WO2025096893A1 - Magnetic microneedle sensor arrays and related aspects for molecular monitoring - Google Patents

Abstract

Provided herein are methods of detecting target molecules using electrochemical sensors that comprise biomolecular receptor-bound redox reporters. Related electrochemical sensor devices, kits, systems, computer readable media, and additional methods are also provided.

Description

MAGNETIC MICRONEEDLE SENSOR ARRAYS AND RELATED ASPECTS FOR MOLECULAR MONITORING

CROSS-REFERENCE TO RELATED APPLICATIONS

[001] This application claims priority to U.S. Provisional Patent Application Ser. No. 63/547,180, filed November 3, 2023, the disclosure of which is incorporated herein by reference.

BACKGROUND

[002] Electrochemical, aptamer-based (E-AB) sensors are analytical platforms that achieve continuous monitoring of specific molecular targets in vivo. E- AB sensors present an architecture typically consisting of three elements (FIG. 1A): 1 ) a self-assembled monolayer (SAM) of target-binding, alkanethiol-functionalized nucleic-acid aptamers or other bioreceptor, 2) an electrode-blocking SAM of alkanethiols to prevent undesired electrochemical reactions and confer biocompatibility to the electrode surface, and 3) a redox reporter sensitive to targetbinding events. The redox reporter, typically methylene blue (MB), is attached to the terminal end of the aptamer, opposite to the electrode attachment terminus. In the presence of target, aptamer molecules reversibly undergo binding-induced conformational changes that presumably bring the reporter closer to the electrode surface, causing a change in the electron transfer rate between the reporter and the electrode (FIG. 1B), which can be easily measured electrochemically. Aptamer binding in E-AB sensors is at dynamic equilibrium, reversibly switching between bound and unbound states at rates of milliseconds. This behavior makes E-AB sensors ideal for continuous monitoring applications. Moreover, because their working principle mimics the binding-induced conformational changes seen in naturally occurring chemoreceptors in the body, E-AB sensors tolerate prolonged measurements in complex matrices such as unprocessed biological fluids.

[003] E-AB sensors can be successfully interrogated via chronoamperometry, differential pulse techniques such as square-wave voltammetry and differential pulse voltammetry, alternating current voltammetry, and electrochemical impedance spectroscopy. Ultimately, the choice of technique is typically determined by the final intended application of the E-AB sensor. For example, the simplicity of the voltage program in chronoamperometry is ideal for drift-free measurements at sub-second interrogation frequencies, which may be needed for the study of fast biological processes like neurotransmitter modulation in the brain. Electrochemical impedance, in contrast, offers the convenience of interrogating E-AB sensors in a label-free manner, without using a redox reporter. However, the vast majority of reported E-AB sensors have been interrogated by pulse techniques and, in particular, by square wave voltammetry. This widespread use likely arose because pulsed techniques differentially remove currents originating from charging the electrode-electrolyte double layer, significantly improving the signal-to-noise ratio of E-AB measurements. Yet, pulsed techniques also remove valuable electrochemical information regarding sensor stability (e.g., the capacitive current reports on monolayer stability) and differential voltage pulsing also strains the E-AB interface causing faster loss of signal.

[004] Cyclic voltammetry (CV) is frequently used for the surface characterization of E-AB sensors, as this technique provides valuable information regarding monolayer stability (by proxy of double layer capacitance) and surface coverage of the redox reporter-modified aptamer (from faradaic peak areas). However, CV is not commonly used for the direct interrogation of E-AB sensors, in part because sensors with defective blocking monolayers or redox reporter-modified aptamers with slow electron transfer kinetics present large capacitive currents that can hide the faradaic waves of methylene blue, resulting in low signal-to-noise E-AB measurements. Moreover, for many E-AB sensors, CV peak currents do not change significantly with increasing target concentrations.

[005] Accordingly, there is a need for additional methods, and related aspects, of interrogating electrochemical sensors, including those using cyclic voltammetry.

SUMMARY

[006] The present disclosure relates, in certain aspects, to methods, systems, kits, and computer readable media of use in detecting target molecules in interstitial fluid (ISF) of or in subjects. Some aspects, for example, provide magnetic wearable microneedle sensor arrays that enable the continuous, minimally invasive sensing of molecules in vivo via the dermal ISF, both in animal models and in humans. Interstitial fluid is a body compartment that rapidly equilibrates with blood, allowing the monitoring of systemic biomarkers, among other target molecules. However, the dermis ISF is not directly irrigated by capillaries nor innervated, allowing for painless, minimally invasive sensing. Some embodiments include the use of the technology in toxicology and pharmacokinetic (PK)/pharmacodynamic (PD) labs for drug development applications. In some embodiments, the approaches disclosed herein are translated into wearable sensor devices to be placed, for example, on the ear lobe of customers for continuous health status monitoring. These and other aspects will be apparent upon a complete review of the present disclosure, including the accompanying figures.

[007] According to various embodiments, an electrochemical sensor device is presented. The electrochemical sensor device includes a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode. The electrochemical sensor device also includes at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected, or connectable, to a detector, and a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer. In addition, the electrochemical sensor device also includes a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule. [008] Various optional features of the above embodiments include the following. The electrochemical sensor device is configured to generate one or more square wave or cyclic voltammograms using square wave voltammetry (SWV) and/or cyclic voltammetry (CV) and determine a change in peak-to-peak separation, EP.T, from the voltammograms to detect the target molecule when the biomolecular receptors bind the target molecule. The electrochemical sensor device is configured to perform drift correction. The drift correction is performed using kinetic differential measurements (KDM). The electrochemical sensor device is configured to generate substantially stable baseline electrochemical signals. A thickness of the electrically conductive layer is less than about 1 pm. The body structure comprises a substantially circular cross-sectional shape, a substantially square cross-sectional shape, or a substantially rectangular cross-sectional shape. The plurality of microneedles is arrayed in a substantially circular format, a substantially square format, or a substantially rectangular format. At least one of the microneedles in the plurality of microneedles has a substantially circular, a substantially square, or a substantially rectangular cross-sectional shape. At least of the microneedles in the plurality of microneedles has a substantially conical shape, a substantially pyramidal shape, or a substantially tapered shape. Edges of at least one pair of microneedles in the plurality of microneedles are separated from one another by about 5 mm or less. The detector comprises an electrochemical analyzer. The electrochemical analyzer comprises a potentiostat. The electrically conductive layer is configured to wirelessly connect to the detector. The detector is configured to measure square wave voltammograms. An electrically insulating cover structure is attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure. The body structure and the plurality of microneedles are fabricated using a stereolithography (SLA) three- dimensional (3D) printing technique.

[009] Various additional optional features of the above embodiments include the following. The plurality of microneedles comprises twelve microneedles, wherein seven of the microneedles are working electrodes, wherein two of the microneedles are reference electrodes, and wherein three of the microneedles are counter electrodes. The body structure comprises a cross-sectional dimension of about 30 mm or less. The magnet comprises a permanent magnet. The magnet comprises a neodymium ring magnet. The magnet comprises a cross-sectional dimension of about 25 mm or less. The magnet comprises a thickness of about 5 mm or less. The plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged aptamers. The subject comprises a mammalian subject. The mammalian subject comprises a human subject. The mammalian subject is from an order Rodentia. The metallic structure is configured to maintain a substantially constant and substantially homogeneous pressure across the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject. The metallic structure is configured to maintain a substantially constant penetration depth of the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject. The metallic structure is disposed at least proximal to an ear of the subject. A metallic plate comprises the metallic structure. The metallic structure is implanted in the subject. The metallic structure comprises a biocompatible plate fabricated from a medical grade, ferromagnetic steel. The plurality of microneedles is fabricated integral with the body structure. The electrically conductive layer comprises a chromium layer and/or a gold layer. The electrically conductive layer of the reference electrode further comprises at least one silver layer. The electrically conductive layer of the counter electrode further comprises at least one platinum layer.

[010] Various other optional features of the above embodiments include the following. At least one microneedle of the plurality of microneedles extends from the surface of the body structure by about 0.5 mm to about 2 mm. A base portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 400 pm. A tip portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 100 pm. The biomolecular receptor comprises an aptamer. The aptamer comprises a 5’- thiol modification and a 3’ methylene blue modification. The biomolecular receptor comprises a nucleic acid molecule (e.g., RNA, DNA, PNA, LNA, L-DNA, etc.). The redox reporters comprise methylene blue (MB) or an osmium-based complex. The target molecule comprises a therapeutic agent. The therapeutic agent comprises an antibiotic. The target molecule comprises a metabolite and/or an electrolyte. The target molecule comprises a biomolecule. The electrochemical sensor device is a wearable device. A kit comprising the electrochemical sensor device.

[011] According to various embodiments, a method of detecting a target molecule in a subject is presented. The method includes positioning a plurality of microneedles of an electrochemical sensor device in contact with a dermal interstitial fluid (ISF) of or in the subject, wherein the electrochemical sensor device comprises: a body structure comprising first and second surfaces; the plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected to a detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals; and a magnet connected to the second surface of the body structure, wherein the magnet magnetically attaches the electrochemical sensor device to a metallic structure disposed in and/or on a body of the subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with the dermal ISF of or in the subject, which dermal ISF comprises the target molecule. The method also includes detecting the electrochemical signals using the detector, thereby detecting the target molecule in the subject.

[012] Various optional features of the above embodiments include the following. The method further comprises implanting the metallic structure in the body of the subject prior to the positioning step. The method comprises implanting the metallic structure in an ear lobe of the subject prior to the positioning step. The method comprises substantially continuously monitoring the target molecule in the subject during a selected duration of time. The method comprises intermittently monitoring the target molecule in the subject during a selected duration of time. The method comprises performing a pharmacokinetic, pharmacodynamic, and/or toxicology assessment of the target molecule in the subject. The method comprises detecting the electrochemical signals in substantially real-time. The method further comprises performing drift correction of the electrochemical sensor device using one or more kinetic differential measurements (KDM). The electrochemical sensor device further comprises an electrically insulating cover structure attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure. The plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged nucleic acids. The redox reporters comprise methylene blue (MB).

[013] According to various embodiments, a method of producing an electrochemical sensor device is presented. The method includes forming a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure, and disposing at least one electrically conductive layer on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connectable to a detector. The method also includes attaching a plurality of biomolecular receptor-bound redox reporters to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters are configured to undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer, and connecting a magnet to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule, thereby producing the electrochemical sensor device.

[014] Various optional features of the above embodiments include the following. The method further comprises attaching an electrically insulating cover structure to the first surface of the body structure such that the plurality of microneedles extends through the electrically insulating cover structure.

[015] According to various embodiments, a system for detecting a target molecule in a subject is presented. The system includes a detector, and an electrochemical sensor device, comprising: a body structure comprising first and second surfaces; a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected to the detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals that are detected by the detector; and, a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule. The system also includes a processor, and a memory communicatively coupled to the processor, the memory storing non-transitory computer executable instructions which, when executed on the processor, perform operations comprising: detecting the electrochemical signals using the detector.

BRIEF DESCRIPTION OF THE DRAWINGS [016] The accompanying drawings, which are incorporated in and constitute a part of this specification, illustrate certain embodiments, and together with the written description, serve to explain certain principles of the methods, systems, and related computer readable media disclosed herein. The description provided herein is better understood when read in conjunction with the accompanying drawings which are included by way of example and not by way of limitation. It will be understood that like reference numerals identify like components throughout the drawings, unless the context indicates otherwise. It will also be understood that some or all of the figures may be schematic representations for purposes of illustration and do not necessarily depict the actual relative sizes or locations of the elements shown.

[017] FIGS. 1A and 1B schematically show that E-AB sensors undergo target binding-induced changes in electron transfer kinetics of the redox reporter that can be monitored in real time via electrochemical interrogation according to one exemplary embodiment. (A) In this work we employed three different DNA aptamers modified at the 5’ terminus with alkanethiol linkers and at the 3’ terminus with the redox reporter methylene blue. We co-deposited these modified oligonucleotides with 6-mercapto-1 -hexanol on the surface of gold electrodes via self-assembly. (B) In the presence of their target molecule, the aptamers undergo a conformational change that, presumably, brings the redox reporter closer to the electrode surface, increasing the electron transfer rate.

[018] FIG. 2A is a flow chart that schematically depicts steps in a method of detecting a target molecule in a subject according to an exemplary embodiment.

[019] FIG. 2B is a flow chart that schematically depicts steps in a method of producing an electrochemical sensor device according to an exemplary embodiment.

[020] FIG. 3 shows an algorithm according to an exemplary embodiment.

[021] FIG. 4 schematically depicts a system according to an exemplary embodiment.

[022] FIGS. 5A-5D. 3D-Printed, aptamer-based microneedle sensor arrays using magnetic skin placement. (A) Schematic representation of a fully assembled sensor patch for skin placement. Unmodified microneedle arrays were 3D-printed via stereolithography. To enable magnetic placement of sensor patches on the skin of live rodents, we encased a ring-shaped permanent magnet at the bottom of the patches. (B) Gold plated MNs allow self-assembly of biosensing monolayers, consisting of electrode-blocking alkanethiols and thiolated and redox- labelled aptamers. (C) Rapid loss of MN sensor signals upon deployment on the skin of rodents, caused by contraction of the animal skin after the device placement. (D) Addressing this problem by using magnetic attachment of MN sensor arrays which leads to stable sensor response during continuous pharmacokinetic measurements.

[023] FIGS. 6A-6C. SLA 3D printing and assembly of microneedle sensor arrays. (A) SEM micrograph of freshly 3D-printed MNs before assembly. (B) Device fabrication steps: (Bi) UV cured and washed MNs; (Bn) MNs with kapton tape mask, ready for sputtering; (Bm) MNs coated with a chromium layer followed by a gold film; (Biv) reference electrode coating with silver ink; (Bv) electrode connections to Dupont pins; and (Bvi) final assembled device, including neodymium magnetic ring at the back and insulating 3D-printed cover at the top. (C) Front and back photographs of the final assembly.

[024] FIGS. 7A and 7B. Skin penetration tests using tapered vs conical microneedle arrays. (A) Photograph of tapered MN sensor array. We tested three tapered microneedle array designs that differed in needle-to-needle lateral spacing: (Ai, Aiv) 2.2 mm; (AH, AV) 2mm; and (Am, Avi) 2.5mm. (B) Photograph of conical microneedle sensor array. We tested three conical microneedle array designs that differed in the needle-to-needle lateral spacing: (Bi, Biv) 0.8 mm; (Bn, Bv) 0.5 mm; and (Biii, Bvi) 1 mm. Colors, contrast and brightness were adjusted by the iPhone used for taking the photographs without further editing.

[025] FIGS. 8A-8C. Comparative in-vivo performance of skin taped, glued, and magnetically placed MNs. (A) Securing the devices via medical tape, we observed rapid skin disinsertion of MN devices over a period of 1 h. (B) Using super glue under and around the devices improves on-skin retention relative to using medical tape, but ultimately we observed ~80% signal loss. (C) In contrast, magnetic placement of the MN sensor arrays via a subcutaneous metallic plate leads to robust retention of sensor signals over the entire measurement period. Center panels show the output of the MN sensors when interrogated via square wave voltammetry at various square wave frequencies (e.g., 80, 150, 240, 300, 600 Hz). The right panels show sensor performance measured at a square wave frequency of 80 Hz. [026] FIGS. 9A-9F. Pharmacokinetics of tobramycin in ISF via magnetically placed MNs. These [tobramycin]isF vs t plots were measured following intravenous dosing of 20 mg/kg tobramycin via tail vein boluses. Panel (A) illustrates the raw data obtained at signal-on frequency, 300 Hz, and signal-off frequency, 30 Hz, as well as the resulting KDM trace obtained from their difference. Panel (B) depicts the calibration curve utilized for conversion of signal gain to tobramycin concentration. Three independent ex-vivo experiments were performed in 1 :1 solution of serum and PBS complemented with 2 mM MgCh Standard deviations are represented by the error bars (n=3). Panels (C), (D) and (E) show the response of three individual MN arrays placed on three independent animals. Colored dashed lines show non-linear regressions to a model of first order absorption and excretion pharmacokinetics, described in detail within the main text. Panel (F) shows the fits from all datasets with x-axes corrected to 0 to remove the offset from baseline measurements. The solid black line represents the mean response, with average time to plateau ~ 46 min post intravenous bolus. Intravenous boluses were dosed in phosphate-buffered saline with full dose delivery within 10 s.

[027] FIGS. 10A-10E. Monitoring therapeutic transport from blood to the dermis interstitial fluid (ISF) via microneedle sensor arrays. (A) Top view of the sensor array employed in this study. The platform features five working electrode (WE) channels for real-time drug monitoring, one counter electrode (CE) channel, and one Ag|AgCI pseudo-reference (RE) channel. Each WE and CE is made of four acupuncture microneedles coated with gold nanoparticles (AuNPs). The reference electrode uses five microneedles coated with silver paint, later treated in bleach to form the AgCI film. Inset: Scanning electron microscopy (SEM) micrograph showing an example microneedle coated with AuNPs. (B) To create drug sensors, the microneedles are functionalized with mixed self-assembled monolayers of hexylthioland redox reporter-modified aptamers, and electrode passivating 6- mercaptohexanol. Upon interacting with their target, the sensors undergo changes in electron transfer (eT), which can be serially interrogated via square wave voltammetry. (C) For this study we placed the microneedle sensors on the abdominal skin of male Sprague-Dawley rats, and delivered drug boluses intravenously (I.V.). (D) To prevent sensor loosening and disconnection from the skin, the microneedle platform contains a permanent magnet ring on its back. A subcutaneously implanted magnetic plate holds the devices in place, as previously reported by our group. (E) Using SACMES, a real-time data processing and visualization software, we can monitor ISF drug uptake and excretion in real time. The innovative multichannel platforms allows redundant measurements in a single rodent, drastically decreasing the number of animals needed for statistical validation.

[028] FIGS. 11A-11C. Setup of microneedle sensor array in vivo and measurement protocol. (A) Sprague-Dawley rats (300 g) were anesthetized under isoflurane. A 1 cm-wide skin cut was made on the left side of their abdomen to subcutaneously (sub-Q) slide a magnetic plate. Additionaly, their right jugular vein was cathetherized for intravenous (i.V.) dosing. Vital signs (heart rate, temperature, oxygen) were monitored throughout the 3-hour measurement periods using an oximeter. The microneedle platform contains a back-mounted magnet for magnetic attahcement of the devices on the animal skin. Electrodes were wired to a potentiostat for external data collection. (B) In this study, our measurement protocol included a 30 min-long baseline data collection (sensor equilibration period in vivo), followed by an I.V. drug bolus. We then continued recording every 120 s for 2.5 h. (C). Drugs investigated in this study, sorted by molecular weight.

[029] FIGS. 12A-12F. Vancomycin pharmacokinetics in the dermal ISF. (A) We first titrated vancomycin into three independent rats, at I.V. doses of 20 mg/kg (black), 40 mg/kg (red), and 83 mg/kg (blue). Solid symbols show the statistical mean, shaded areas the standard deviation across four WE channel measurements. To assess animal-to-animal variability, we repeated the 40 mg/kg dose with a fourth rat (red triangles vs squares). (B) Sensor calibration in 50% rat serum in phosphate- buffered saline, as a proxy for ISF. The y-axis corresponds to the gain of kinetic differential measurements. (C) Following a 20 mg/kg dose, the ISF vancomycin profile can be modeled via first order absorption and excretion (red line), with CMAX = 1.3 ± 0.4 pM, kabs = 2.68 ± 1.09 min^-1, and keiim = 2.68 ± 1.04 min^-1 [i.e., flip-flop pharmacokinetics]. The higher vancomycin doses (40 mg/kg and 83 mg/kg) caused sensor signal saturation (i.e., measurements much above the ECso), resulting in noisy plateau concentrations (D, E) or oversaturation (F). The solid black lines are point connectors to illustrate trends.

[030] FIGS. 13A-13D. Tobramycin pharmacokinetics in the dermal ISF. (A)

We dosed tobramycin into two independent rats, at I.V. doses of 28 mg/kg (black) and 40 mg/kg (red). Solid symbols show the statistical mean, shaded areas the standard deviation across four WE channel measurements. The upper data shows vehicle doses of volume identical to the positive controls, delivered to two additional, independent animals to demonstrate that sensor responses are not caused by the dose vehicle. (B) Sensor calibration in 50% rat serum in phosphate-buffered saline, as a proxy for ISF. The y-axis corresponds to the gain of kinetic differential measurements. (C) Following a 28 mg/kg dose, the ISF tobramycin profile can be modeled via first order absorption and excretion (red line), with CMAX = 67 ± 18 pM, kabs = 4.0 ± 0.8 min’¹, and keiim = 1.8 ± 0.3 min’¹. (D) The higher dose (40 mg/kg) presented a biphasic concentration plateau after the absorption phase, which may be attributed to competing excretion kinetics such as those related to efflux pumps.

[031] FIGS. 14A-14E. Microneedle patch monitors hydrophobic molecules in the dermal ISF. (A) We tested a hydrophobic amino acid phenylalanine with two doses (30 mg/kg and 60 mg/kg) injected intravenously. Sensor response over time for each dose is shown in the graph. (B) We converted these sensor readings into actual analyte concentrations with the assistance of an ex vivo calibration curve obtained in 50% rat serum solution. (C) Using a one-compartment pharmacokinetic model, we analyzed the analyte concentration profiles for both phenylalanine doses. Compared to the lower dose (30 mg/kg), the higher dose (60 mg/kg) showed a faster absorption and excretion of the amino acid in the skin fluid. (D) We tested if the microneedle patch could monitor other hydrophobic drugs. A single injection of irinotecan (31 mg/kg) intravenously did not result in a remarkable change, which may be due to limited movement of the drug from the bloodstream to the skin fluid. (E) Similar to irinotecan, injections of another chemotherapeutic drug doxorubicin at two doses (5 mg/kg and 10.7 mg/kg) did not produce a significant change in the sensor signal. Solid lines represent the average of measurements from multiple sensors (N=3 or 4), and shaded areas represent the variations.

DEFINITIONS

[032] In order for the present disclosure to be more readily understood, certain terms are first defined below. Additional definitions for the following terms and other terms may be set forth through the specification. If a definition of a term set forth below is inconsistent with a definition in an application or patent that is incorporated by reference, the definition set forth in this application should be used to understand the meaning of the term.

[033] As used in this specification and the appended claims, the singular forms “a,” “an,” and “the” include plural references unless the context clearly dictates otherwise. Thus, for example, a reference to “a method” includes one or more methods, and/or steps of the type described herein and/or which will become apparent to those persons skilled in the art upon reading this disclosure and so forth.

[034] It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments only and is not intended to be limiting. Further, unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this disclosure pertains. In describing and claiming the methods, systems, and component parts, the following terminology, and grammatical variants thereof, will be used in accordance with the definitions set forth below.

[035] About: As used herein, “about” or “approximately” or “substantially” as applied to one or more values or elements of interest, refers to a value or element that is similar to a stated reference value or element. In certain embodiments, the term “about” or “approximately” or “substantially” refers to a range of values or elements that falls within 25%, 20%, 19%, 18%, 17%, 16%, 15%, 14%, 13%, 12%, 11 %, 10%, 9%, 8%, 7%, 6%, 5%, 4%, 3%, 2%, 1 %, or less in either direction (greater than or less than) of the stated reference value or element unless otherwise stated or otherwise evident from the context (except where such number would exceed 100% of a possible value or element).

[036] Bind: As used herein, “bind,” in the context of pathogen detection, refers to a state in which a first chemical structure (e.g., a therapeutic agent) is sufficiently associated a second chemical structure (e.g., a bioreceptor) such that the association between the first and second chemical structures can be detected.

[037] Detecting: As used herein, “detecting,” “detect,” or “detection” refers to an act of determining the existence or presence of one or more target analytes in a sample. [038] Biomolecule. As used herein, "biomolecule” refers to an organic molecule produced by a living organism. Examples of biomolecules, include macromolecules, such as nucleic acids, proteins, carbohydrates, and lipids.

[039] Bioreceptor. As used herein, “bioreceptor” refers to a biochemical structure that receives or binds other chemical structures (e.g., therapeutic agents, nucleic acids, proteins, metabolites, and the like).

[040] Sample: As used herein, “sample” means anything capable of being analyzed using a device or system disclosed herein. Exemplary sample types include environmental samples and biological samples. In some embodiments, subjects exhale, spit, sneeze, cough, and/or the like to produce aerosolized samples.

[041] Specifically Bind: As used herein, "specifically bind,” in the context of pathogen detection, refers to a state in which substantially only target chemical structures (e.g., biomolecules) are sufficiently associated with a corresponding or cognate binding agent, to the exclusion of non-target chemical structures, such that the association between the target chemical structures and the binding agent can be detected.

[042] System: As used herein, "system" in the context of analytical instrumentation refers a group of objects and/or devices that form a network for performing a desired objective.

[043] Subject: As used herein, “subject” refers to an animal, such as a mammalian species (e.g., human) or avian (e.g., bird) species. More specifically, a subject can be a vertebrate, e.g., a mammal such as a mouse, a primate, a simian or a human. Animals include farm animals (e.g., production cattle, dairy cattle, poultry, horses, pigs, and the like), sport animals, and companion animals (e.g., pets or support animals). A subject can be a healthy individual, an individual that has or is suspected of having a disease or a predisposition to the disease, or an individual that is in need of therapy or suspected of needing therapy. The terms “individual” or “patient” are intended to be interchangeable with “subject.”

DETAILED DESCRIPTION

[044] Molecular monitoring in the dermal interstitial fluid (ISF) is an attractive approach to painlessly screen markers of health and disease status on the go. One promising strategy for accessing ISF involves the use of wearable patches containing microneedle sensor arrays. To date, such microneedle sensors have been fabricated via various manufacturing strategies based on injection molding, machining, and advanced lithography to name a few. 3D-printed microneedles have previously been reported as a convenient and scalable approach to sensor fabrication that, when combined with aptamer-based molecular measurements, can support continuous molecular monitoring in ISF. However, the previous platform suffered from poor patch stability when deployed on the skin of rodents in vivo. As disclosed herein, the inventors have identified that this problem was due to the rheological properties of, for example, the test animal (e.g., rodent) skin, which can contract post microneedle placement, physically pushing the microneedles out of the skin. This sensor retraction caused a loss of electrical contact between working and reference needles, irreversibly damaging the sensors. To address this problem in the present disclosure, we provide innovative approaches that allow magnetic placement of microneedle sensor arrays on the skin of live rodents or other subjects, affixing the patches under light pressure that prevents needle retraction. Using this strategy, we achieved sensor signaling baselines that drift at rates comparable to those seen with other in vivo deployments of electrochemical, aptamer-based sensors. To illustrate, we show real-time pharmacokinetic measurements in live Sprague-Dawley rats using SLA-printed, aptamer-functionalized microneedles and demonstrate their ability to support drift correction via kinetic differential measurements. These and other attributes of the present disclosure will be apparent upon a complete review of the specification, including the accompanying figures.

[045] To illustrate, in some aspects, the present disclosure provides electrochemical sensor devices. An electrochemical sensor device typically includes a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure. The plurality of microneedles includes at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode. The electrochemical sensor device also includes at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles in which the electrically conductive layer is operably connected, or connectable, to a detector. The electrochemical sensor device also includes a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer in which biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer. In addition, the electrochemical sensor device also includes a magnet connected, or connectable, to the second surface of the body structure, in which the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule. These and other elements or aspects of the electrochemical sensor devices of the present disclosure are further described and depicted in the Example below.

[046] In some embodiments, the electrochemical sensor device is a wearable device. The wearable device may be one wearable device or a plurality of wearable devices (e.g., comprising electrochemical sensor device that detect/monitor the same or different target molecules). The wearable device is typically configured to be wearable on or in proximity to the patient’s or subject’s body. For example, the wearable device may be attached by a strap or other means to a portion of the patient’s body such as to an arm, a leg, a waist, a neck, groin, etc. In alternative embodiments, the wearable device may be configured to be attached in proximity to a particular portion of the patient’s body. For example, the wearable device may be configured to be attached to clothes worn by a person. The wearable device may also be integrated into or attached to another device worn by a person. For example, the wearable device can be configured to attach to a watch, to a belt, to jewelry, to glasses, to undergarment, etc. In some embodiments, the electrochemical sensor device (e.g., configured as a wearable device, etc.) is packaged as a component of a kit.

[047] In some embodiments, the wearable device comprises a metallic structure. In some embodiments, the metallic structure is disposed at least proximal to an ear of the subject. In some embodiments, a metallic plate comprises the metallic structure, the metallic structure is implanted in the subject. In some embodiments, the metallic structure comprises a biocompatible plate fabricated from a medical grade, ferromagnetic steel. Typically, the metallic structure is configured to maintain a substantially constant and substantially homogeneous pressure across the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject. In some embodiments, the metallic structure is configured to maintain a substantially constant penetration depth of the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject. In some embodiments, the magnet comprises a permanent magnet. In some embodiments, wherein the magnet comprises a neodymium ring magnet. In some embodiments, the magnet comprises a cross- sectional dimension of about 25 mm or less. In some embodiments, the magnet comprises a thickness of about 5 mm or less.

[048] In some embodiments, the electrochemical sensor device is configured to generate one or more square wave or cyclic voltammograms using square wave voltammetry (SWV) and/or cyclic voltammetry (CV) and determine a change in peak-to-peak separation, EP.T, from the voltammograms to detect the target molecule when the biomolecular receptors bind the target molecule. In some embodiments, the electrochemical sensor device is configured to perform drift correction. In some embodiments, the drift correction is performed using kinetic differential measurements (KDM). In some embodiments, the electrochemical sensor device is configured to generate substantially stable baseline electrochemical signals. In some embodiments, a thickness of the electrically conductive layer is less than about 1 pm. In some embodiments, the body structure comprises a substantially circular cross-sectional shape, a substantially square cross-sectional shape, or a substantially rectangular cross-sectional shape.

[049] In some embodiments, the plurality of microneedles is arrayed in a substantially circular format, a substantially square format, or a substantially rectangular format. In some embodiments, at least one of the microneedles in the plurality of microneedles has a substantially circular, a substantially square, or a substantially rectangular cross-sectional shape. In some embodiments, at least of the microneedles in the plurality of microneedles has a substantially conical shape, a substantially pyramidal shape, or a substantially tapered shape. In some embodiments, edges of at least one pair of microneedles in the plurality of microneedles are separated from one another by about 5 mm or less. In some embodiments, the plurality of microneedles is fabricated integral with the body structure.

[050] In some embodiments, the electrically conductive layer comprises a chromium layer and/or a gold layer. In some embodiments, the electrically conductive layer of the reference electrode further comprises at least one silver layer. In some embodiments, the electrically conductive layer of the counter electrode further comprises at least one platinum layer. In some embodiments, at least one microneedle of the plurality of microneedles extends from the surface of the body structure by about 0.5 mm to about 2 mm. In some embodiments, a base portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 400 pm. In some embodiments, a tip portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 100 pm.

[051] In some embodiments, the detector comprises an electrochemical analyzer. In some embodiments, the electrochemical analyzer comprises a potentiostat. In some embodiments, the electrically conductive layer is configured to wirelessly connect to the detector. In some embodiments, the detector is configured to measure square wave voltammograms. In some embodiments, an electrically insulating cover structure is attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure.

[052] In some embodiments, the body structure and the plurality of microneedles are fabricated using a stereolithography (SLA) three-dimensional (3D) printing technique. In some embodiments, the plurality of microneedles comprises twelve microneedles, wherein seven of the microneedles are working electrodes in which two of the microneedles are reference electrodes, and in which three of the microneedles are counter electrodes. In some embodiments, the body structure comprises a cross-sectional dimension of about 30 mm or less.

[053] In some embodiments, the plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged aptamers. In some embodiments, the biomolecular receptor comprises an aptamer. In some embodiments, the aptamer comprises a 5’- thiol modification and a 3’ methylene blue modification. In some embodiments, the biomolecular receptor comprises a nucleic acid molecule. In some embodiments, the redox reporters comprise methylene blue (MB) or an osmium-based complex. In some embodiments, the target molecule comprises a therapeutic agent. In some embodiments, the therapeutic agent comprises an antibiotic. In some embodiments, the target molecule comprises a metabolite and/or an electrolyte. In some embodiments, the target molecule comprises a biomolecule.

[054] In some embodiments, the subject comprises a mammalian subject. In some embodiments, the mammalian subject comprises a human subject. In some embodiments, the mammalian subject is from an order Rodentia.

[055] To illustrate some of these aspects, FIG. 2A is a flow chart that schematically depicts steps in a method of detecting a target molecule in a subject according to an exemplary embodiment. As shown, method 200, includes positioning a plurality of microneedles of an electrochemical sensor device (as described further herein) in contact with a dermal interstitial fluid (ISF) of or in the subject (e.g., piercing the skin of an ear lobe of a human subject, etc.) (step 202). In some embodiments, the biomolecular receptor comprises an aptamer. Method 200 also includes detecting the electrochemical signals using the detector, thereby detecting the target molecule in the subject (step 204).

[056] In some embodiments, method 200 further includes implanting the metallic structure in the body of the subject prior to the positioning step (i.e., step 202). In some embodiments, method 200 includes implanting the metallic structure in an ear lobe of the subject prior to the positioning step (i.e., step 202). In some embodiments, method 200 includes substantially continuously monitoring the target molecule in the subject during a selected duration of time. In some embodiments, method 200 includes intermittently monitoring the target molecule in the subject during a selected duration of time. In some embodiments, method 200 includes performing a pharmacokinetic, pharmacodynamic, and/or toxicology assessment of the target molecule in the subject. In some embodiments, method 200 includes detecting the electrochemical signals in substantially real-time. In some embodiments, method 200 further includes performing drift correction of the electrochemical sensor device using one or more kinetic differential measurements (KDM). In some embodiments, the electrochemical sensor device further includes an electrically insulating cover structure attached to the first surface of the body structure. In these embodiments, the plurality of microneedles typically extends through the electrically insulating cover structure. In some embodiments, the plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged nucleic acids. In some embodiments, the redox reporters comprise methylene blue (MB). To further illustrate, FIG. 3 shows algorithm 400 that can be used to implement aspects of method 200 according to an exemplary embodiment.

[057] In some embodiments, method 200 includes comparing the AEP.T to a no target peak-to-peak separation, AEP.NT, determined from one or more cyclic voltammograms generated from the electrochemical sensor in the absence of the target molecule. In some embodiments, method 200 includes determining a concentration of the target molecule in the sample by comparing the AEP.T to a standard curve. In some embodiments, method 200 includes determining the AEP.T from at least a first cyclic voltammogram and at least a second cyclic voltammogram generated from the electrochemical sensor. In some embodiments, method 200 includes correlating at least two currents with corresponding peak potentials and calculating a separation between the peak potentials.

[058] In some embodiments, method 200 includes determining a concentration of the target molecule in the sample via the change in the target peak- to-peak separation, AEP.T. In some embodiments, the electrochemical sensor is substantially resistant to drift. In some embodiments, method 200 includes determining the change in the target peak-to-peak separation, AEP.T, from the cyclic voltammograms with about 900 milliseconds, about 800 milliseconds, about 700 milliseconds, about 600 milliseconds, about 500 milliseconds, about 400 milliseconds, about 300 milliseconds, about 200 milliseconds, about 100 milliseconds, or less of contacting the electrochemical sensor with the sample. In some embodiments, method 200 includes generating the cyclic voltammograms from the electrochemical sensor using a voltage scanning rate of about 5 V s’¹ or more. In some embodiments, the voltage scanning rate is between about 5 V s’¹ and about 10 V s’¹. In some embodiments, method 200 includes continuously monitoring the change in the target peak-to-peak separation, EP.T over time from multiple cyclic voltammograms generated from the electrochemical sensor.

[059] To further illustrate, FIG. 2B is a flow chart that schematically depicts steps in a method of method of producing an electrochemical sensor device according to an exemplary embodiment. As shown, method 210 includes forming a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure (step 212) and disposing at least one electrically conductive layer on at least a portion of the plurality of microneedles, wherein electrically conductive layer is operably connectable to a detector (step 214). Method 210 also includes attaching a plurality of biomolecular receptor-bound redox reporters to the electrically conductive layer in which biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters are configured to undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer (step 216). In addition, method 210 also includes connecting a magnet to the second surface of the body structure in which the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule (step 218). In some embodiments, method 210 further includes attaching an electrically insulating cover structure to the first surface of the body structure such that the plurality of microneedles extends through the electrically insulating cover structure.

[060] The present disclosure also provides various systems and computer program products or machine-readable media. In some aspects, for example, the methods described herein are optionally performed or facilitated at least in part using systems, distributed computing hardware and applications (e.g., cloud computing services), electronic communication networks, communication interfaces, computer program products, machine readable media, electronic storage media, software (e.g., machine-executable code or logic instructions) and/or the like. To illustrate, FIG. 4 provides a schematic diagram of an exemplary system suitable for use with implementing at least aspects of the methods disclosed in this application. As shown, system 600 includes at least one controller or computer, e.g., server 602 (e.g., a search engine server), which includes processor 604 and memory, storage device, or memory component 606, and one or more other communication devices 614, 616, (e.g., client-side computer terminals, telephones, tablets, laptops, other mobile devices, etc. (e.g., for receiving data for further analysis, etc.)) positioned remote from electrochemical sensor device 618, and in communication with the remote server 602, through electronic communication network 612, such as the Internet or other internetwork. Communication devices 614, 616 typically include an electronic display (e.g., an internet enabled computer or the like) in communication with, e.g., server 602 computer over network 612 in which the electronic display comprises a user interface (e.g., a graphical user interface (GUI), a web-based user interface, and/or the like) for displaying results upon implementing the methods described herein. In certain aspects, communication networks also encompass the physical transfer of data from one location to another, for example, using a hard drive, thumb drive, or other data storage mechanism. System 600 also includes program product 608 stored on a computer or machine readable medium, such as, for example, one or more of various types of memory, such as memory 606 of server 602, that is readable by the server 602, to facilitate, for example, a guided search application or other executable by one or more other communication devices, such as 614 (schematically shown as a desktop or personal computer). In some aspects, system 600 optionally also includes at least one database server, such as, for example, server 610 associated with an online website having data stored thereon searchable either directly or through search engine server 602. System 600 optionally also includes one or more other servers positioned remotely from server 602, each of which are optionally associated with one or more database servers 610 located remotely or located local to each of the other servers. The other servers can beneficially provide service to geographically remote users and enhance geographically distributed operations. [061] As understood by those of ordinary skill in the art, memory 606 of the server 602 optionally includes volatile and/or nonvolatile memory including, for example, RAM, ROM, and magnetic or optical disks, among others. It is also understood by those of ordinary skill in the art that although illustrated as a single server, the illustrated configuration of server 602 is given only by way of example and that other types of servers or computers configured according to various other methodologies or architectures can also be used. Server 602 shown schematically in FIG. 4, represents a server or server cluster or server farm and is not limited to any individual physical server. The server site may be deployed as a server farm or server cluster managed by a server hosting provider. The number of servers and their architecture and configuration may be increased based on usage, demand and capacity requirements for the system 600. As also understood by those of ordinary skill in the art, other user communication devices 614, 616 in these aspects, for example, can be a laptop, desktop, tablet, personal digital assistant (PDA), cell phone, server, or other types of computers. As known and understood by those of ordinary skill in the art, network 612 can include an internet, intranet, a telecommunication network, an extranet, or world wide web of a plurality of computers/servers in communication with one or more other computers through a communication network, and/or portions of a local or other area network.

[062] As further understood by those of ordinary skill in the art, exemplary program product or machine readable medium 608 is optionally in the form of microcode, programs, cloud computing format, routines, and/or symbolic languages that provide one or more sets of ordered operations that control the functioning of the hardware and direct its operation. Program product 608, according to an exemplary aspect, also need not reside in its entirety in volatile memory, but can be selectively loaded, as necessary, according to various methodologies as known and understood by those of ordinary skill in the art.

[063] As further understood by those of ordinary skill in the art, the term "computer-readable medium" or “machine-readable medium” refers to any medium that participates in providing instructions to a processor for execution. To illustrate, the term "computer-readable medium" or “machine-readable medium” encompasses distribution media, cloud computing formats, intermediate storage media, execution memory of a computer, and any other medium or device capable of storing program product 608 implementing the functionality or processes of various aspects of the present disclosure, for example, for reading by a computer. A "computer-readable medium" or “machine-readable medium” may take many forms, including but not limited to, non-volatile media, volatile media, and transmission media. Non-volatile media includes, for example, optical or magnetic disks. Volatile media includes dynamic memory, such as the main memory of a given system. Transmission media includes coaxial cables, copper wire and fiber optics, including the wires that comprise a bus. Transmission media can also take the form of acoustic or light waves, such as those generated during radio wave and infrared data communications, among others. Exemplary forms of computer-readable media include a floppy disk, a flexible disk, hard disk, magnetic tape, a flash drive, or any other magnetic medium, a CD-ROM, any other optical medium, punch cards, paper tape, any other physical medium with patterns of holes, a RAM, a PROM, and EPROM, a FLASH-EPROM, any other memory chip or cartridge, a carrier wave, or any other medium from which a computer can read.

[064] Program product 608 is optionally copied from the computer-readable medium to a hard disk or a similar intermediate storage medium. When program product 608, or portions thereof, are to be run, it is optionally loaded from their distribution medium, their intermediate storage medium, or the like into the execution memory of one or more computers, configuring the computer(s) to act in accordance with the functionality or method of various aspects. All such operations are well known to those of ordinary skill in the art of, for example, computer systems.

[065] To further illustrate, in certain aspects, this application provides systems that include one or more processors, and one or more memory components in communication with the processor. The memory component typically includes one or more instructions that, when executed, cause the processor to provide information that causes at least one result, data, and/or the like to be displayed or otherwise indicated (e.g., via a result indicator of electrochemical sensor device 618 and/or via communication devices 614, 616 or the like) and/or receive information from other system components and/or from a system user (e.g., via communication devices 614, 616, or the like).

[066] EXAMPLE 1 : 3D-Printed, Aptamer-based Microneedle Sensor Arrays Using Magnetic Placement on Live Rats for Pharmacokinetic Measurements in Interstitial Fluid

[067] 1. INTRODUCTION

[068] Microneedle (MN)-based electrochemical sensors are emerging as promising platforms for monitoring, minimally invasively, systemic biomarkers in the dermal interstitial fluid (ISF). With sharp tips down to few micrometers and length below one millimeter, MN-supported electrochemical sensors painlessly penetrate the epidermis to tap into the dermal ISF. Because this biofluid is highly irrigated by capillary plexuses and drained by lymphatic capillaries within the dermis, the dermal ISF represents a body compartment that dynamically equilibrates with, and therefore reflects, systemic molecular concentrations. Additionally, given that dermal ISF is available throughout the entire surface of the human body, dermis-implanted MN sensors represent an ideal tool for real-time, on-body biomarker monitoring. While early MN biosensing applications have focused on the continuous monitoring of metabolites and electrolytes, the recent use of bioaffinity receptors, particularly aptamers, has expanded the scope of such MN sensors towards a broader range of important target analytes.

[069] Significant efforts have been devoted over the past decade to the development of MN fabrication techniques that meet the size, shape and mechanical properties required for continuous molecular monitoring in dermal ISF. These techniques include etching, micromolding, UV lithography, drawing lithography, micromachining, stereolithography (SLA 3D printing), and injection molding. Among such methods, SLA 3D-printing is attractive because of its high resolution (~10 pm over the z-axis), ease of rapid prototyping, and versatility in the chemical composition of resins that can be employed for printing MNs. Early studies have demonstrated that 3D printing is suitable for the fabrication of microneedle sensing patches. For example, another group created a 3D-printed master mold to generate hollow MN patches via micromolding. The patches were later interfaced with a paper-based colorimetric assay for glucose in the back, allowing rapid extraction of ISF and visual determination of glycemic states. However, this MN fabrication strategy was limited by the fact that serial casting of MN patches from a master mold results in progressive deformation and, consequently, in gradual loss of resolution of the casted microneedles. In contrast, it is possible to employ SLA directly for rapid manufacture of MN masters. [070] In this example we describe an aptamer-based, 3D-printed microneedle sensor array for drift-corrected, real-time measurements of therapeutic drug pharmacokinetics in live rodents, highlighting a new magnetic device placement approach that ensures a highly stable baseline signal (Figure 5). Unlike benchmark micromolding strategies, in which a negative mold is fabricated for subsequent casting of the MN positive, here we directly 3D printed the MN arrays via SLA (Figure 5A). We demonstrate that the design of the resulting MNs can be optimized in terms of optimal geometries for in-vivo skin penetration and for post-fabrication coating with conductive gold films. The resulting gold-sputtered MN arrays allow facile functionalization with bioaffinity receptors, such as thiolated redox-tagged aptamers, via self-assembly of monolayers to create electrochemical biosensors for continuous drug monitoring at the MN tips (Figure 5B).

[071] In a previous study from our group, we reported significant loss of 3D- printed MN sensor signaling over time when the MNs were deployed on the skin of rodents. This signal loss was due to progressive retraction of the microneedles caused by the rheological properties of the rodent skin, which can contract after placing the device and push the microneedles outwardly. To overcome this issue, we demonstrate here a new device placement approach in which the microneedles are held on the skin via magnetic attraction. Specifically, we developed an easy surgical procedure to implant a metallic plate under the subcutaneous tissue of the rodents, while the microneedle platform contains a ring-shaped permanent magnet on the back (Figure 5A). Such new magnetically supported devices may offer significant benefits to biomedical research: (1 ) pre-measurement implantation of the subcutaneous plate would allow the preparation of large animals cohorts for molecular measurements in ISF; (2) after recovery from surgery, the microneedle devices could be placed and removed as needed, to allow multiday measurements in the same animals; (3) the magnetic attachment should be sturdy enough to allow continuous measurements in awake animals; and finally, (4) the magnetic skin attachment could facilitate translation to other model systems, and even to humans, if placement is intended in discrete skin regions, such as the ear lobes. Here, however, we focus on a proof-of-concept demonstration in anesthetized rodents to highlight the significant improvement in sensor baseline stability relative to common non-magnetic skin placement approaches that rely on medical adhesives or glue (Figure 5D vs C). Moreover, the enhanced signal stability associated with the magnetic placement enables the use, for the first time, of kinetic differential measurements (KDM), a strategy that uses dual square wave frequency measurements to correct for sensor-to-sensor manufacture variability and signaling drift in real time. Our results highlight the potential of SLA-based, 3D-printed, magnetically supported MN sensor arrays for continuous in-vivo pharmacokinetic studies, including considerable promise for improved drug monitoring in awake animals.

[072] 2. MATERIALS AND METHODS

[073] Chemicals and Materials. Gold cleaning solution (PN: 667978), 6- mercapto-1 -hexanol (MCH), and Tris(2-carboxyethyl) phosphine hydrochloride (TCEP) were purchased from Sigma-Aldrich (St. Louis, MO). 200 proof ethanol, sodium chloride (NaCI), sodium hydroxide (NaOH), and trace-metal grade sulfuric acid (H2SO4) were purchased from Fisher Scientific (Waltham, MA). Tobramycin sulfate was ordered from Spectrum Pharmacy Products (New Brunswick, NJ). N42 neodymium ring magnets (42NEG554012-NI) were purchased from Integrated Technologies Group (Culver City, CA). Neodynium magnetic discs (20 mm diameter, 3 mm thickness) were ordered from the MIN Cl store and Silver conductive epoxy adhesive was ordered from MG Chemicals store (Amazon). All aqueous solutions were prepared using deionized water from a Milli-Q® Direct purification system, with a resistivity of 18 MW. Photocurable resin (clear resin) was purchased from Formlabs (Somerville, MA), and (3D printing UV sensitive resin) from Anycubic (Brighton, Co). Silver ink was purchased from Ercon Inc (Wareham, MA). Deidentified human serum was purchased from BiolVT (Washington, D.C.). The oligonucleotide sequence for tobramycin measurements had dual modifications (5’ hexanethiol and 3’ methylene blue) and was purchased HPLC-purified from Sigma- Aldrich (Houston, TX).

[074] Microneedle Patch Fabrication: The MN patches were designed in SolidWorks software. MN arrays consisted of 12 MNs, with 7 designated as working electrodes, 2 as reference electrodes, and 3 as counter electrodes. For magnet placement, a void space was included on the back of MN patches, with dimensions: ID = 0.625 in, OD = 0.866 in, and thickness = 0.193 in. Design files in .stl format are provided as part of the Supporting Information. All patches were SLA printed using a Form3+ printer (formlabs™). After printing, the parts were washed inside an isopropanol ultrasonic bath for 60 min, then UV-cured for 60 min. As a final step, we outgassed the patches in a vacuum oven at 80°C overnight. The mask for sputtering was designed in Keynote (Mac OS) and fabricated on Kapton tape sheets using a Cricut Explore Air® 2 machine with a 60° deep replacement blade. The mask was placed on each of the 3D printed patches for metal deposition. We sputtered gold on the MN arrays using 65% rotation and 2.41 mTorr Ar gas pressure in a Denton Discovery 635 instrument. We first deposited a Cr adhesion layer for 6 mins at 300W, then the Au layer for 15 min at 100W, unless otherwise noted in the experimental discussion. To make electrical connections, conductive silver epoxy was carefully placed on connection channels, as shown in Figure 6B_V, and lined over the straight angle header of Dupont pins. We cured the silver epoxy at 80°C for 30 min. To create a pseudo-reference electrode, we added a layer of silver ink only on the 2 MNs located at the center of each MN array and cured it at 80°C for 30 min. SEM micrographs of the silver-coated pseudo-reference microneedles reveal a slight increase in microneedle body and tip diameter of 40 pm caused by the coating process. For final patch assembly, we pipetted in a thin layer of SLA resin at the base of the MNs, then inserted the cover piece to control microneedle height and achieve electrical insulation. We cured the resin in this last step using a UV-lamp (Peopoly 450 nm UV) for 2 h.

[075] Thickness Analysis of Sputtered Cr and Au. To determine the thicknesses of sputtered Cr and Au layers on the MNs, we employed glass slides as substrate analogs. The slides were cleaned with acetone, isopropyl alcohol, and deionized water, then dried under nitrogen gas. Afterwards, we applied the same 1/2” Kapton tape mask used on the MN patches. We sputtered with Cr and Au using various deposition protocols. After sputtering metals, we gently removed the Kapton tape. The thickness of the various metalic layers was then measured by comparing the glass slide surface baseline vs. the height of the sputtered layers using a stylus profiler (Dektak XT) with two-point cursors (200pm and 300pm). The cursors applied a force of 3 mg, and performed linear scans at a speed of 10 pm/s. We obtained a Cr layer of ~90 nm, ~200 nm and 350 nm with deposition times of 4, 6 and 12 min, respectively. For Au, we measured thicknesses of ~500 nm using a deposition time of 15 min, respectively. [076] Cell for Microneedle Cleaning and Surface Functionalization. We designed and 3D printed a chamber to allow cleaning with ethanol and functionalization of the MN surfaces with aptamer and alkylthiol monolayers To ensure the MNs were fully immersed in cleaning and deposition baths, we leveraged the magnetic attraction between the ring-shaped magnet on the back of the MNs, and a plate magnet placed underneath the chamber. The chamber contains an internal edge that is 4 mm taller than the bottom of the chamber which, coupled to the magnetic force, safely held the MNs in solution without risk of mechanical damage.

[077] Microneedle Cleaning Protocol. All MN arrays were treated prior to sensing monolayer depositions following a three-stage procedure. Stage 1: First, the MNs were incubated in pure ethanol for 10 min. Stage 2: An additional layer of gold was electroplated onto their surface from a gold chloride solution consisting of: 1 .2 mg/mL HAuCk, 1.5 wt% HCI, and 0.1 M NaCI, at 50° C. We electroplated via amperometry at -0.4 V with a pseudo-reference AgCI wire for 300 s. Stage 3: The electroplated devices were electrochemically cleaned via cyclic voltammetry in (I) 0.5 M NaOH and (II) 0.5 M H2SO4, following previously reported protocols (Arroyo- Curras et al., 2017a). Briefly, we first voltammetrically interrogated the MNs in 0.5 M NaOH, from -0.4 to -1.4 V vs Ag/AgCI, for 100 cycles at 4 V/s. Next, we voltammetrically interrogated the MNs in 0.5 M H2SO4, from 0.2 to -1.6 V vs Ag/AgCI at 4 V/s. We measured cyclic voltammograms in 0.05 M H2SO4 from 0.2 to -1 .6 V vs Ag/AgCI at 0.1 V/s to determine the final microscopic surface area of the as-cleaned MN arrays prior to monolayer depositions.

[078] Deposition of Sensing Monolayer. 1 pL of 100 pM thiolated MB- modified DNA with 1 pL of 100 mM TCEP were mixed for 1 h to reduce disulfide bonds. Afterwards, the aptamer solution was diluted to 200 nM using Tris buffer (20 mM Tris, 100 mM NaCI, 5 mM MgCl2, pH 7.4). The concentration of aptamer solutions was measured using UV-Vis spectroscopy employing an Implen Nanophotometer NP80 (Implen, Westlake Village, CA) prior to depositions. The MN sensor arrays were immersed in aptamer solution for 1 h in the incubation chamber, then incubated in a solution of 30 mM mercaptohexanol at 25° C for 3 h to form the final sensing monolayer. Surface concentration of aptamers was determined via the AUC of cyclic voltammograms performed at 100 mV/s. [079] Mechanical Durability of 3D-Printed MNs. The durability of gold- sputtered MNs was evaluated via full penetration of dissected porcine skin: 100, 500, and 1000 penetrations. We purchased porcine skin from a local supermarket and cleaned it by washing with deionized water. The porcine skin was fixed on a benchtop. Then, we pressed the MNs with force of 50 g/cm² (0.71 PSI) until they fully penetrated the skin. Then, we retracted the MNs and exposed the wounds to cyanine 7, a dye that reacts with free amines in the biomolecules present at the exposed dermis. SEM images of MN arrays before and after repeated skin penetration showed minor changes in the MN geometry or metal film homogeneity, and indicated that the 3D-printed microneedles are robust enough for skin insertion. Additionally, we electrochemicallly evaluated the active microscopic surface area of the microneedles after each penetration. For this, we determined double-layer capacitances (Cdi) by cyclic voltammetry at different scan rates (50, 100,1 50, 200, 250, and 300 mV/s) using a potential window from 0 V to 0.3 V after each insertion condition (0, 1 , 100, 500, and 1000 insertions). By using the equations: Cdl = Aj / 2v and Aj = ja - jc, where ja is the anodic and jc the cathodic current densities, we built current density vs scan rate plots in which the slope corresponds to Cdi. To build the plots, we sampled the current at 0.15 V across the different scan rates. The slopes were Cdi,o = 12.38 pF, Cdi = 10.20 pF, Cdi.wo = 9.22 pF, Cdi.soo = 8.16 pF, and Cdi.woo = 8.11 pF. These measurements indicated the microneedle surface area decreased following skin insertion by 17.60% after initial placement, to 34.49% after 1 ,000 insertions. However, we note these measurements were conducted using bare gold microneedles. An analogous experiment performed in vivo in rats using aptamer- coated microneedle arrays showed no signficant changes in surface aptamer concentration across ten insertions.

[080] Electrochemical Measurements. For benchtop electrochemical treatments and measurements, we employed 8-Channel Electrochemical Analyzers (CHI 1040C, Austin, TX). For in vivo measurements, we employed a hand-held potentiostat (CHI 1242, Austin, TX). All raw measurements were controlled via commercial CHI software. All in vitro measurements were carried out in a three- electrode cell configuration consisting of gold working, platinum wire counter, and Ag/AgCI (saturated KCI) reference electrodes. Cyclic voltammogram was recorded at a scan rate of 100 mV/s for evaluating the cleanliness of the gold surface of microneedles. Square wave voltammetry was performed using a square-wave amplitude of 25mV, step size of 1 mV, and various frequencies.

[081] In vivo Skin Penetration and ISF Drug Measurements. All in-vivo measurements were conducted on male Sprague-Dawley rats (300 - 400 g weight), in accordance to the JHU ACUC-approved protocol number RA22M242. Given the pilot nature of this study, a priori power analysis of animal numbers was not considered. However, the results presented here will inform power analyses for future in-vivo protocols with the goal of establishing accuracy of molecular transients and for technology validation. Rats were anesthetized using isoflurane gas inhalation (induction at 2.5-3%, maintenance at 2%) and monitored throughout the experiment using a pulse oximeter (Starr Life Sciences, Oakmont, PA) to measure heart rate, temperature, and %SpO2. We shaved the hair located around the abdomen area and top back area of rats to complete skin penetration tests using tapered and conical MNs. We photographed the penetration marks using an Apple iPhone 11 , without any additional photo editing changes. We have reported earlier successful penetration into the skin dermis of rats for the visual tracks shown in Figure 7 in a previous publication (Wu et al., 2022). For subcutaneous placement of the magnetic plate, we made a lateral skin incision adjacent to the abdominal area of about 1 cm in length. We then slid a round magnetic plate under the subcutaneous tissue, to a final location exactly at the center of the abdominal area of the rats. We closed the wound using hydrocolloid gel bandages. The magnet-supported MN patches were then placed above the metallic plate and softly snapped into place. The magnetic pressure between the plate and microneedle array was 50 g/cm² (0.71 PSI) as measured via a flexible pressure sensor from Qinlorgo (purchased via Amazon). We waited 30 min before serially measuring square wave voltammograms at various frequencies (i.e., 30, 80, 150, 240, 300, 400, and 600 Hz). We adjusted the data collection to achieve a rate of one voltammogram every ~1.5 min at any given frequency. After obtaining a one to two hour-long baseline (waiting for sensor baseline to reach a steady state), we performed an intravenous bolus of 20 mg/kg tobramycin via the tail vein. After completing molecular measurements for 2 h, all animals were euthanized via anesthesia overdose followed by decapitation, in compliance with American Veterinary Medical Association (AVMA) guidelines.

[082] Data Analysis. We employed a previously reported, open-source Python script called SACMES for the batch processing of our electrochemical measurements (Curtis et al., 2019). SACMES allows us to extract specific capacitive and faradaic currents from our electrochemical measurements, including peak currents, in real time. SACMES also supports the kinetic differential measurements (KDM) employed in this work for drift correction, as illustrated in Figure 9A. MN sensor calibrations were performed in a 1 :1 solution of serum with phosphate- buffered saline complemented with 2 mM MgCl2 as a proxy for ISF, as shown in Figure 9B.

[083] 3. RESULTS AND DISCUSSION

[084] We fabricated our MN sensor arrays using an SLA 3D printer from Formlabs (Model: Form3+), which has a z-axis resolution of 25 pm. The blueprint design of the microneedles consisted of twelve MNs, seven acting as working electrodes, three as counter electrodes, and two as the reference electrode. In the resulting 3D-printed array, the MNs were ~3mm long before device assembly (Figure 6A). After curing and washing the arrays, we sequentially applied Kapton tape masks to deposit metallic layers on the MNs (Figure 6 Bi-Biv). We first deposited a chromium adhesion layer on all the microneedles (~200 nm), followed by the sputtering of a gold layer to support aptamer self-assembly (~500 nm, see Methods for additional details). We coated the two reference electrode needles with silver ink, and then incubated them in pure bleach (using a microdroplet from a 26- gauge needle) for 20 min to generate the AgCI film prior to in vivo experiments. For final assembly, we made connections to the MNs using Dupont pins (Figure 6B_V), insulated them with a top 3D-printed lid delimiting the final dimensions of the MNs (1.5 mm long, 389 pm-wide, ~65 pm for the tip diameter, and mounted a magnetic neodymium ring (grade N42) on their back (Figure 6Bvi. Figure 6C shows front and back of the final device assembly ready for in vivo measurements.

[085] Effect of MN geometry on skin penetration. We initially evaluated two MN array geometries to determine the best strategy to access the dermal ISF in rats (Figure 7). The first design was round, with tapered MNs (as shown in Figure 6A) located at the center of the patch, with needle-to-needle lateral spacings of ~ 2 mm (Figure 7Ai, iv), ~ 2.2 mm (Figure 7AH, _V), and ~ 2.5 mm (Figure 7Am, _v . We evaluated two skin placement locations for our devices: the rat abdomen and the back, over the scapula. In this test, performed on one animal per design, our tapered microneedle devices were able to successfully penetrate the abdomen skin, regardless of lateral spacing, as revealed by hollow markings in Figure 7Ai-Aiu. In contrast, penetration of the back skin was not always successful, with fewer MN tracks appearing on the skin, and in average appearing as more superficial (Figure 7Aiv-Avi). These results indicate that the device placement on the rat’s back may require sharper needles than the ones employed in this work. In average, the applied pressure in each test was 50 g/cm² (0.71 PSI).

[086] The second design was rectangular, with conical MNs located at the center of the patch (Figure 7B), with needle-to-needle lateral spacings ~ 0.8 mm (Figure 7Bi, iv), ~ 0.5 mm (Figure 7BH, _v), and ~ 1.0 mm (Figure 7Bm, _v . These devices had the counter and reference electrodes separated from the working electrode array. Like the circular devices, the rectangular arrays performed better on the abdomen skin relative to the back. However, the conical microneedles had inferior skin penetration relative to the tapered microneedles. We believe this inferior performance originated because it is easier to apply equal pressure across circular devices vs rectangular ones, which tend to tilt sideways upon skin placement. Given these observations, we decided to continue to use the tapered MNs with the circular base, as shown in Figures 6C and 7A. All subsequent measurements were performed on the abdomen skin of rats.

[087] Microtip Stability as a Function of Chromium and Gold Layer Thicknesses. To support electrochemical aptamer-based (E-AB) sensing, the MNs must be coated with a homogeneous gold film. Additionally, the film must have high purity to allow dense packing of aptamer and blocking alkylthiol monolayer elements. Finally, the film must be robust enough to tolerate chemical cleaning in ethanol and electrochemical cleaning in dilute sodium hydroxide and sulfuric acid, as required prior to monolayer deposition (see Methods section). Thus, a critical aspect of MN patch prototyping was tuning the chromium adhesion layer and gold layer thicknesses to tolerate cleaning steps prior to sensor fabrication. We evaluated both tapered and conical MN designs to determine the most suitable for cleaning and posterior aptamer and alkanethiol monolayer formation. For this purpose, we tuned the deposition time of each metal layer to increase thickness, with the goal of identifying the best conditions for sensor fabrication. We empirically determined that tapered MNs achieved better adhesion of metal layers than conical MNs, as indicated by our ability to successfully deposit aptamer and alkanethiol monolayers. At this stage we are not sure why the conical microneedles did not achieve stable metal coatings that could tolerate chemical and electrochemical cleaning; however, we speculate this is a 3D printing issue related to roughness in the plastic caused by the progressively decreasing radius of the conical features. Additionally, through testing different metal deposition times, we determined that a 200 nm-thick chromium adhesion layer and a 500 nm-thick gold layer tolerated best the chemical and electrochemical cleaning procedures. Finally, these films were conductive enough to allow homogeneous electrodeposition of gold to increase the microscopic surface area of the microneedles and improve the signal-to-noise of E-AB measurements.

[088] Effect of magnetic placement on sensor baseline. In our previously reported work, we held microneedle sensor arrays on the skin of rodents via adhesive tape (Figure 8A). This strategy was not robust (i.e. , the devices tended to lose electrical contact). As an example, typical in-vivo sensor performance from this strategy is shown in Figure 8A, center. We observed linear sensor signaling decay using all square wave frequencies tested over a 20 min period. We believe the inflexion point seen in the data at 20 min indicates that the needles are no longer in the dermis, but rather superficially resting at the top of the skin. Individual analysis of the square-wave voltammograms (Figure 8A, right) revealed a clear disappearance of the methylene blue reporter redox peak after 1 h of serial sensor interrogation.

[089] To overcome the on-skin stability issue of our platform, we tested two additional strategies. The first was to deploy super glue around the edges of the devices, carefully enough to not contact the microneedle sensors with the glue (Figure 8B). This strategy slowed down the microneedle retraction from the skin (Figure 8B, center), but did not prevent microneedle retraction over the measurement period. This is presumably because of the difficulty of applying glue to the microneedle surroundings without damaging the sensor, which leaves glue-free islands around each individual microneedle that are big enough to let the skin contract, pushing the microneedles outwardly. Although this strategy retained more of the original sensor signal relative to taped devices (Figure 8B, right), the signal decay was still significant, close to 80%.

[090] The second strategy we tested involved magnetic placement of the microneedle sensor arrays via a subcutaneous metallic plate. To evaluate this strategy, we performed a small skin incision to the left of the rat’s abdominal area (shown in Figure 8C). Through this small incision (~1 cm wide), we placed a flat metallic disc with round edges subcutaneously. The flat disc plate served two functions: (1 ) it held the pressure constant and homogeneous across the microneedle array, and (2) it maintained the skin stretched under the microneedle array to ensure constant microneedle penetration depth. Using this strategy, we achieved flat sensor responses (Figure 8C, center) that decayed at rates that are comparable to those seen in other in vivo implementations of electrochemical, aptamer-based sensors. Individual comparison of voltammograms after 1 h of serial sensor interrogation in vivo showed minimal decay of the reporter signal (Figure 8C). We note that drift in these sensors is square wave frequency dependent. Some frequencies are more sensitive to electron transfer from methylene blue-modified aptamers than others. Thus, the extent to which the sensor signal decays at any given frequency is typically not the same at other frequencies, as seen for the glued and magnetically attached sensors. However, we observed less dispersion across frequencies for the skin-glued devices because sensor disconnection seemed to occur faster, and more evenly, than with the other two strategies.

[091] In this example, we used a neodymium magnet for the subcutaneous plate; however, since this material is cytotoxic for chronic deployments. Future use of this magnetic placement approach will rely on biocompatible plates made of medical grade ferromagnetic steels commonly used in medical devices, or magnetic soft materials with rheological properties that better match those of the skin. Although a comprehensive evaluation of steel types, magnetic materials, and long-term biocompatibility studies are needed, these measurements fall beyond the scope of the current work and will be pursued in the future. Overall, the highly stable sensor baselines achieved with this magnetic strategy pave the way for continuous molecular measurements in rodent cohorts, both in anesthetized and awake animals. A second, independent repetition of the measurements shown in Figure 8 was performed to illustrate the reproducibility of the results based on the different MN placement strategies. Finally, the sensor baseline stabilizing effect of the magnetic approach is not limited to one hour-long measurements. We repeated three additional measurements to illustrate that, when correcting for natural E-AB sensor drift as discussed in Figure 9 below, the magnetic approach allows reproducible measurements of flat sensor baselines over periods of hours.

[092] Pharmacokinetic measurements in dermal ISF following intravenous bolus. To demonstrate that our magnetic placement strategy supports drift-free molecular measurements, we performed continuous monitoring of tobramycin in ISF following intravenous dosing via the tail vein. We carried out these measurements in three separate animals, using three independently fabricated microneedle sensor arrays, delivering an intravenous bolus of 20 mg/kg. A previous publication on microneedle-supported electrochemical aptamer-based sensors reported single square wave frequency measurements and corrected for drift by empirically constructing a baseline function based on an exponential equation. However, sensor drift in vivo can be best corrected via kinetic differential measurements (KDM), an approach that simultaneously collects sensor data at low and high square wave frequencies. The low frequency measurements are subtracted in real time from the high frequency data, correcting the drift while simultaneously amplifying the sensor signals. This approach has been discussed once for microneedle aptamer-based sensors in the context of vancomycin ISF measurements ex vivo. Here, in contrast, we demonstrate the first implementation of this drift correction strategy in vivo.

[093] For our in vivo measurements, we empirically determined that square wave voltammetric measurements collected at 300 Hz and 30 Hz presented identical drift in vivo across the three different devices employed (black and gray data in Figure 9A). Subtraction of the low frequency data from the high frequency measurements produced KDM-corrected, drift-free pharmacokinetic measurements (lower trace in Figure 9A). Calibrating our sensors at the same frequencies in 50% dilute serum solutions ex vivo, we built dose-response curves (Figure 9B). By performing non-regression analysis of such curves using the Hill isotherm, we determined a best-fit curve that was used to translate sensor signaling output to drug concentration in ISF. Using this calibration, we measured tobramycin ISF levels in three independent animals (Figure 9C-9E). A negative control experiment, in which we injected phosphate buffered saline without tobramycin, did not produce a positive response to the drug bolus. Instead, it showed negative drift over time, potentially due to changes in hydration of the animal skin. [094] We made two important observations in the pharmacokinetic profiles obtained: tobramycin disposition into the dermal ISF and excretion do not follow a conventional two compartment pharmacokinetic model with rapid, first order absorption and slow, first order elimination kinetics. Instead, the slow absorption phase in our profiles (increasing concentration rate in Figure 9C-9E) indicates that tobramycin transport into the dermal ISF is better explained via flip-flop pharmacokinetics. This pharmacokinetic model assumes that the rate of excretion is faster than, or as fast as, the rate of absorption into the dermal ISF. Thus, the drug absorption is the rate-limiting step, as opposed to drug elimination. Fit parameters after individual curve-fitting are shown in Table 1 with <10% error in the estimation of kinetic rates. The data consistently show absorption rates, bs, that are greater than or equal to the excretion rates, taiim, when the precision of these measurements is considered. These results look like those from the only other previous report based on acupuncture microneedle-based aptamer sensors. However, neither flip-flop kinetics nor specific values for kinetic rate constants were discussed in that prior publication. Further investigation of how such kinetics reflect physiological processes will be the subject of future research.

Table 1. Pharmacokinetic parameters derived from Figure 9

[095] The second important observation is that, a though we administered the same body weight-corrected doses, tobramycin ISF levels varied across the three independent animals. We currently attribute this variability to (1 ) tail vein dosing errors and (2) variability in hydration status of the animals. It is also possible that the observed differences in pharmacokinetic parameters are due to genetic variations across individual subjects, since the Sprague-Dawley rats employed in this work are an outbred strain subject to natural genetic mutations. We will devote future efforts to studying these confounding effects on our pharmacokinetic measurements.

[096] 4. CONCLUSIONS

[097] This example reports the development of MN aptamer-based sensor arrays that were directly printed via SLA, metal coated via sputtering and electroplating, functionalized with alkylthiol- and redox reported-modified aptamers, and deployed in vivo on the skin of rodents via an innovative magnetic approach. The magnetic approach consists of subcutaneously implanting a metallic plate to allow stable affixing of a magnet-containing MN sensor patch on the skin of rodents. The magnetic force applied between the plate and the MN sensor patch can be adjusted by tuning the thickness of the patch, with thicker patches decreasing magnetic attraction and, therefore, pressure on the skin. This innovative sensor deployment approach paves the way for future measurements in awake animals using chronic paradigms, in which measurements could be repetitively performed on the same animals by placing and removing the microneedles on demand. The microneedle sensors tolerate repeated skin reinsertions without significant loss of aptamers or general damage to the biosensor interface. Additionally, the magnetic strategy could enable sensing applications in humans when discreet molecular monitoring is warranted such as, for example, on the ear lobes.

[098] The wearable MN patches reported here achieved two critical innovations relative to previously published in-vivo MN-based platforms. First, the magnetic skin placement approach allowed the measurement of stable E-AB signal baselines, with drift that is on par with other previously published in vivo deployments of E-AB sensors. Second, the stable signaling baselines we achieved supported, for the first time with in-vivo MN sensors arrays, drift-corrected molecular tracking in the dermal ISF using kinetic differential measurements (KDM). This strategy relies on dual square wave frequency measurements, collected serially on the devices and in real time to correct for sensor-to-sensor manufacturing variability, amplify sensor signal gain in response to increasing target concentration, and correct drift to achieve accurate pharmacokinetic measurements. Leveraging these innovations, we demonstrated continuous molecular measurements of the therapeutic tobramycin in the dermal ISF of three independent animals following intravenous administration. The resulting pharmacokinetic measurements revealed processes of drug absorption into and excretion from dermal ISF that are best described by a model of Flip-Flop pharmacokinetics, in which drug excretion is faster or as fast as drug absorption. Future research across statistically powered animal studies should shed further insights into the implications of such drug transport processes on molecular levels in dermal ISF and their implications in health and disease monitoring.

[099] Our MN sensor patches are not without limitations in their current format. The resolution of the 3D-printed MN tips is sufficiently sharp to penetrate the abdominal skin of rats but not to reproducibly penetrate the thicker skin on the animals’ back. To address this issue in future work, we have acquired a new 3D printer with better z resolution relative to the Form Labs printer used in this work. Preliminary prototypes have achieved tapered tip diameters of <20 pm. A second limitation in the current design arises from the fact that the entire body of the MNs is functionalized with aptamers, therefore providing molecular measurements averaged across the length of the MNs. We will devote future efforts to exposing only the tapered end of the MNs for achieving molecular measurements truly localized in the dermis. Finally, future measurements should address heterogeneity in drug dosing via automated tail vein drug dosing, and control over hydration level to ensure the pharmacokinetic variation observed in our measurements does not arise from variations in personnel or experimental setups.

[0100] EXAMPLE 2: Aptamer-Based Therapeutic Drug Monitoring in the Dermis Interstitial Fluid

[0101] 1. INTRODUCTION

[0102] The dermal interstitium is a fluid-filled elastic space that has been highlighted as an ideal body compartment for diagnostic development. This is because the interstitial fluid (ISF) contained within the interstitium is rich in molecular diversity, reflecting both skin-local and systemic compounds that could be leveraged as health and disease biomarkers. The potential value of molecules in ISF for diagnostic applications has sparked intense biosensor development efforts focused not just on sampling ISF, but also on directly interrogating it via wearable biosensor platforms. To date, ISF-probing sensors have been developed for the tracking of metabolites such as lactate, glucose and ketone bodies. Additionally, our group and others have highlighted the potential value of ISF-based measurements for therapeutic drug monitoring (TDM). In TDM, therapeutic dosing is based on serially measured plasma drug concentrations, typically performed via blood draws followed by in vitro quantification. However, the availability of skin-interfaced wearable monitors could significantly shorten the turnaround time of TDM, from hours and sometimes days to instantaneous, real-time drug tracking in ISF. This ability, in turn, should improve patient outcomes by precisely modulating therapeutic dosing to achieve the right in vivo concentration at the right time, while simultaneously minimizing drug toxicity.

[0103] To enable real-time TDM, microneedle sensor patches with built-in electrochemical, aptamer-based (E-AB) sensors have been previously reported for continuous molecular monitoring in ISF. These platforms consist of microneedle or microtip electrode arrays that can pierce the epidermis to then sit in the dermis, where they can interact with ISF. The electrodes can be fabricated at scale via various methods such as 3D printing, polymer casting, and micromolding, metal- coated via sputtering, and later integrated with wearable electronic boards for wireless interrogation and data transmission. To achieve highly specific molecular sensing, the electrodes are functionalized with structure-switching nucleic acid aptamers, which can reversibly bind and dissociate from their targets at sub-second scales. Target interactions with the aptamers change the electron transfer rate of aptamer-bound redox reporters, a phenomenon that can be probed in real time via electrochemistry. As promising as this platform is, to date microneedle sensor arrays have been limited to the sensing of only two therapeutic agents, tobramycin and vancomycin. Both of these targets are soluble in aqueous media at the doses typically delivered in vivo [their predicted solubilities are 1000 g/L (~ 2.14 M) for tobramycin and 0.4 g L’¹ (~276 pM) for vancomycin]. Additionally, tobramycin has negligible protein binding in plasma while vancomycin is ~ 50% protein bound, leaving a significant protein-free fraction that can diffusively cross capillary barriers to go from blood to ISF. However, an open question remains as to whether microneedle sensors could be used for ISF-based TDM of more hydrophobic targets such as chemotherapeutics used in cancer treatment, which tend to be >60% protein bound in plasma.

[0104] This example sought to expand our general understanding of molecular transport from blood to dermal ISF for small molecule therapeutics and to evaluate the value of microneedle sensor arrays for the TDM of agents that mainly exist bound to proteins in plasma. To complete our study, we employed a microneedle platform that blended two previously published strategies, acupuncture microneedles and magnetically-attached microneedle sensors, to create a novel multichannel sensor allowing five redundant, real-time in vivo measurements in rodents. We functionalized the devices with aptamers binding the two antibiotics vancomycin and tobramycin (as positive controls), the two chemotherapeutic drugs irinotecan and doxorubicin, and the amino acid phenylalanine (relevant to metabolic disorders such as phenylketonuria). Using the resulting aptamer-based sensors we demonstrate that targets that exist at considerable concentrations in their free form in plasma like tobramycin and vancomycin readily transport to the dermal ISF and can be monitored via microneedle aptamer-based sensors. However, the targets irinotecan and doxorubicin [predicted solubilities of 0.1 g L’¹ (~170 pM) and 0.41 g L- 1 (~750 pM)], which are >60% protein-bound in plasma, fall below the limit of detection of current aptamer-based sensors in ISF, even following high intravenous doses. Additionally, we discuss the special case of phenylalanine [predicted solubility for phenylalanine is 41 g L’¹ (~248 mM)], which is known to be bound to proteins in plasma but was successfully monitored by our sensors in ISF, opening a potential opportunity to develop wearable monitors for individuals with phenylketonuria.

[0105] 2. RESULTS

[0106] The microneedle sensor platform we employed in this study (Figure 10) consisted of two main components: a stereolithographically printed shell containing microneedle insertion holes, and commercially purchased acupuncture microneedles. The shell included twenty four concentric holes (700 pm in diameter) dedicated to five working electrode channels and one counter electrode. At the center, the shell also included five holes for the reference electrode (800 pm in diameter). Each of the working and counter electrode channels consisted of four gold microneedles electrochemically coated with gold nanoparticles (AuNPs, Figure 10A). The reference electrode consisted of five silver paint-coated microneedles, which were bleached post assembly to form a Ag|AgCI pseudo-reference. On the back end of the device, the four microneedles of each channel were shorted together and connected to electrically insulated copper wires, which were linked to a potentiostat for electrochemical interrogation. The four shorted microneedles per channel allowed the measurement of larger baseline currents without compromising the microscopic nature of the electrodes, significantly improving signal-to-noise in our in vivo measurements. Additionally, the five working electrode channels provided rigorous statistical redundancy during each in vivo measurement, eliminating the need to interrogate a large number of animals.

[0107] To functionalize the sensor platform with aptamers, we included in the 3D design a 400 pm-thick breakable wall that created a well to hold monolayer solutions. We first incubated the microneedles with any one of the thiol- and reporter- modified aptamer sequences, then backfilled the resulting monolayer with mercaptohexanol, as is standard in the field. All aptamers employed were previously engineered to undergo target-binding induced changes in electron transfer from the reporter attached to their 3’ end (i.e., methylene blue), which can be interrogated in serial mode via square wave voltammetry (Figure 10B). Details regarding the microneedle cleaning procedures employed prior to aptamer deposition are described in detail in the Materials and Methods section.

[0108] The microneedle electrodes were electrochemically coated with AuNPs to minimize exposure of their nickel core to buffer or biofluids, and to increase the loading capacity of aptamers on their surface. Scanning electron micrographs (SEM) confirmed the presence of uniform AuNPs coatings (inset in Figure 10A). Relative to bare electrodes, the AuNPs-coated microneedles exhibited a significant increase in gold abundance under energy-dispersive spectroscopy (EDS) and in microscopic surface area as measured via cyclic voltammetry. Additionally, square wave voltammograms of the aptamer-functionalized surfaces revealed a 10-fold increase in reporter reduction currents and lower capacitive currents. Overall, these combined results demonstrate that the electrochemical deposition of AuNPs on the microneedle electrodes successfully produced homogeneous gold surfaces and allowed functionalization of the microneedle electrodes with thiol- and reporter-modified aptamers and blocking monolayers (Figure 10B).

[0109] To deploy our microneedle platform in vivo we targeted the abdominal skin of rats (Figure 10C and 11 A). The back of the platform contains a ring-shaped permanent magnet that allows magnetic attachment to the abdominal skin via a subcutaneously implanted magnetic plate (Figure 10D). Prior to skin deployment, the breakable walls used for aptamer functionalization are removed, leaving any microscopic debris inside a circular channel that separates the wall from the microneedles (gray ring in Figure 10A). This debris chamber minimizes the risk of sensor damage when removing the walls. Finally, air can be blown over the microneedles to remove any debris and the devices can be placed on the animals, connected to a potentiostat, and serially interrogated to support real-time molecular monitoring in the dermal ISF (Figure 10E). The in vivo protocol used in this study involved intravenous (I.V.) delivery of drug boluses (Figure 10C), which was accomplished by cannulating the right jugular vein of the rodents. A photograph of the final setup displaying the connected microneedle platform and the cannulated vein is shown in Figure 11 A. Our measurement protocol (Figure 11 B) following device placement consisted of serially interrogating the sensors every 120 s for 30 min to establish a baseline. This period covers the acute inflammation phase post microneedle insertion and allows for sensor baseline stabilization. We then delivered the target dose of any of the five drugs considered in this study (Figure 11 C), and continued recording with the same frequency for 2.5 h. At the end of this period, the animals were euthanized, and skin samples were dissected to confirm microneedle placement in the dermis via histological analysis.

[0110] We first used our microneedle sensor platform to monitor vancomycin transport from blood to the dermis ISF. Vancomycin is a glycopeptide antibiotic with a narrow therapeutic window and for which TDM is standard of care. For these measurements, we employed a vancomycin-binding aptamer sequence previously optimized by our laboratory. After sensor fabrication, we identified optimal square wave voltammetry parameters for sensor interrogation by building square wave frequency maps in 50% rat serum in phosphate-buffered saline. Measurements at square wave frequencies of 150 Hz and 15 Hz provided the largest ON and OFF sensor responses, respectively, while also allowing the in vivo implementation of kinetic differential measurements (KDM)- a drift correction and signal amplification strategy. After the baseline signal stabilization period (Figure 11 B, first 30 min), we dosed three independent rats with intravenous vancomycin boluses of 20 mg/kg, 40 mg/kg, and 83 mg/kg, respectively. In all cases, we observed an immediate rise in signal output (within a few minutes post dosing, Figure 12A), corresponding to drug absorption into the dermis ISF. The excretion phase for the lowest dose (black trace, Figure 12C) had a time constant 2.68 ± 1.04 min’¹, which relative to the absorption time constant of 2.68 ± 1.09 min’¹, reflects flip-flop pharmacokinetics. The implication of flip-flop behavior is that the residence time of vancomycin in the ISF is absorptionlimited in rats, a factor that may be animal model specific and not directly translatable to TDM in human patients but should be excluded in any pilot clinical studies for ISF-based TDM.

[0111] The subject-to-subject variability in pharmacokinetics observed with our platform falls within statistical error of the measurements. To demonstrate this, we repeated the intravenous vancomycin bolus of 40 mg/kg into a fourth independent animal. Doing so we observed a delay of a few minutes in the onset of the absorption phases between the two comparative measurements, but the plateau concentrations and pharmacokinetic curves looked otherwise identical (Figure 12A, red squares vs triangles). The delay in absorption onset is not surprising given potential differences in fat content between the animals (their experimental weights were 320 g to 400 g, or ~20%), which can affect the response time of our ISF-based measurements (see, for example, response onset variability in Figure 12A). To convert the sensorgrams in Figure 12A to pharmacokinetic profiles (i.e., [vancomycin]isF vs, time), we built a dose-response curve in 50% rat serum in phosphate-buffered saline as a proxy to ISF (Figure 12B). Fitting the curve to the Hill equation (trace) allowed us to extract the KDM-based ECso and gain of the sensors, which we could then use to translate the sensorgrams into drug concentrations over time (Figure 12C-F). This process allowed us to observe that vancomycin can readily transfer from blood to ISF regardless of dose delivered, without any obvious transport limitations. In fact, the concentrations measured in ISF reflect previously measured plasma concentrations following a similar intravenous dose. Additionally, at intravenous doses of 40 mg/kg or higher, the vancomycin concentration reached in ISF is so high the sensor signal saturates. For example, the signal scatter observed in Figure 12E, D reflects concentrations above 90% sensor gain based on our dose-response curve (Figure 12B). Additionally, it is interesting to note that the curves smoothen out at the end of the measurement period because ISF vancomycin starts to decrease via excretion, bringing the measurements back to below sensor saturating concentrations. Finally, a dose of 83 mg/kg grossly saturates the sensor; thus, the flat data in Figure 12F reflects our inability to quantify vancomycin concentrations above 4 pM.

[0112] We next tested the ability of our microneedle sensors to monitor the blood-to-ISF transport of tobramycin, an aminoglycoside antibiotic that also has a narrow therapeutic window. Following the same process, we determined the best square wave interrogation frequencies by constructing frequency maps. Based on these measurements, 300 Hz and 30 Hz were determined as optimal parameters for in vivo KDM-based quantification and drift correction. We recorded a 30 min baseline on two independent animals, after which we dosed intravenous tobramycin boluses of 28 mg/kg and 40 mg/kg, respectively (Figure 13A). During both experiments, we again observed sensor responses within a few minutes post drug bolus administration. However, the two doses created statistically distinct profiles. While the lower dose generated a pharmacokinetic profile characteristic of first order absorption and first order elimination kinetics (Figure 13A, black), the larger dose reached a steady plateau that lasted approximately 60 min before excretion kinetics became visible in the profile (Figure 13A, lighter, upper trace). By building a doseresponse curve in our ISF proxy (Figure 13B), we confirmed that the plateau (at ~25% gain) was indeed not reflecting sensor saturation. Instead, we speculate the signal saturation observed in this case could reflect an active process of drug excretion (e.g., via an efflux pump), with the rate of excretion matching the rate of absorption. Then, when absorption stops, the ISF tobramycin levels begin to reflect systemic excretion. A final observation is that the ISF concentrations measured for tobramycin (Figure 13C-D), which agree in magnitude with previously reported plasma levels following a similar dose, are much higher than those measured for vancomycin (Figure 12C-E), potentially reflecting differences in free drug availability in plasma between the two therapeutics.

[0113] As final controls to demonstrate that sensor responses are due to the drugs themselves and not to changes in, for example, dermis hydration following intravenous dosing, we repeated measurements in two additional animals with the same vehicle boluses shown in Figure 13A, but without tobramycin. Regardless of volume injected (squares and triangles, Figure 13A), the sensors showed minimal response to the bolus delivery. The results in Figures 12 and 13 demonstrate the ability of vancomycin and tobramycin to freely diffuse between blood and the dermis ISF. Probing blood concentrations with our microneedles by drawing blood from the animals and pipetting the blood into the devices further shows that the two compartments are in rapid equilibrium (i.e., the ISF vs. blood levels are quantitatively similar). However, it is still unclear if the same is true for less soluble therapeutics with higher extent of protein binding, which we studied next. [0114] Microneedle Patch Monitors Hydrophobic Analytes in Rodent Dermal ISF. Having successfully demonstrated in vivo measurements of hydrophilic small molecules with our microneedle E-AB sensors, we next investigated the detection of a hydrophobic molecule: phenylalanine. This essential amino acid is found in many foods, but its safe intake is restricted for individuals with phenylketonuria (PKU), a rare genetic disorder causing phenylalanine buildup in the body. Recent research demonstrated E-AB sensors for rapid phenylalanine detection in the bloodstream of rodents. Here, we present the first microneedle E-AB sensors for detecting phenylalanine in the dermal ISF.

[0115] Like previous measurements, we first identified the optimal frequencies for phenylalanine detection using a frequency map in 50% rat serum solution with and without saturating phenylalanine. Following a 15-minute baseline collection, two individual animals received intravenous bolus injections of 30 mg/kg (black dot) and 60 mg/kg (red square) phenylalanine (Figure 14A). To convert the sensor signal into ISF concentrations, we constructed a dose-response calibration curve ex vivo using the same frequencies (300 Hz and 10 Hz) in 50% serum solution (Figure 14B). This curve, together with the Hill equation, was then employed to translate the in vivo sensor signals into drug concentrations, allowing us to obtain complete pharmacokinetic profiles for both bolus injections. Fitting these profiles with a pharmacokinetic model revealed a rapid absorption rate constant and a slow elimination rate constant for both phenylalanine doses (Table 1 ). Notably, this behavior resembles the rate constants observed for the 28 mg/kg tobramycin injection. We attribute this similarity in disposition kinetics to the small size (molecular weight under 500 Da) of both tobramycin and phenylalanine.

Table 1. Pharmacokinetic data derived from Figure 12-14.

Rat Small IV bolus CMax, ISF tplateau abs Xelim AUC

# molecules dose (pM) (pM*h) injected

1 Vancomycin 20 0.5 ± 0.1 39 2.7 ± 1.1 2.7 ± 1.0 0.4 ± 0.1

2 Tobramycin 28 67 ± 18 43 4.0 ± 0.8 1.8 ± 0.3 58.0 ± 25.5 3 Phenylalanine 30 267 ± 70 36 5.0 ± 0.7 1.0 ± 0.1 301.6 ± 145.9

4 Phenylalanine 60 287 ± 2 26 17.5 ± 1.6 1.3 ± 0.1 186.6 ± 16.9

[0116] Hydrophobic chemotherapeutic drugs are often administered to cancer patients through intravenous infusions to maintain stable drug levels in the blood. However, current dosing strategies based on body surface area (BSA) don't account for the complex ways these drugs are eliminated. Advanced tools for precise, real-time drug level measurement would provide doctors with a better way to prescribe effective chemotherapeutic dosages. Here, we investigated the ability of our microneedle E-AB sensors to detect two hydrophobic anticancer drugs: irinotecan and doxorubicin.

[0117] We administered intravenous bolus injections of 31 mg/kg (Figure 14D) and 34 mg/kg irinotecan to separate rodents. Despite using KDM for improved signal processing, we did not observe significant changes in sensor signal. Similarly, two different doses of doxorubicin administered intravenously to other rodents resulted in no significant sensor response (Figure 14E). We believe the minimal response to both drugs is due to their protein-binding properties and hydrophobicity, which likely hinder their transport and partitioning from blood into the dermal ISF.

[0118] 3. DISCUSSION

[0119] Continuous and real-time monitoring of specific molecules within dermal ISF could revolutionize medicine. This technology would allow both doctors and patients to tailor treatment plans for precision medicine, leading to better outcomes. Motivated by this goal, we report the development of the first microneedle patch equipped with five E-AB sensors for continuous, real-time monitoring of small molecules in rodents. The E-AB sensors were functionalized on the AuNPs-coated microneedles with self-assembled monolayers of a mixture of mercaptohexanol and thiolated- and redox reporter-modified aptamers. Due to the optimized electroplating and cleaning protocol, the E-AB sensors immobilized on the AuNPs-coated microneedles surface are not only reproducible but also robust, allowing multiple skin reinsertions while retaining the immobilized aptamers on the surface. Additionally, we employed the magnetic placement approach from our previous work to secure the microneedle patch on rodent skin, preventing needle retraction during measurements. Moreover, to ensure a stable sensor baseline prior to drug administration, we applied kinetic differential measurements that enable drift correction based on the measurements performed at two frequencies. The groundbreaking microneedle E-AB sensors enable continuous, real-time, and multiredout monitoring of diverse drugs, laying the groundwork for personalized treatment strategies and improved patient outcomes.

[0120] To interpret the pharmacokinetic profiles of various small molecules, we employed two-compartment pharmacokinetics models describing drug absorption into and excretion from dermal ISF. Vancomycin, a hydrophilic antibiotic (20 mg/kg, I.V. bolus), exhibited a pharmacokinetic profile consistent with flip-flop pharmacokinetics, in which slow drug absorption is the rate-limiting step. In contrast, tobramycin (28 mg/kg, I.V. bolus), another hydrophilic antibiotic, displayed a typical two-compartment model with rapid, first-order drug absorption. The observed difference in pharmacokinetics might be attributed to the size disparity between vancomycin (1449 Da) and tobramycin (467 Da), with vancomycin’s large size hindering absorption. Among the hydrophobic molecules, the small, nonpolar amino acid phenylalanine (I.V. administration) showed pharmacokinetic profiles best described by a two-compartment model. However, intravenous injection of the hydrophobic chemotherapeutic drugs irinotecan and doxorubicin resulted in minimal signal response from the microneedle E-AB sensors. This is likely due to the high degree of plasma protein binding and hydrophobicity of these drugs, limiting the free fraction available for diffusion into the ISF.

[0121] By leveraging the multi-redout capacity of our microneedle patch, we minimized the influence of sensor-to-sensor variability in our investigation of rat-to- rat variability. Although the overlapping pharmacokinetic profiles from the 40 mg/kg vancomycin bolus suggest minimal variability in this study, statistically powered studies with larger animal groups are needed. These future studies should focus on factors like age, weight, body fat, hydration, and sensor placement to gain a more comprehensive understanding of ISF-based drug pharmacokinetics in anesthetized rodents and their potential applications in preclinical research. Ultimately, we will focus on applying our microneedle E-AB sensors for chronic monitoring of small molecules in awake animals.

[0122] 4. MATERIALS AND METHODS

[0123] As described herein, the materials and methods included chemicals and materials, microneedle patch fabrication, electroplating and e-cleaning of the microneedles, E-AB sensor fabrication and calibration, electrochemical methods and data analysis software, and regression analysis of pharmacokinetic data.

[0124] Some further aspects are defined in the following clauses:

[0125] Clause 1 : An electrochemical sensor device, comprising: a body structure comprising first and second surfaces; a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected, or connectable, to a detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer; and, a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule.

[0126] Clause 2: The electrochemical sensor device of Clause 1 , wherein the electrochemical sensor device is configured to generate one or more square wave or cyclic voltammograms using square wave voltammetry (SWV) and/or cyclic voltammetry (CV) and determine a change in peak-to-peak separation, EP.T, from the voltammograms to detect the target molecule when the biomolecular receptors bind the target molecule.

[0127] Clause 3: The electrochemical sensor device of Clause 1 or Clause 2, wherein the electrochemical sensor device is configured to perform drift correction.

[0128] Clause 4: The electrochemical sensor device of any one of the preceding Clauses 1-3, wherein the drift correction is performed using kinetic differential measurements (KDM).

[0129] Clause 5: The electrochemical sensor device of any one of the preceding Clauses 1-4, wherein the electrochemical sensor device is configured to generate substantially stable baseline electrochemical signals.

[0130] Clause 6: The electrochemical sensor device of any one of the preceding Clauses 1-5, wherein a thickness of the electrically conductive layer is less than about 1 pm.

[0131] Clause 7: The electrochemical sensor device of any one of the preceding Clauses 1-6, wherein the body structure comprises a substantially circular cross-sectional shape, a substantially square cross-sectional shape, or a substantially rectangular cross-sectional shape.

[0132] Clause 8: The electrochemical sensor device of any one of the preceding Clauses 1-7, wherein the plurality of microneedles is arrayed in a substantially circular format, a substantially square format, or a substantially rectangular format.

[0133] Clause 9: The electrochemical sensor device of any one of the preceding Clauses 1-8, wherein at least one of the microneedles in the plurality of microneedles has a substantially circular, a substantially square, or a substantially rectangular cross-sectional shape.

[0134] Clause 10: The electrochemical sensor device of any one of the preceding Clauses 1-9, wherein at least of the microneedles in the plurality of microneedles has a substantially conical shape, a substantially pyramidal shape, or a substantially tapered shape.

[0135] Clause 11 : The electrochemical sensor device of any one of the preceding Clauses 1-10, wherein edges of at least one pair of microneedles in the plurality of microneedles are separated from one another by about 5 mm or less.

[0136] Clause 12: The electrochemical sensor device of any one of the preceding Clauses 1-11 , wherein the detector comprises an electrochemical analyzer.

[0137] Clause 13: The electrochemical sensor device of any one of the preceding Clauses 1-12, wherein the electrochemical analyzer comprises a potentiostat.

[0138] Clause 14: The electrochemical sensor device of any one of the preceding Clauses 1-13, wherein the electrically conductive layer is configured to wirelessly connect to the detector.

[0139] Clause 15: The electrochemical sensor device of any one of the preceding Clauses 1-14, wherein the detector is configured to measure square wave voltammograms.

[0140] Clause 16: The electrochemical sensor device of any one of the preceding Clauses 1-15, wherein an electrically insulating cover structure is attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure.

[0141] Clause 17: The electrochemical sensor device of any one of the preceding Clauses 1-16, wherein the body structure and the plurality of microneedles are fabricated using a stereolithography (SLA) three-dimensional (3D) printing technique.

[0142] Clause 18: The electrochemical sensor device of any one of the preceding Clauses 1-17, wherein the plurality of microneedles comprises twelve microneedles, wherein seven of the microneedles are working electrodes, wherein two of the microneedles are reference electrodes, and wherein three of the microneedles are counter electrodes.

[0143] Clause 19: The electrochemical sensor device of any one of the preceding Clauses 1-18, wherein the body structure comprises a cross-sectional dimension of about 30 mm or less.

[0144] Clause 20: The electrochemical sensor device of any one of the preceding Clauses 1-19, wherein the magnet comprises a permanent magnet. [0145] Clause 21 : The electrochemical sensor device of any one of the preceding Clauses 1-20, wherein the magnet comprises a neodymium ring magnet.

[0146] Clause 22: The electrochemical sensor device of any one of the preceding Clauses 1-21 , wherein the magnet comprises a cross-sectional dimension of about 25 mm or less.

[0147] Clause 23: The electrochemical sensor device of any one of the preceding Clauses 1-22, wherein the magnet comprises a thickness of about 5 mm or less.

[0148] Clause 24: The electrochemical sensor device of any one of the preceding Clauses 1-23, wherein the plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrodeblocking alkanethiols, one or more alkanethiols, and one or more redox reporter- tagged aptamers.

[0149] Clause 25: The electrochemical sensor device of any one of the preceding Clauses 1-24, wherein the subject comprises a mammalian subject.

[0150] Clause 26: The electrochemical sensor device of any one of the preceding Clauses 1-25, wherein the mammalian subject comprises a human subject.

[0151] Clause 27: The electrochemical sensor device of any one of the preceding Clauses 1-26, wherein the mammalian subject is from an order Rodentia.

[0152] Clause 28: The electrochemical sensor device of any one of the preceding Clauses 1-27, wherein the metallic structure is configured to maintain a substantially constant and substantially homogeneous pressure across the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject.

[0153] Clause 29: The electrochemical sensor device of any one of the preceding Clauses 1-28, wherein the metallic structure is configured to maintain a substantially constant penetration depth of the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject.

[0154] Clause 30: The electrochemical sensor device of any one of the preceding Clauses 1-29, wherein the metallic structure is disposed at least proximal to an ear of the subject.

[0155] Clause 31 : The electrochemical sensor device of any one of the preceding Clauses 1-30, wherein a metallic plate comprises the metallic structure.

[0156] Clause 32: The electrochemical sensor device of any one of the preceding Clauses 1-31 , wherein the metallic structure is implanted in the subject.

[0157] Clause 33: The electrochemical sensor device of any one of the preceding Clauses 1-32, wherein the metallic structure comprises a biocompatible plate fabricated from a medical grade, ferromagnetic steel.

[0158] Clause 34: The electrochemical sensor device of any one of the preceding Clauses 1-33, wherein the plurality of microneedles is fabricated integral with the body structure.

[0159] Clause 35: The electrochemical sensor device of any one of the preceding Clauses 1-34, wherein the electrically conductive layer comprises a chromium layer and/or a gold layer.

[0160] Clause 36: The electrochemical sensor device of any one of the preceding Clauses 1-35, wherein the electrically conductive layer of the reference electrode further comprises at least one silver layer.

[0161] Clause 37: The electrochemical sensor device of any one of the preceding Clauses 1-36, wherein the electrically conductive layer of the counter electrode further comprises at least one platinum layer.

[0162] Clause 38: The electrochemical sensor device of any one of the preceding Clauses 1-37, wherein at least one microneedle of the plurality of microneedles extends from the surface of the body structure by about 0.5 mm to about 2 mm.

[0163] Clause 39: The electrochemical sensor device of any one of the preceding Clauses 1-38, wherein a base portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 400 pm.

[0164] Clause 40: The electrochemical sensor device of any one of the preceding Clauses 1-39, wherein a tip portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 100 pm. [0165] Clause 41: The electrochemical sensor device of any one of the preceding Clauses 1-40, wherein the biomolecular receptor comprises an aptamer.

[0166] Clause 42: The electrochemical sensor device of any one of the preceding Clauses 1-41 , wherein the aptamer comprises a 5’- thiol modification and a 3’ methylene blue modification.

[0167] Clause 43: The electrochemical sensor device of any one of the preceding Clauses 1-42, wherein the biomolecular receptor comprises a nucleic acid molecule.

[0168] Clause 44: The electrochemical sensor device of any one of the preceding Clauses 1-43, wherein the redox reporters comprise methylene blue (MB) or an osmium-based complex.

[0169] Clause 45: The electrochemical sensor device of any one of the preceding Clauses 1-44, wherein the target molecule comprises a therapeutic agent.

[0170] Clause 46: The electrochemical sensor device of any one of the preceding Clauses 1-45, wherein the therapeutic agent comprises an antibiotic.

[0171] Clause 47: The electrochemical sensor device of any one of the preceding Clauses 1-46, wherein the target molecule comprises a metabolite and/or an electrolyte.

[0172] Clause 48: The electrochemical sensor device of any one of the preceding Clauses 1-47, wherein the target molecule comprises a biomolecule.

[0173] Clause 49: The electrochemical sensor device of any one of the preceding Clauses 1-48, wherein the electrochemical sensor device is a wearable device.

[0174] Clause 50: A kit comprising the electrochemical sensor device of any one of the preceding Clauses 1-49.

[0175] Clause 51 : A method of detecting a target molecule in a subject, the method comprising: positioning a plurality of microneedles of an electrochemical sensor device in contact with a dermal interstitial fluid (ISF) of or in the subject, wherein the electrochemical sensor device comprises: a body structure comprising first and second surfaces; the plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected to a detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals; and a magnet connected to the second surface of the body structure, wherein the magnet magnetically attaches the electrochemical sensor device to a metallic structure disposed in and/or on a body of the subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with the dermal ISF of or in the subject, which dermal ISF comprises the target molecule; and, detecting the electrochemical signals using the detector, thereby detecting the target molecule in the subject.

[0176] Clause 52: The method of Clause 51 , further comprising implanting the metallic structure in the body of the subject prior to the positioning step.

[0177] Clause 53: The method of Clause 51 or Clause 52, comprising implanting the metallic structure in an ear lobe of the subject prior to the positioning step.

[0178] Clause 54: The method of any one of the preceding Clauses 51-53, comprising substantially continuously monitoring the target molecule in the subject during a selected duration of time.

[0179] Clause 55: The method of any one of the preceding Clauses 51-54, comprising intermittently monitoring the target molecule in the subject during a selected duration of time.

[0180] Clause 56: The method of any one of the preceding Clauses 51-55, comprising performing a pharmacokinetic, pharmacodynamic, and/or toxicology assessment of the target molecule in the subject.

[0181] Clause 57: The method of any one of the preceding Clauses 51-56, comprising detecting the electrochemical signals in substantially real-time.

[0182] Clause 58: The method of any one of the preceding Clauses 51-57, further comprising performing drift correction of the electrochemical sensor device using one or more kinetic differential measurements (KDM).

[0183] Clause 59: The method of any one of the preceding Clauses 51-58, wherein the electrochemical sensor device further comprises an electrically insulating cover structure attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure.

[0184] Clause 60: The method of any one of the preceding Clauses 51-59, wherein the plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged nucleic acids.

[0185] Clause 61 : The method of any one of the preceding Clauses 51-60, wherein the redox reporters comprise methylene blue (MB).

[0186] Clause 62: A method of producing an electrochemical sensor device, the method comprising: forming a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure; disposing at least one electrically conductive layer on at least a portion of the plurality of microneedles, wherein electrically conductive layer is operably connectable to a detector; attaching a plurality of biomolecular receptor-bound redox reporters to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters are configured to undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer; and connecting a magnet to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule, thereby producing the electrochemical sensor device.

[0187] Clause 63: The method of Clause 62, further comprising attaching an electrically insulating cover structure to the first surface of the body structure such that the plurality of microneedles extends through the electrically insulating cover structure.

[0188] Clause 64: A system for detecting a target molecule in a subject, the system comprising: a detector; an electrochemical sensor device, comprising: a body structure comprising first and second surfaces; a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein electrically conductive layer is operably connected to the detector; a plurality of biomolecular receptorbound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals that are detected by the detector; and, a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule; a processor; and a memory communicatively coupled to the processor, the memory storing non-transitory computer executable instructions which, when executed on the processor, perform operations comprising: detecting the electrochemical signals using the detector.

[0189] While the foregoing disclosure has been described in some detail by way of illustration and example for purposes of clarity and understanding, it will be clear to one of ordinary skill in the art from a reading of this disclosure that various changes in form and detail can be made without departing from the true scope of the disclosure and may be practiced within the scope of the appended claims. For example, all the methods, devices, systems, computer readable media, and/or component parts or other aspects thereof can be used in various combinations. All patents, patent applications, websites, other publications or documents, and the like cited herein are incorporated by reference in their entirety for all purposes to the same extent as if each individual item were specifically and individually indicated to be so incorporated by reference.

Claims

WHAT IS CLAIMED IS:

1 . An electrochemical sensor device, comprising: a body structure comprising first and second surfaces; a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected, or connectable, to a detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer; and, a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule.

2. The electrochemical sensor device of claim 1 , wherein the electrochemical sensor device is configured to generate one or more square wave or cyclic voltammograms using square wave voltammetry (SWV) and/or cyclic voltammetry (CV) and determine a change in peak-to-peak separation, EP.T, from the voltammograms to detect the target molecule when the biomolecular receptors bind the target molecule.

3. The electrochemical sensor device of claim 1 , wherein the electrochemical sensor device is configured to perform drift correction.

4. The electrochemical sensor device of claim 3, wherein the drift correction is performed using kinetic differential measurements (KDM).

5. The electrochemical sensor device of claim 1 , wherein the electrochemical sensor device is configured to generate substantially stable baseline electrochemical signals.

6. The electrochemical sensor device of claim 1 , wherein a thickness of the electrically conductive layer is less than about 1 pm.

7. The electrochemical sensor device of claim 1 , wherein the body structure comprises a substantially circular cross-sectional shape, a substantially square cross-sectional shape, or a substantially rectangular cross-sectional shape.

8. The electrochemical sensor device of claim 1 , wherein the plurality of microneedles is arrayed in a substantially circular format, a substantially square format, or a substantially rectangular format.

9. The electrochemical sensor device of claim 1 , wherein at least one of the microneedles in the plurality of microneedles has a substantially circular, a substantially square, or a substantially rectangular cross-sectional shape.

10. The electrochemical sensor device of claim 1 , wherein at least of the microneedles in the plurality of microneedles has a substantially conical shape, a substantially pyramidal shape, or a substantially tapered shape.

11 . The electrochemical sensor device of claim 1 , wherein edges of at least one pair of microneedles in the plurality of microneedles are separated from one another by about 5 mm or less.

12. The electrochemical sensor device of claim 1 , wherein the detector comprises an electrochemical analyzer.

13. The electrochemical sensor device of claim 12, wherein the electrochemical analyzer comprises a potentiostat.

14. The electrochemical sensor device of claim 1 , wherein the electrically conductive layer is configured to wirelessly connect to the detector.

15. The electrochemical sensor device of claim 1 , wherein the detector is configured to measure square wave voltammograms.

16. The electrochemical sensor device of claim 1 , wherein an electrically insulating cover structure is attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure.

17. The electrochemical sensor device of claim 1 , wherein the body structure and the plurality of microneedles are fabricated using a stereolithography (SLA) three-dimensional (3D) printing technique.

18. The electrochemical sensor device of claim 1 , wherein the plurality of microneedles comprises twelve microneedles, wherein seven of the microneedles are working electrodes, wherein two of the microneedles are reference electrodes, and wherein three of the microneedles are counter electrodes.

19. The electrochemical sensor device of claim 1 , wherein the body structure comprises a cross-sectional dimension of about 30 mm or less.

20. The electrochemical sensor device of claim 1 , wherein the magnet comprises a permanent magnet.

21 . The electrochemical sensor device of claim 1 , wherein the magnet comprises a neodymium ring magnet.

22. The electrochemical sensor device of claim 1 , wherein the magnet comprises a cross-sectional dimension of about 25 mm or less.

23. The electrochemical sensor device of claim 1 , wherein the magnet comprises a thickness of about 5 mm or less.

24. The electrochemical sensor device of claim 1 , wherein the plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged aptamers.

25. The electrochemical sensor device of claim 1 , wherein the subject comprises a mammalian subject.

26. The electrochemical sensor device of claim 25, wherein the mammalian subject comprises a human subject.

27. The electrochemical sensor device of claim 25, wherein the mammalian subject is from an order Rodentia.

28. The electrochemical sensor device of claim 1 , wherein the metallic structure is configured to maintain a substantially constant and substantially homogeneous pressure across the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject.

29. The electrochemical sensor device of claim 1 , wherein the metallic structure is configured to maintain a substantially constant penetration depth of the plurality of microneedles when the plurality of microneedles is disposed in contact with the dermal ISF of the subject.

30. The electrochemical sensor device of claim 1 , wherein the metallic structure is disposed at least proximal to an ear of the subject.

31 . The electrochemical sensor device of claim 1 , wherein a metallic plate comprises the metallic structure.

32. The electrochemical sensor device of claim 1 , wherein the metallic structure is implanted in the subject.

33. The electrochemical sensor device of claim 1 , wherein the metallic structure comprises a biocompatible plate fabricated from a medical grade, ferromagnetic steel.

34. The electrochemical sensor device of claim 1 , wherein the plurality of microneedles is fabricated integral with the body structure.

35. The electrochemical sensor device of claim 1 , wherein the electrically conductive layer comprises a chromium layer and/or a gold layer.

36. The electrochemical sensor device of claim 1 , wherein the electrically conductive layer of the reference electrode further comprises at least one silver layer.

37. The electrochemical sensor device of claim 1 , wherein the electrically conductive layer of the counter electrode further comprises at least one platinum layer.

38. The electrochemical sensor device of claim 1 , wherein at least one microneedle of the plurality of microneedles extends from the surface of the body structure by about 0.5 mm to about 2 mm.

39. The electrochemical sensor device of claim 1 , wherein a base portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 400 pm.

40. The electrochemical sensor device of claim 1 , wherein a tip portion of at least one microneedle of the plurality of microneedles has a diameter of less than about 100 pm.

41 . The electrochemical sensor device of claim 1 , wherein the biomolecular receptor comprises an aptamer.

42. The electrochemical sensor device of claim 41 , wherein the aptamer comprises a 5’- thiol modification and a 3’ methylene blue modification.

43. The electrochemical sensor device of claim 1 , wherein the biomolecular receptor comprises a nucleic acid molecule.

44. The electrochemical sensor device of claim 1 , wherein the redox reporters comprise methylene blue (MB) or an osmium-based complex.

45. The electrochemical sensor device of claim 1 , wherein the target molecule comprises a therapeutic agent.

46. The electrochemical sensor device of claim 45, wherein the therapeutic agent comprises an antibiotic.

47. The electrochemical sensor device of claim 1 , wherein the target molecule comprises a metabolite and/or an electrolyte.

48. The electrochemical sensor device of claim 1 , wherein the target molecule comprises a biomolecule.

49. The electrochemical sensor device of claim 1 , wherein the electrochemical sensor device is a wearable device.

50. A kit comprising the electrochemical sensor device of claim 1 .

51 . A method of detecting a target molecule in a subject, the method comprising: positioning a plurality of microneedles of an electrochemical sensor device in contact with a dermal interstitial fluid (ISF) of or in the subject, wherein the electrochemical sensor device comprises: a body structure comprising first and second surfaces; the plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected to a detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals; and a magnet connected to the second surface of the body structure, wherein the magnet magnetically attaches the electrochemical sensor device to a metallic structure disposed in and/or on a body of the subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with the dermal ISF of or in the subject, which dermal ISF comprises the target molecule; and, detecting the electrochemical signals using the detector, thereby detecting the target molecule in the subject.

52. The method of claim 51 , further comprising implanting the metallic structure in the body of the subject prior to the positioning step.

53. The method of claim 52, comprising implanting the metallic structure in an ear lobe of the subject prior to the positioning step.

54. The method of claim 51 , comprising substantially continuously monitoring the target molecule in the subject during a selected duration of time.

55. The method of claim 51 , comprising intermittently monitoring the target molecule in the subject during a selected duration of time.

56. The method of claim 51 , comprising performing a pharmacokinetic, pharmacodynamic, and/or toxicology assessment of the target molecule in the subject.

57. The method of claim 51 , comprising detecting the electrochemical signals in substantially real-time.

58. The method of claim 51 , further comprising performing drift correction of the electrochemical sensor device using one or more kinetic differential measurements (KDM).

59. The method of claim 51 , wherein the electrochemical sensor device further comprises an electrically insulating cover structure attached to the first surface of the body structure, and wherein the plurality of microneedles extends through the electrically insulating cover structure.

60. The method of claim 51 , wherein the plurality of biomolecular receptorbound redox reporters operably attached to the electrically conductive layer comprise one or more self-assembling biosensing monolayers that comprise one or more electrode-blocking alkanethiols, one or more alkanethiols, and one or more redox reporter-tagged nucleic acids.

61 . The method of claim 51 , wherein the redox reporters comprise methylene blue (MB).

62. A method of producing an electrochemical sensor device, the method comprising: forming a body structure comprising first and second surfaces, and a plurality of microneedles extending from the first surface of the body structure; disposing at least one electrically conductive layer on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connectable to a detector; attaching a plurality of biomolecular receptor-bound redox reporters to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters are configured to undergo conformational changes when the biomolecular receptors bind a target molecule to produce one or more electrochemical signals that are detected by the detector when the detector is operably connected to the electrically conductive layer; and connecting a magnet to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule, thereby producing the electrochemical sensor device.

63. The method of claim 62, further comprising attaching an electrically insulating cover structure to the first surface of the body structure such that the plurality of microneedles extends through the electrically insulating cover structure.

64. A system for detecting a target molecule in a subject, the system comprising: a detector; an electrochemical sensor device, comprising: a body structure comprising first and second surfaces; a plurality of microneedles extending from the first surface of the body structure, wherein the plurality of microneedles comprises at least one microneedle configured as a working electrode, at least one microneedle configured as a reference electrode, and at least one microneedle configured as a counter electrode; at least one electrically conductive layer disposed on at least a portion of the plurality of microneedles, wherein the electrically conductive layer is operably connected to the detector; a plurality of biomolecular receptor-bound redox reporters operably attached to the electrically conductive layer, wherein biomolecular receptors of the plurality of biomolecular receptor-bound redox reporters undergo conformational changes when the biomolecular receptors bind the target molecule to produce one or more electrochemical signals that are detected by the detector; and, a magnet connected, or connectable, to the second surface of the body structure, wherein the magnet is configured to magnetically attach the electrochemical sensor device to a metallic structure disposed in and/or on a body of a subject such that the electrochemical sensor device is reversibly retained in position on a surface of the body of the subject at least proximal to the metallic structure and a least some of the plurality of microneedles are disposed in contact with a dermal interstitial fluid (ISF) of or in the subject, which dermal ISF comprises the target molecule; a processor; and, a memory communicatively coupled to the processor, the memory storing non-transitory computer executable instructions which, when executed on the processor, perform operations comprising: detecting the electrochemical signals using the detector.