K. L. Ploense, A. T. Csordas, M. Kuwahara, T. E. Kippin Dfg, K. W. Plaxco, Chem Sci 2019, 10, 8164.

[0033] H. Chang, M. Zheng, X. Yu, A. Than, R. Z. Seeni, R. Kang, J. Tian, D. P. Khanh, L. Liu, P. Chen, C. Xu, Advanced Materials 2017, 29, 1.

[0034] P. Makvandi, M. Kirkby, A. R. J. Hutton, M. Shabani, C. K. Y. Yiu, Z. Baghbantaraghdari, R. Jamaledin, M. Carlotti, B. Mazzolai, V. Mattoli, R. F. Donnelly, Nanomicro Lett 2021, 13,DOI 10.1007 / s40820-021-00611-9.

[0035] F. Bakhshandeh, S. Saha, P. Sen, S. Sakib, R. Maclachlan, F. Kanji, E. Osman, L. Soleymani, Biosens Bioelectron 2023, 235, 115359.

[0036] K. P. Foley, S. Zlitni, E. Denou, B. M. Duggan, R. W. Chan, J. C. Stearns, J. D. Schertzer, Nat Commun 2018, 9, DOI 10.1038 / s41467-018-07146-5.

[0037] A. P. Drabovich, M. Berezovski, V. Okhonin, S. N. Krylov, Anal Chem 2006, 78, 3171.

[0038] S. Akgönüllü, S. Koyun, H. Yavuz, A. Erdem, A. Denizli, in Biomedical Engineering Technologies: Volume 1 (Eds.: M. R. Ossandon, H. Baker, A. Rasooly), Springer US, New York, NY, 2022, pp.37–56.

[0039] Y. Wu, F. Tehrani, H. Teymourian, J. Mack, A. Shaver, M. Reynoso, J. Kavner, N. Huang, A. Furmidge, A. Duvvuri, Y. Nie, L. M. Laffel, F. J. Doyle, M. E. Patti, E. Dassau, J. Wang, N. Arroyo-Currás, Anal Chem 2022, 94, 8335.

[0040] M. Dervisevic, N. H. Voelcker, ACS Mater Lett 2023, 5, 1851.

[0041] K. L. Wolkowicz, E. M. Aiello, E. Vargas, H. Teymourian, F. Tehrani, J. Wang, J. E. Pinsker, F. J. Doyle, M. E. Patti, L. M. Laffel, E. Dassau, Bioeng Transl Med 2021, 6, DOI10.1002 / btm2.10201.

[0042] A. Victorious, Z. Zhang, D. Chang, R. Maclachlan, R. Pandey, J. Xia, J. Gu, T. Hoare, L. Soleymani, Y. Li, Angewandte Chemie 2022, 61, 1.

[0043] S. Lin, X. Cheng, J. Zhu, B. Wang, D. Jelinek, Y. Zhao, T.-Y. Wu, A. Horrillo, J. Tan, J. Yeung, W. Yan, S. Forman, H. A. Coller, C. Milla, S. Emaminejad, Sci. Adv 2022, 8, 4539.

Claims

CLAIMS:

1. An electrochemical biosensor for detecting at least one target analyte in interstitial fluidof a subject, comprising: a) a microneedle array comprising a highly-crosslinked hydrogel polymer;b) at least one electrochemical chip comprising at least one transducer surfaceconfigured for generating a signal in the presence of the at least one target analyte; and c) a low crosslinked hydrogel polymer,wherein the low crosslinked hydrogel polymer is configured to couple the microneedle array with the at least one transducer surface of the at least one electrochemical chip.

2. The biosensor of claim 1, further comprising one or more probes functionalized on theat least one transducer surface that bind to the at least one target analyte, the one or more probes configured to generate a measurable signal in the presence of the at least one target analyte.

3. The biosensor of claim 1, wherein the transducer surface is configured for reacting withthe at least one target analyte to generate a measurable signal in the presence of the at least one target analyte.

4. The biosensor of claim 1, further comprising (i) a first probe that binds to a first targetanalyte functionalized on a first transducer surface and (ii) the first probe and a second probe that binds to a second target analyte functionalized on a second transducer surface, each transducer surface configured to generate measurable signals, wherein a difference between the measurable signal on the first transducer surface and the measurable signal on the second transducer surface is indicative of the presence of the second target analyte.

5. The biosensor of any one of claims 1 to 4, wherein the highly crosslinked hydrogelpolymer has a crosslinking density that is at least 10% greater than that of the low-crosslinked hydrogel polymer.

6. The biosensor of any one of claims 1 to 4, wherein the highly crosslinked hydrogelpolymer has a crosslinking density that is at least 20% greater than that of the low-crosslinked hydrogel polymer.

7. The biosensor of any one of claims 1 to 4, wherein the highly crosslinked hydrogelpolymer has a crosslinking density that is at least 30% greater than that of the low-crosslinked hydrogel polymer.

8. The biosensor of any one of claims 1 to 4, wherein the highly crosslinked hydrogelpolymer has a crosslinking density that is at least 40% greater than that of the low-crosslinked hydrogel polymer.

9. The biosensor of any one of claims 1 to 8, wherein the highly-crosslinked and the lowcrosslinked hydrogel polymer, each independently, comprise gelatin, hyaluronic acid, alginate, chitosan, collagen, or combinations thereof.

10. The biosensor of any one of claims 1 to 9, wherein the highly-crosslinked and the lowcrosslinked hydrogel polymer, each independently, comprise methacrylated gelatin, methacrylated hyaluronic acid, methacrylatedalginate, methacrylated chitosan, methacrylated collagen, dopamine-functionalized hyaluronic acid or combinations thereof.

11. The biosensor of any one of claims 1 to 10, wherein the highly-crosslinked and the lowcrosslinked hydrogel polymer, each independently, comprise methacrylated hyaluronic acid.

12. The biosensor of any one of claims 1 to 11, wherein the highly-crosslinked and the lowcrosslinked hydrogel polymer, each independently, further comprise a crosslinking agent.

13. The biosensor of claim 12, wherein the crosslinking agent comprises N,Nʹ-methylenebisacrylamide (MBA), PEG–maleimide, HA–thiol, enzymes, Carbodiimide-based (EDC / NHS), or combinations thereof.

14. The biosensor of any one of claims 1 to 13, wherein the highly-crosslinked hydrogelpolymer comprises a porosity of about 25% to about 50%.

15. The biosensor of any one of claims 1 to 14, wherein the low crosslinked hydrogelpolymer comprises a porosity of above about 50% up to about 75%.

16. The biosensor of any one of claims 1 to 15, wherein the at least one electrochemicalchip comprises one or more working electrodes.

17. The biosensor of claim 16, wherein the at least one electrochemical chip furthercomprises a counter electrode.

18. The biosensor of claim 16, wherein the at least one electrochemical chip furthercomprises a reference electrode.

19. The biosensor of claim 16, wherein the counter electrode is a reference electrode.

20. The biosensor of any one of claims 1 to 19, wherein the at least one transducer surfacecomprises metals, metal alloys, metal oxides, superconductors, semi-conductors, carbon-based materials, conductive polymers, or combinations thereof.

21. The biosensor of claim 20, wherein the at least one transducer surface comprises metals,metal alloys, carbon-based materials, or combinations thereof.

22. The biosensor of claim 20 or 21, wherein the transducer surface comprises gold.

23. The biosensor of any one of claims 1 to 22, wherein the one or more probes isfunctionalized on the at least one transducer surface via chemical bonding, an intermediate linker, physical adsorption, embedded in inks, embedded in polymers, or a combination thereof.

24. The biosensor of claim 23, wherein the chemical bonding occurs via thiol chemistry,gold chemistry, or a combination thereof.

25. The biosensor of any one of claims 1 to 24, wherein the one or more probes is a nucleicacid probe, an antibody probe, or a combination thereof.

26. The biosensor of claim 25, wherein the one or more nucleic acid probes comprises asingle-stranded DNA, single-stranded RNA, double-stranded DNA, double stranded RNA, DNAzyme, an aptamer, or a combination thereof.

27. The biosensor of any one of claims 25 or 26, wherein the one or more nucleic acidprobes comprises an aptamer.

28. The biosensor of any one of claims 25 to 27, wherein the one or more nucleic acidprobes comprises a detection moiety.

29. The biosensor of claim 28, wherein the detection moiety comprises an electroactivespecies, a redox active species, or a combination thereof.

30. The biosensor of claim 29, wherein the redox active species comprises methylene blue.

31. The biosensor of claim 25, wherein the one or more antibody probes comprises an anti-insulin antibody, an anti-troponin antibody, or an anti-bovine serum albumin antibody, or a combination thereof.

32. The biosensor of claim 25 or 31, wherein the one or more antibody probes comprises adetection moiety.

33. The biosensor of claim 32, wherein the detection moiety comprises an electroactivespecies, a redox active species, or a combination thereof.

34. The biosensor of claim 33, wherein the redox active species comprises methylene blue.

35. The biosensor of any one of claims 1 to 34, wherein the signal is a current, a potentialor an impedance.

36. The biosensor of any one of claims 1 to 35, wherein the at least one target analyte is abiomolecule.

37. The biosensor of any one of claims 1 to 36, wherein the at least one target analyte isglucose, lactate, or a combination thereof.

38. The biosensor of any one of claims 1 to 37, wherein the microneedles are arranged in awearable patch configured for transdermal sensing.

39. The biosensor of any one of claims 1 to 38, for use in health monitoring, screening,diagnostics, or a combination thereof.

40. A transdermal patch comprising the biosensor of any one of claims 1 to 39.

41. A method for transdermal biosensing of at least one target analyte in interstitial fluid ofa subject, the method comprising:a) applying a transdermal patch comprising the biosensor of any one of claims1 to 39 onto the subject; and b) detecting the signal from the at least one electrochemical chip of thebiosensor, whereby the at least one transducer surface generates an electrochemical signal in the presence of the at least one target analyte, wherein the electrochemical signal is indicative of the concentration of the at least one target analyte in the subject.

42. The method of claim 41, wherein the signal is monitored continuously over time.

43. The method of any one of claim 41 or 42, wherein the signal comprises a current, apotential or an impedance.

44. The method of any one of claims 41 to 43, wherein the signal comprises current.

45. The method of any one of claims 41 to 44, wherein the signal is measured using squarewave voltammetry.

46. The method of any one of claims 41 to 45, wherein the signal is used to calculate akinetic differential measurement (KDM).

47. A method for fabricating the electrochemical biosensor of any one of claims 1 to 39,the method comprising: a) providing a microneedle array comprising a highly-crosslinked hydrogelpolymer, the microneedle array prepared by drying the highly-crosslinked hydrogel polymer in a mold; b) providing at least one electrochemical chip comprising at least onetransducer surface; and c) coupling the microneedle array with the at least one transducer surface ofthe electrochemical chip using a low crosslinked hydrogel polymer.

48. The method of claim 47, further comprising adjusting pH of the highly-crosslinkedhydrogel polymer before drying in the mold.

49. The method of claim 47, further comprising exposing the highly-crosslinked hydrogelpolymer to ultraviolet (UV) light after drying in the mold.

50. The method of any one of claims 47 to 49, further comprising crosslinking the highly-crosslinked and low crosslinked hydrogel polymers after coupling the microneedle array with the at least one transducer surface of the at least one electrochemical chip.

51. The method of any one of claims 47 to 50, further comprising calibrating the biosensorby measuring a signal upon increasing and decreasing target analyte concentrations.

52. The method of claim 51, wherein a KDM is calculated to generate a calibration curve.

Images