Active reset method for biosensor probes

Abstract

Provided herein is a novel biosensor integrated into a microdevice to collect biological samples and continuously detect the target analytes. More specifically, provided herein is a method of actively resetting biosensors used to in situ and reagentless sensor regeneration after target binding, enabling continuously detection of analytes in a biological sample. The Active- reset methodology can be applied to any aptamer-based biosensors that employ a reagentless methodology to detect both electroactive target analytes (using a monolayer transporter biosensor) and non-electroactive target analytes (using a redox reporter), with various sizes and affinities.

Description

ACTIVE RESET METHOD FOR BIOSENSOR PROBESPRIORITY

[0001] This application claims priority to U.S. provisional application No.63 / 563,662 filed March 11, 2024, which is incorporated herein by reference in its entirety.TECHNICAL FIELD

[0002] This disclosure relates to an implantable biosensor that continuously detects the presence and concentration of molecular analytes in a biological sample. More specifically, this disclosure relates to a reagentless electrochemical biosensor integrated into a microdevice and methods of use thereof for in vivo detection of biomolecules and proteins in different sizes and affinities in real time.BACKGROUND

[0003] The following description of the background of the present technology is provided simply as an aid in understanding the present technology and is not admitted to describe or constitute prior art to the present technology.

[0004] In order to promote specificity and sensitivity, many reagentless detection methods rely on receptors that have a high degree of affinity with the target analyte. This presents an additional problem when attempting to integrate these systems into continuous monitoring mechanisms, as the receptors may remain bound to the target analyte between rounds of detection, leading to inaccurate readouts of analyte presence and concentration. Thus, development of versatile sensors for reagentless analyte detection that can track molecular analytes in biological fluids and can be rapidly reset or regenerated between rounds of detection is desirable.SUMMARY

[0005] The present disclosure provides a method for actively resetting biosensor probes in a biosensor system. In some embodiments, the active resetting method allows biosensor probes in a biosensor system to provide more accurate detection of analytes than biosensor systems that rely only on passive resetting of biosensor probes. It is to be understood that thedisclosed embodiments are merely exemplary, and accordingly, the invention may be embodied in various and alternative forms. The specific structural and functional details disclosed herein are not to be interpreted as limiting but merely as a basis for the claims and as a representative basis for teaching one skilled in the art to variously employ the embodiments described herein.

[0006] In a first aspect, the present disclosure provides a method actively resetting the immobilized biosensor probes in a biosensor system comprising rapidly oscillating the biosensor probes by inducing a change in electric current and / or applying alternating electric field to the biosensor probes, wherein the biosensor probe comprises an analyte receptor that can bind to the analyte and wherein rapid oscillation of the biosensor probe results in the release and fast dissociation of an analyte bound by the biosensor probe.

[0007] In a second aspect, the present disclosure provides a method of detecting an analyte using a biosensor system, comprising allowing a biosensor probe to contact a biological sample that may contain the analyte, wherein the immobilized biosensor probe is configured to bend and transfer an electron with the electrode in response to a change in electric current and / or an applied alternating electric field and, after a first detecting step is performed, applying a change in electric current or alternating electric field to the biosensor probe to cause rapid oscillation to shake off the captured analyte before performing subsequent detection steps with the biosensor probe.

[0008] In a third aspect, the present disclosure provides a method of determining a concentration of an analyte, the method comprising (a) applying a biological sample to a biosensor system, wherein the biosensor system comprises: (i) an electrode, and (ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus and an analyte receptor attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure, wherein the analyte present in the biological sample binds to the analyte receptors of the plurality of the biosensor probes; (b) applying an electric current to the biological sample and the plurality of the biosensor probes via the electrode of the biosensorsystem, wherein the biosensor probes and analyte bound by the analyte receptors attached to the biosensor probes are displaced toward the surface of the electrode; (c) measuring the chronoamperometric response of the electric current by translating the difference between an unbound electron transfer rate and a bound electron transfer rate; (d) calculating the concentration of the analyte in the biological sample based on a chronoamperometric response based on difference between the electrochemical signal produced by the biosensor probes that have bound the analyte and by biosensor probes that have not bound the analyte; and (e) actively resetting the biosensor probe by inducing rapid oscillation of the biosensor probes via the electrode using changes in electric current and / or applied electric field.

[0009] In a fourth aspect, the present disclosure provides a device for continuously detecting an analyte in a biological sample from a subject, the device comprising (a) a collector component for collecting the biological sample, wherein the collector component comprises a microneedle having three inlet channels that connect to a main channel in the interior of the microneedle; (b) a detection component attached to the collector component comprising (i) an electrode, and (ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus, wherein each biosensor probe comprises an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure and wherein each biosensor probe is bound to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure; (c) an electrochemical measurement component connected to the electrode of the detection component, wherein the electrochemical measurement component detects changes in the electric current at the surface of the electrode.

[0010] Further objects and advantages of the present invention will be clear from the description that follows.FIGURES

[0011] FIGS. 1A, IB, and 1C depict the active-reset sensor for continuous protein monitoring in vivo. FIG. 1A shows a schematic of a microdevice inserted into the skin (left panel). The device delivers interstitial fluid (ISF) to the electrodes upon insertion into theskin. Aptamer-based molecular pendulum sensors (middle panel) enable reagentless analysis of protein biomarkers. These constructs consist of two DNA strands of different lengths. One strand contains a thiol modification to facilitate gold binding and ferrocene (e- donor) modification for electrochemical measurements. The other strand hybridizes to the modified strand and is elongated to include an aptamer sequence for protein recognition. The hybridized region of the construct remains rigid such that the application of a positive current drags the construct towards the electrode surface in an arc-like motion. When the terminal end of the construct falls sufficiently close to the electrode, ferrocene is oxidized to ferrocenium, and faradaic current is recorded. The difference in electron transfer rates between unbound and bound construct configurations due to differences in hydrodynamic radius forms the basis for quantifying target protein concentration. Potential-mediated oscillation (right panel) enables active reset / regeneration of sensors. Application of an alternative potential results in oscillation of the molecular pendulum constructs and subsequent release of target proteins. FIG. IB shows the dynamics of oscillatory-based sensor regeneration method. This method enables continuous monitoring of increasing and decreasing protein levels. FIG. 1C shows the microdevice design and application. FIG. 1 C(i) shows the electrode-integrated microdevice and internal features. FIG. lC(ii) shows the design of the capillary -based microdevice along with subcomponent. FIG. lC(iii) shows the schematic of the implanted sensor on a rat for real-time, in vivo cytokine monitoring.

[0012] FIG. 2 shows the chronoamperometry for unbound (P1-P2 only) and bound (Pl- P2 + targeted protein) sensors. The sensors were tested for the detection of different proteins with varying morphologies and sizes, including a cardiovascular biomarker, Myeloperoxidase (MPO), and two inflammatory cytokines, Interleukin 6 (IL-6) and Tumor Necrosis Factor a (TNF-a). The sensors effectively distinguish between the bound and unbound states when the proteins are present at 100 pg / mL concentration in artificial interstitial fluid (ISF). The magnitude of signal change is influenced by the molecular mass and the concentration of the target protein. Notably, the signal change for MP, which has a molecular mass approximately five times higher than TNF-a, is considerably greater in the bound state compared to TNF-a.

[0013] FIGS. 3 A, 3B, and 3C show calibration curves for IL-6, MPO, and TNF-a proteins in simulated ISF. The sensor results show the sensitivities and wide dynamic ranges of the sensors with the limit of detection of 0.65 pg / ml, 0.67 pg / ml, and 0.88 pg / ml for IL-6, MPO, and TNF-a, respectively. Error bars depict standard deviations, with signal change percentage computed based on the baseline signal (lb) using the following formula: (I - lb) / lb x 100. Here, lb represents the mean value of the unbound signals (n=6). The centers of error bars indicate mean values. Signal change values from each concentration show statistical significance compared to one another (P values are indicated by stars, **P<0.01, ***p<0.001, ****P<0.0001; two-sided, two-sample t-tests). Inset: chronoamperometry signal for unbound and bound states with various concentrations. It is noteworthy that for the MPO aptamer employed KD=160 pM a concentration of 1000 pg / ml of MPO proteins is equal to 12.5 pM, which is lower than the saturation concentration (9XKD = 1.6 nM). Additionally, the saturation concentrations for these sensors are indeed higher than the physiologically relevant concentrations of the target proteins.

[0014] FIG. 4 depict the selectivity evaluation of MPO, IL-6, and TNF-a sensors. The specificity of the selected biorecognition elements is challenged with a panel of various proteins, such as BNP, IgG, Cysteine, CRP, and glucose, confirming that the sensors are significantly selected to their target protein. The data is mean±SD, with n=3

[0015] FIG. 5 shows the rheological test for the simulated ISF viscosity. The results demonstrate that a mixture comprising 10% glycerin and 90% water closely matches the viscosity of ISF, approximately 1.2 mPa s. The 1 :9 glycerin-to-simulated ISF ratio is employed to mimic the viscosity, composition, and pH of subcutaneous ISF

[0016] FIGS. 6A, 6B, 6C, 6D, 6E, 6F, 6G, 6H, 61, 6J, 6K, 6L, and 6M ©depict activereset sensor development. FIG. 6A depicts the different detected currents for bound versus unbound analytes. Chronoamperometry is used for the detection of target proteins that slow electron transfer kinetics upon binding. FIG. 6B shows the measurement of a model protein (myeloperoxidase (MPO)) at a concentration of 200 pg / mL using aptamers with dissociation constants (KD) of 27 nM, 2 nM, and 160 pM, representing low, medium, and high affinities, respectively. FIG. 6C shows oscillation-based regeneration / reset of the sensor in less thanone minute. FIG. 6D depicts the sensor signal behavior during regeneration. This behavior varies depending on the voltage and frequency applied. Complete regeneration is achieved using a frequency of 95 Hz and a voltage range of 0 to 0.3 V. Very low frequency (20 Hz) and low voltage (0. IV) do not significantly affect the sensor signal. When employing a voltage range of 0 to 0.2V and a frequency of 50 Hz, the sensor undergoes partial regeneration. FIG. 6E shows the potentials and frequencies required for various affinities to dissociate target proteins. FIG. 6F shows continuous measurement with changing MPO levels from 10 pg / mL to 1000 pg / mL concentrations, over a 3+ hour experiment. FIG. 6G shows a schematic showing oscillation-based sensor regeneration. Application of an alternating potential oscillates the sensor, which leads to decreasing aptamer-protein interactions and subsequent protein release. FIG. 6H shows MD analysis of aptamer-protein in the absence and presence of oscillation. Computed numbers of interactions between aptamer and protein are shown in the absence (i) and presence (ii) of oscillation. FIG. 6H(iii) shows the Solvent Accessible Surface Area (SASA) of P2 aptamer in the complex state. FIG. 6H(iv) shows the RMSF values of P2 aptamer in the complex state. The hydrogen bond number between protein and water molecules are shown in FIG. 6(v), and the hydrogen bond number between the P2 aptamer and water molecules are shown in FIG. 6(vi). FIG. 61 shows a confirmation of protein dissociation using quartz crystal microbalance (QCM) measurements. FIG. 6J shows the regeneration percentage based on frequency and applied voltage level tor the MPO aptamers with low, medium, and high affinities PIG. K shows m

[0017] FIG. 7 depicts the repeatability evaluation of the regeneration methodology, visualizing the real-time monitoring of MPO protein with a 100 pg / mL concentration using the high-affinity aptamer for two hours.

[0018] FIG. 8 shows the effect of repetitive regeneration on sensor functionality. Consecutive regeneration was induced by applying an alternating potential between 0 to 0.36 V with 100 Hz frequency using high-affinity MPO receptor targeting 100 pg / mL. The signal reduced by around 7% after 12 rounds of regeneration.

[0019] FIGS. 9A and 9B shows the frequency effect on the ferrocene peak. The ferrocene reporter functionality was tested in various oscillation frequencies of the bound molecular pendulum with 200 pg / mL MPO protein using square wave voltammetry (step 1 mV, amplitude 25 mV, and frequency 60 Hz) in the simulated ISF. Considering a maximum 5% peak reduction as the threshold, the results show that for frequencies of 100 Hz and less, the ferrocene molecule performance is over 95%.

[0020] FIGS. 10A and 10B show the molecular dynamic (MD) analyses of aptamerprotein in the absence and presence of oscillation. Computer numbers of interactions in 100 nanoseconds between the aptamer and protein in the absence (FIG. 10A) and presence (FIG. 10B) of oscillation are shown.

[0021] FIGS. 11A, 1 IB, 11C, 1 ID, and 1 IE show MD analysis of aptamer-protein complexes in the absence and presence of oscillation during the simulation. FIG. 11 A shows the Rg values of the Pl aptamer. FIG. 1 IB shows the RMSD values of the Pl aptamer. FIG. 11C shows the RMSF values of the Pl aptamer. FIG. 1 ID shows the Rg values of the P2 aptamer. FIG. 1 IE shows the RMSD value of the P2 aptamer.

[0022] FIGS. 12A, 12B, and 12C depicts the MD analysis of the aptamer-protein complex in the absence and presence of oscillation during the simulation. FIG. 12A shows the RMSF values of the protein. FIG. 12B shows the RMSD value of the protein. FIG. 12C shows the Rg value of the protein.

[0023] FIGS. 13A, 13B, and 13C show the radial distribution function (RDF) values. FIG. 13 A shows the RDF between P2 aptamer and water. FIG. 13B shows the RDF valuesbetween protein and P2 aptamer. FIG. 13C shows the RDF values between protein and water molecules.

[0024] FIGS. 14A, 14B, 14C, 14D, 14E, 14F, 14G, and 14H all show the interaction mode between different residues of the P2 aptamer and protein in the absence of oscillation based on cluster analysis. The H-bonds and salt-bridge interactions are shown as dashed yellow and purple lines. The secondary structure of the protein is shown as grey ribbon.

[0025] FIGS. 15A and 15B show the interaction mode between the P2 aptamer and the protein in the presence of oscillation. FIG. 15A shows the orientation and distance of residues T79, G80, and T81 towards the protein. FIG. 15B shows the P2 aptamer interacting residues with the protein residues. The H-bonds are shown as dashed yellow lines. The secondary structure of the protein is shown as grey ribbon.

[0026] FIGS. 16A and 16B depict snapshots of the gold-aptamer-protein system at various times. The sensor snapshots show the tethered DNA orientations during the 100 ns MD simulation in the absence (FIG. 16A) and presence (FIG. 16B) of oscillation.

[0027] FIGS. 17A, 17B, 17C, 17D, 17E, 17F, 17G, and 17H depict the implantable microdevice architecture, fluid dynamics, and sensing performance. FIG. 17A shows the schematic of the fabricated microdevice. FIG. 17B shows microscopy images of the device. FIG. 17B(i) shows a side view of the device with a scale bar of 500 pm. FIG. 17B(ii) shows the printed inlet features, with a scale bar of 500 pm. FIG. 17B(iii) shows the microdevice tip, with a scale bar of 50 pm. FIG. 17C shows an illustration in interstitial fluid (ISF) flowing into the channels through capillary force. FIG. 17D shows the computational fluid dynamic (CFD) simulation of the ISF flow when the dermis pressure is at its minimum (-4 mmHg), requiring less than 5 ms to fdl the needle. The inset shows a fluorescence microscopy image of the filled channel; the inset image’s scale bar is 300 pm. FIG. 17E shows CFD simulation of the protein concentration change. The average protein level in the main channel is calculated over time when the concentration in the microdevice is lower or higher than in the dermis layer. This causes the proteins to diffuse into or out and equilibrate in less than 20 minutes. FIG. 17F shows implanted microdevice, (i) The device mounted with housing for dermal implant, (ii) Administration of the sensing platforms on rat dorsal skin.(iii) Indentation caused by the microneedle penetration, (iv) H&E stained tissue section of the rat after microneedle penetration, illustrating the extent of penetration. FIG. 17G shows H&E stained sections of healthy rat organs with and without administration of the microdevice, indicating excellent biocompatibility of the device. Scale bar, 500 m. FIG. 17H shows qPCR results of a comprehensive panel of 66 genes spanning crucial inflammatory pathways that are systematically interrogated to discern any potential activation of inflammatory pathways. The 66 genes are ALOX12, ALOX5, ANXA1, ANXA3, ANXA5, CACNA1C, CACNA2D1, CD40, CD40LG, CYSLTR1, ES22, HPGD, HRH1, HRH2, HRH3, ICAM1, IL1R1, IL1R2, IL1RL1, IL2RB, IL2RG, ITGAL, ITGAM, ITGB1, ITGB2, KLK14, KLK15, KLK1B3, KLK7, KNG1, LTA4H, LTB4R1, LTB4R2, LTC4S, MAPK1, MAPK14, MAPK3, MAPK8, NFKB1, NR3C1, PDE4B, PDE4D, PLA2G2A, PLA2G2D, PLA2G5, PLA2G7, PLCB2, PLCB3, PLCB4, PLCD1, PLCE1, PLCG1, PLCG2, PTGDR, PTGER3, PTGFR, PTGIR, PTGIS, PTGS2, SCYE1, TBXA2R, TBXAS1,TNFRSF1A, TNFRSF1B, TNFSF13B and VCAM1.

[0028] FIGS. 18A, 18B, and 18C show the characteristics of the electrode-integrated microdevice. FIG. 18A shows a 3D and top view of the electrode-integrated microdevice depicting the locations of the gold working electrode (WE, D = 100 pm), platinum counter electrode (CE, D = 250 pm), and silver reference electrode (RE, D = 100 pm) within the microdevice. FIG. 18B shows the scanning electron microscopy (SEM) image of the working electrode’s surface. The rod-shaped working electrode is adorned with nanoneedles to enhance the sensor’s sensitivity. FIG. 18C shows the dynamic contact angle of the microdevice material, demonstrating the hydrophilicity of the printed microdevice, which facilitates the capillary force-driven drawing of interstitial fluid (ISF) from the dermis layer.

[0029] FIGS. 19A, 19B, 19C, 19D, 19E, and 19F show the internal design of the microdevice. FIGS. 19A and 18B depict side views of the microdevice displaying the electrode. FIG. 19C depicts the top view of the microdevice. FIGS. 19D and 19E depict the cross-section of the needle, illustrating the main channel and inlet nozzles. FIG. 19F shows the bottom view of the microdevice, showing the three inlet nozzles. The size of the nozzles (approximately 300 pm * 400 pm) and the channel (D ~ 300 pm) is designed to be smallenough to induce capillary force upon needle insertion into the skin and large enough to prevent clogging during measurements.

[0030] FIG. 20 shows access to the central channel. The integrated electrodes have access to the main channel through the side channel, which is approximately 10 pm in width. The needle is secured in a nut that provides access to the embedded electrodes.

[0031] FIGS. 21A, 21B, 21C, 21D, 21E, 21F, 21G, 21H, 211, and 21J shows the computational fluid dynamic simulation of ISF flow in the microdevice. Since the needle’s geometry is axisymmetric, an axisymmetric cross-section of the microdevice was simulated. The results illustrate the changes in pressure (FIGS. 21A-21E) and velocity (FIGS. 21F-21J) over time as the needle draws dermal ISF through capillary force.

[0032] FIGS. 22A and 22B show the velocity profdes of the final stages of ISF flow. Comparing the velocity profiles show how the stop valve halts ISF capillary flow, ensuring a consistent fluid volume over the integrated electrodes.

[0033] FIGS. 23A and 23B depict the biocompatibility testing results. The biocompatibility of the printed and unprinted microdevice resin (IP-Q photoresin) is assessed using cytotoxicity measurements on human dermal fibroblast cells (HDFs) with the MTT assay. FIG. 23A shows the cytotoxicity tests on pure resin with different concentrations for 48 hours, which indicates the microdevice’s biocompatibility with concentrations below 0.078 mg / mL. Given the minimal leaching of the printed microdevice (less than 0.001 mg / mL), this microdevice is non-cytotoxic in this application and can be used on the skin for at least 48 hours. FIG. 23B shows the viability of fibroblast cells exposed to the printed microdevice for 24 and 48 hours, which shows that HDF viability remains largely unaffected and confirms the material’s biocompatibility.

[0034] FIGS. 24A and 24B show the diffusion-based concentration change in the microdevice. The schematics of FIGS. 24A and 24B illustrate how proteins diffuse in or out of the microdevice, depending on the concentration difference between the microdevice channel and the dermis layer. When dermal concentration decreases, proteins diffuse out of the needle, leading to a rapid balance in protein levels within the microdevice (FIG. 24A).Conversely, when dermal protein concentration increases, proteins move into the microdevice (FIG. 24B).

[0035] FIGS. 25 A, 25B, 25C, 25D, and 25E show continuous monitoring of IL-6 and TNF-a in diabetic rats. FIG. 25A and 25B show real-time monitoring of IL-6 and TNF- a for about 6 hours for the biological target concentration range of 10 pg / ml to 500 pg / ml. The statistical analysis was performed for the last five signal points of each concentration, representing the statistically significant difference between various concentrations (P values are indicated by stars, ***P<0.001 and ****P<0.0001; two-sided, two-sample t-tests and ANOVA). The statistical non-significant differences are not shown. Fig. 25C shows the impact of fasting on pro-inflammatory cytokines: fasting enhances autophagy, increases production of adiponectin and ghrelin, reduces reactive oxygen species (ROS), and lowers immune cell activity. This combination leads to the downregulation of IL-6 and TNF-a. FIG. 25D shows the design of the animal study: longitudinal microdevice administration on diabetic rats and blood / ISF collection. Diabetic rats undergo fasting for 6 hrs, and the levels of IL-6 and TNF-a are measured every 20 min during the experiment. Different animals undergo fasting alone, fasting + insulin injection at 20 min, and the and fasting + LPS injection at the time 220 min. FIG. 25E shows the results of the animal study- IL-6 and TNF- a levels measured over time in ISF using the electrochemical platform for the fasting-only model, fasting+insulin model, and fasting+LPS model. The results are compared with ELISA measurement in ISF (extracted hydrogel microneedle patch) and in serum.

[0036] FIG. 26 shows the effect of insulin injection: Protein change rate of the microdevice compared with ELISA results obtained from serum and extracted ISF for the fasting-only model and fasting+insulin model. Insulin injection in ex vivo (ELISA) and in vivo (microdevice) measurements enhance the protein level decrease rate (anti-inflammation behavior) for both TNF-oc and IL-6 cytokines compared to the fasting-only model.

[0037] FIGS. 27A, 27B, and 27C depicts the blood glucose measurement over time in Type-1 diabetic rats that undergo insulin injection. FIG. 27A shows that fasting leads to a slight decrease in the glucose level. FIG. 27B shows that three-unit insulin injection leads toa fast reduction of the glucose level in the blood, reaching euglycemia and hypoglycemia in about 180 min. FIG. 27C shows that LPS injection leads to glucose level enhancement

[0038] FIGS. 28A, 28B, and 28C show quartz crystal microbalance (QCM) analysis.FIG. 28A depicts active-reset methodology using a sensor without ferrocene redox molecule for continuous IL-6 protein measurement. FIG. 28B Shows an evaluation of non-specific binding using scrambles sequence for P2 probe in the presence of IL-6 protein. FIG. 28C shows an evaluation of the stability of the sensor in the buffer.

[0039] FIG. 29. Evaluating the effect of negative alternating potential. Results indicate that applying negative potential ranges of (i) No. 1 (0 to -0.1V), (ii) No. 2 (0 to -0.2V), and (iii) No. 3 (0 to -0.3V) at a frequency of 100 Hz does not reset the bound signal.

[0040] FIG. 30 shows a comparison of active reset and passive reset for MPO aptamers with low, medium and high affinities.

[0041] FIGS. 31 A and 3 IB shows a stability test of sensors in unbound and bound states for over two weeks. The sensor is interrogated for over two weeks by detecting the target protein with the same concentration. The sensor is kept at 4°C in the buffer after each measurement.

[0042] FIG. 32 shows the stability of the immobilized probes and blocking layer during oscillation. The effect of various oscillation parameters, including (FIG. 32A) frequencies, (FIG. 32B) voltages, and (FIG. 32C) time, on the functionalized unbound electrode surface was evaluated. The stability of the unbound chronoamperometry results confirms that there is no significant change in the signal for oscillations with frequencies up to 10 kHz, voltages up to 350 mV, and times up to 120 seconds.

[0043] FIGS. 33A and 33B shows real-time monitoring of various target analytes possessing different sizes using the monolayer biosensor transporter and molecular pendulum methods. The serotonin biomarker is continuously monitored using monolayer transporter biosensors with aptamer-based receptors. The insulin biomarker is continuously monitored using antibody-conjugated receptors through molecular pendulum bioanalysis.

[0044] FIG. 34 shows that different targets, from small molecules to large-size proteins, are continuously monitored using the active reset approach.

[0045] FIG. 35 A and 35B show MD simulation of pendulum oscillation with two different energies. Evaluation of various covalent (angle, dihedral, and stretch) and non- covalent (Electrostatic and VdW) energies of (FIG. 35A) P2 and (FIG. 35B) protein for high energy oscillation (solid line, voltage = 0.5 V, frequency = 1 / (5 ns)) and low energy oscillation (dotted line, voltage = 0.3 V, frequency = 1 / (20 ns)) cases showing that applying oscillation with higher energy enhances the target protein dissociation. This is confirmed by the P2 and protein energy, water hydrogen bonding, and RMSD measurements.

[0046] FIGS. 36A and 36B show a P2-water H-bond comparison: MD simulation of probe oscillation with two different energies. P2 -water h-bond (FIG. 36A) and protein-water h-bond (FIG. 36B) for high energy oscillation (solid line, voltage = 0.5 V, frequency = 1 / (5 ns)) and low energy oscillation (dotted line, voltage = 0.3 V, frequency = 1 / (20 ns)) cases. The results show that the number of water h-bond for both P2 and protein is higher in the high-energy cases with respect to low-energy ones, indicating a higher protein dissociation rate.

[0047] FIGS. 37A and 37B show a RMSD comparison: MD simulation of probe oscillation with two different energies. RMSD of P2 strand (FIG. 37A) and protein (FIG. 37B) over time for high energy oscillation (solid line, voltage = 0.5 V, frequency = 1 / (5 ns)) and low energy oscillation (dotted line, voltage = 0.3 V, frequency = 1 / (20 ns)) cases, which shows how the P2 and protein becomes unstable over time for the higher energy system.

[0048] FIGS. 38A and 38B show a P2 SAS A and RMSF comparison: MD simulation of probe oscillation with two different energies. FIG. 38A shows solvent Accessible Surface Area (SASA) of P2. FIG. 38B shows RMSF of P2. The SASA and RMSF of P2 for high energy (black line, voltage=0.5V, frequency=l / (5ns)) is higher than the low energy (blue line, voltage=0.3V, frequency=l / 20ns).

[0049] FIG. 39 show the viscosity effect on sensor reset. When the viscosity of a solution is very high, it restricts the angular displacement of the pendulum, leading to decreasedregeneration performance compared to a low-viscosity solution (blue part). In a high- viscosity solution, the drag force is also high. By decreasing the frequency, we can increase the angular displacement of the pendulum, consequently increasing the inertia force, which enhances the regeneration performance (red and green parts).

[0050] FIG. 40 shows the calculation of probe density using chronocoulometry, Chronocoulometry of a probe-modified sensor in a buffer solution in the presence and absence of 100 pM of [Ru(NH3)6]3+, used for the determination of probe density.

[0051] FIG. 41. shows the Biolayer interferometry (BLI) assessment of IL-6 and TNF-a in PBS buffer (top) and ISF (bottom). The dissociation constants (KDS) of 5.45 nM and 51.6 nM for rat IL-6 and 5.22 nM and 40.7 nM for rat TNF-a in PBS and ISF are calculated, respectively.

[0052] FIG. 42A and 42B shows IL-6 (FIG. 42A) and TNF-oc (FIG. 42B) in vitro continuous monitoring with raw current changes and error analysis. Data points shown are plateau values for each concentration. Error bars represent standard deviations.

[0053] FIGS. 43 A, 43B, 43C, 43D, 43E, and 43F show chronoamperometry results from animal studies. The level of IL-6 and TNF-a cytokines in the rat dermal ISF is measured in three scenarios 1(F1GS. 43A and 43B)) fasting, 2 (FIGS. 43C and 43D)) fasting+insulin, and 3(FIGS. 45E and 45F)) fasting+LPS.

[0054] FIG. 44 shows MeHA microneedle patch administration on the skin of the rat. The MeHA patch is applied on the rat skin and remained for 10 minutes to absorb dermal ISF.

[0055] FIGS. 45A and 45B show a comparison of active reset and no reset signals for detecting IL-6 and TNF-a in diabetic rats undergoing fasting.

[0056] FIGS. 46A, 46B, 46C, 46D, 46E, and 46F show a Bland-Altman analysis of animal study results. The electrochemical and ELISA measurements using microdevice and MeHA patch have been compared at different time points, representing that there is anagreement between two data sets in three scenarios 1 (FIG. 46A and 46B)) fasting, 2(FIG.46C and 46D)) fasting+insulin, and 3(FIG. 46E and 46F)) fasting+LPS.

[0057] FIG. 47 shows the effect of LPS injection: Protein change rate of the microdevice compared with ELISA results obtained from serum and extracted ISF for the fasting-only, fasting+insulin, and fasting+LPS models. LPS injection causes an increase in IL-6 and TNF- a cytokines in ex vivo (ELISA) and in vivo (microdevice) measurements (inducing inflammation behavior).DETAILED DESCRIPTION

[0058] The present application incorporates by reference in its entirety the PCT application “MONOLAYER TRANSPORTER BIOSENSOR SYSTEM” filed March 10, 2025. That PCT application claims priority to U.S. provisional application No.63 / 563,658 filed March 11, 2024, which is also incorporated herein by reference in its entirety.

[0059] Provided herein is a method to actively reset a biosensor system, along with devices using said active reset method. The active reset mechanism relies on rapid oscillation of the biosensor molecules of the biosensor system to facilitate dissociation of the bound analyte from the analyte receptor. This oscillation is induced by changes in the electric current and applied electric field generated by the electrode.

[0060] In a first aspect, the present disclosure provides a method actively resetting the immobilized biosensor probes in a biosensor system comprising rapidly oscillating the biosensor probes by inducing a change in electric current and / or applying alternating electric field to the biosensor probes, wherein the biosensor probe comprises an analyte receptor that can bind to the analyte and wherein rapid oscillation of the biosensor probe results in the release and fast dissociation of an analyte bound by the biosensor probe. In some embodiments of this first aspect, the change in electric current or electric field is a change of up to 0.36V. In some embodiments of this first aspect, the frequency of the rapid oscillation is between 50 and 95 Hz. In some embodiments of this first aspect, the biosensor system comprises (a) an electrode, and (b) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminusand a second terminus and the analyte receptor attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure. In some embodiments of this first aspect, the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA). In some embodiments of this first aspect, the electric current and / or applied alternating electric field is applied to the biosensor probe via the electrode of the biosensor system. In some embodiments of the first aspect, each biosensor probe further comprises a redox reporter attached to the first terminus of the negatively-charged rigid rod structure. In some embodiments, the analyte receptor of the biosensor probe can bind a small molecule, a protein, or other biological entity. In some embodiments of the first aspect, the surface of the electrode is further bound by a chemical monolayer, and the analyte receptor binds to an electroactive analyte. In some further embodiments of this first aspect, the analyte is selected from an ion, a small molecule, a protein, or other biological entity that is electroactive or comprises electroactive moieties.

[0061] In a second aspect, the present disclosure provides a method of detecting an analyte using a biosensor system, comprising allowing a biosensor probe to contact a biological sample that may contain the analyte, wherein the immobilized biosensor probe is configured to bend and transfer an electron with the electrode in response to a change in electric current and / or an applied alternating electric field and, after a first detecting step is performed, applying a change in electric current or alternating electric field to the biosensor probe to cause rapid oscillation to shake off the captured analyte before performing subsequent detection steps with the biosensor probe. In some embodiments of this second aspect, the change in electric current or alternating electric field is a change of up to 0.36V. In some embodiments of this second aspect, the frequency of the rapid oscillation is between 50 and 95 Hz. In some embodiments of this second aspect, the biosensor system comprises (a) an electrode, and (b) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus and an analyte receptor attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rodstructure. In some embodiments of this first aspect, the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA). In some embodiments of this second aspect, the electric current and / or applied electric field is applied to the biosensor probe via the electrode of the biosensor system. In some embodiments of the second aspect, each biosensor probe further comprises a redox reporter attached to the first terminus of the negatively-charged rigid rod structure. In some embodiments, the analyte receptor of the biosensor probe can bind a small molecule, a protein, or other biological entity. In some embodiments of the second aspect, the surface of the electrode is further bound by a chemical monolayer, and the analyte receptor binds to an electroactive analyte. In some further embodiments of this second aspect, the analyte is selected from an ion, a small molecule, a protein, or other biological entity that is electroactive or comprises electroactive moieties.

[0062] In a third aspect, the present disclosure provides a method of determining a concentration of an analyte, the method comprising (a) applying a biological sample to a biosensor system, wherein the biosensor system comprises: (i) an electrode, and (ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus and an analyte receptor attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively- charged rigid rod structure, wherein the analyte present in the biological sample binds to the analyte receptors of the plurality of the biosensor probes; (b) applying an electric current to the biological sample and the plurality of the biosensor probes via the electrode of the biosensor system, wherein the biosensor probes and analyte bound by the analyte receptors attached to the biosensor probes are displaced toward the surface of the electrode; (c) measuring the chronoamperometric response of the electric current by translating the difference between an unbound electron transfer rate and a bound electron transfer rate; (d) calculating the concentration of the analyte in the biological sample based on a chronoamperometric response based on difference between the electrochemical signal produced by the biosensor probes that have bound the analyte and by biosensor probes that have not bound the analyte; and (e) actively resetting the biosensor probe by inducing rapid oscillation of the biosensor probes via the electrodeusing changes in electric current and / or applied electric field. In some embodiments of this third aspect, the method further comprises repeating steps (a) through (f) at least once. In some embodiments of the third aspect, the change in electric current or alternating electric field applied to the biosensor probes to induce rapid oscillation is a change up to 0.36V. In some embodiments of the third aspect, the frequency of rapid oscillation is between 50 to 95 Hz. In some embodiments of this third aspect, the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA). In some embodiments of the third aspect, the biosensor probe further comprises a redox reporter attached to the first terminus of the negatively-charged rigid rod structure, and the analyte receptor of the biosensor probe can bind a small molecule, a protein, or other biological entity. In some further embodiments of third aspect, upon application of the electric current, the electrode induces an electron transfer reaction with the redox reporter as the biosensor probes are displaced toward the surface of the electrode. In further embodiments of this third aspect, the chronoamperometric response is dependent on a time rate at which the biosensor probes are displaced. In some further embodiments of the third aspect, the unbound electron transfer rate is dependent on a time rate at which biosensor probes that have not bound an analyte at the analyte receptor are displaced. In some further embodiments of the third aspect, the bound electron transfer rate is dependent on a time rate at which biosensor probes that have bound an analyte at the analyte receptor are displaced. In some further embodiments of the third aspect, upon application of the electric current, the redox reporter approaches the electrode and the electron transfer is based on a redox reaction or electron tunneling current. In some embodiments of this third aspect, the biosensor system further comprises a chemical monolayer bound to the surface of the electrode. In some further embodiments of this third aspect, the biosensor probes, the analyte bound by the analyte receptors, and the redox reporter attached to the biosensor probes are displaced through the chemical monolayer and toward the surface of the electrode. In some further embodiments of this third aspect, the redox reporter approaches the electrode, and the electron transfer rate is based on a redox reaction or electron tunneling current. In some embodiments of the third aspect, the biosensor system further comprises a chemical monolayer bound to the surface of the electrode, and the analyte is an electroactive analyte. In some furtherembodiments of the third aspect, the electroactive analyte is selected from an ion, a small molecule, a protein, or other biological entity that is electroactive or comprises electroactive moieties. In some further embodiments of the third aspect, upon application of the electric current, the biosensor probes and the analyte bound by the analyte receptors attached to the biosensor probes are displaced through the chemical monolayer and toward the surface of the electrode. In some further embodiments of the third aspect, the analyte bound by the analyte receptors and displaced toward the surface of the electrode through the chemical monolayer are oxidized by the electrode. In some further embodiments of this third aspect, the chronoamperometric response is dependent upon the change in electrochemical signal caused by oxidation of the analyte bound by the analyte receptors and displaced toward the surface of the electrode through the chemical monolayer. In some embodiments of the third aspect, the biological sample is from a human subject. In some further embodiments of this third aspect, the human biological sample comprises a sweat sample, saliva sample, interstitial fluid sample, or serum sample. In some embodiments of this third aspect, the method is used to continuously monitor the concentration of the analyte in a subject.

[0063] In a fourth aspect, the present disclosure provides a device for continuously detecting an analyte in a biological sample from a subject, the device comprising (a) a collector component for collecting the biological sample, wherein the collector component comprises a microneedle having three inlet channels that connect to a main channel in the interior of the microneedle; (b) a detection component attached to the collector component comprising (i) an electrode, and (ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus, wherein each biosensor probe comprises an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure and wherein each biosensor probe is bound to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure; (c) an electrochemical measurement component connected to the electrode of the detection component, wherein the electrochemical measurement component detects changes in the electric current at the surface of the electrode. In some embodiments of this fourth aspect, the main channel of the collector component introduces the biological sample to the electrode and biosensorprobes bound to the surface of the electrode via a side channel connecting the collector component and the detection component. In some embodiments of this fourth aspect, the microneedle of the collector component induces capillary force upon the biological sample from the subject upon insertion of the microneedle into the skin of the subject. In some embodiments of this fourth aspect, the size of the inlet channels of the microneedle are 300 pm x 400 pm. In some embodiments of the fourth aspect, the diameter of the main channel of the microneedle is 300 pm. In some embodiments of this fourth aspect, the electrochemical measurement component comprises a display that provides the results of the analyte detection. In some embodiments of the fourth aspect, the electrochemical measurement component is connected to a separate component that displays the results of the analyte detection. In some embodiments of the fourth aspect, the negatively-charged rigid rod structure comprises dsDNA. In some embodiments of the fourth aspect, the analyte receptor comprises an antibody, an antibody fragment, a nanobody, or an aptamer. In some embodiments of the fourth aspect, the analyte is selected form a small molecule, a protein, or other biological entity. In some embodiments of the fourth aspect, the detection component further comprises a chemical monolayer bound to the surface of the electrode. In further embodiments of this fourth aspect, the chemical monolayer comprises a self-assembling chemical monolayer. In further embodiments of this fourth aspect, the self-assembling chemical monolayer comprises 6-mercaptohexanol (MCH). In some embodiments of the fourth aspect, the subject is a human subject. In some embodiments of the fourth aspect, the biological sample is an interstitial fluid sample.

[0064] It is to be appreciated that certain aspects, modes, embodiments, variations and features of the present methods are described below in various levels of detail in order to provide a substantial understanding of the present technology. It is to be understood that the present disclosure is not limited to particular uses, methods, reagents, compounds, compositions or biological systems, which can, of course, vary. It is also to be understood that the terminology used herein for the purpose of describing particular embodiments only and is not intended to be limiting.I. Definitions

[0065] Unless defined otherwise, all technical and scientific terms used herein generally have the same meaning as commonly understood by one of ordinary skill in the art to which this technology belongs. As used in this specification and the appended claims, the singular forms “a,” “an,” and “the” include plural referents unless the context clearly dictates otherwise. For example, reference to a “a cell” includes a combination of two or more cells, and the like. Generally, the nomenclature used herein and the laboratory procedures in cell culture, molecular genetics, organic chemistry, analytical chemistry and nucleic acid chemistry and hybridization described below are those well-known and commonly employed in the art.

[0066] As used herein, the term “about” in reference to a number is generally taken to include numbers that fall within a range of 1%, 5%, or 10% in either direction (greater than or less than) of the number unless otherwise stated or otherwise evident from the context (except where such number would be less than 0% or exceed 100% of a possible value).

[0067] As used herein, the term “analyte” refers to a molecule that may be present in biological fluids. The analyte comprises the entity to be detected using the present biosensor. More particularly, the analyte can bind to the biosensor via a receptor and its detection can be performed by an electrochemical signal, as will be explained in more detail below. In some embodiments, the analyte can be small molecules, macromolecules, prokaryotic or eukaryotic cell-derived components such as nucleic acid material, proteins, viruses, bacteria, antibodies, or cellular extracts, or any combination thereof. In some embodiments, the analyte may be neurotransmitters, hormones, lipids, and drugs that have compatible oxidation potentials. The analytes may be endogenous targets or exogenous targets. In some embodiments, the analyte can include a virus. In some embodiments, the analyte can be an antibody that is specific to a virus. In some embodiments, the analyte can include a bacterium. In a preferred embodiment, the analyte is a drug, a small molecule, a protein, or other biological entity. For example, the target analytes can include TNF-a, IL-6, and MPO. In other embodiments, the analyte is an ion, a drug, a small molecule, a protein, or other biological entity that is electroactiveor includes electroactive moi eties. In some embodiments the analyte can include serotonin, dopamine, epinephrine, doxorubicin, melatonin, and insulin.

[0068] As used herein, the term “electroactive” can mean responsive or active when subjected to electric potentials.

[0069] As used herein, the term “small molecule” can refer to a natural or synthetic molecule having a molecular mass of less than about 900 Da.

[0070] As used herein, the term “macromolecule” can refer to, but is not limited to, a large molecule composed of thousands of covalently connected atoms such as carbohydrates, lipids, proteins, and nucleic acids. A macromolecule can be formed of repeating monomer units, forming a polymer. Macromolecules may also include non- polymeric large molecules, such as lipids (including phospholipids) and large macrocycles.

[0071] As used herein, the term “receptor” refers to a molecular entity that is capable of binding an analyte, i.e., the target analyte. The receptor can also be referred to as “recognition agent”, “recognition element” or “capture agent”. The receptor is different than the target analyte and complementary to the target analyte. The receptor can be bound to at least a portion of the target analyte when the sensor comes into contact with the target analyte. In some embodiments, the receptor can comprise antibodies, nanobodies, antigens, aptamers, molecular imprints, protein receptors, DNA, microorganisms or protein / enzyme substrates. In one embodiment, the receptor can be a protein, an antibody or an aptamer.

[0072] As used herein, the term “electrode” refers to any electrode or electrochemical system that is sufficient for detecting the change in electrochemical current when an electron transfer from the redox reporter to the electrode occurs, namely upon binding of the target analyte to the receptor. The electrode can be a nanostructured electrode, a non- nano structured electrode, a micro-patterned electrode, or an array thereof. In some embodiments, the electrode can be a glassy carbon electrode, a carbon nanotube-modified electrode, an indium tin oxide (ITO) electrode, a platinum electrode, a gold electrode, asilver electrode, or a palladium electrode. The electrode can be fabricated on solid substrates including glass and silicon, or on flexible substrates including polyester (PET), polyimide (PI), polyethylene naphthalate (PEN), polyetherimide (PEI), optionally along with various fluoropolymers (FEP), copolymers, and paper. In some embodiments, the electrode can be a gold nanostructured microelectrode. In some embodiments, the electrode can be in the form of a wire. The wire can for example have a diameter between about 0.05 to about 0.2 millimeter. In other embodiments, the wire electrode can be embedded within a matrix of a non-conductive material.

[0073] As used herein, the term “electrochemical signal” refers to the signal generated upon the oxidation of the analyte bound to the biosensor at the electrode surface (i.e., change in current). In some embodiments, the electrochemical signal can represent the change in electron transfer decay rate between bound and unbound target analyte following application of a step voltage to the electrode. In some embodiments, the change in electron transfer decay can be performed by chronoamperometry.

[0074] As used herein, the term “sample” means any sample comprising or being tested for the presence of one or more target analyte. In some embodiments, the sample can be any biological fluid (e.g., from the body of a mammal or any other animal), including but not limited to interstitial fluid, blood, plasma, serum, saliva, urine, tears, sweat, to name a few examples.II. Biosensor Systems and Biosensor Probes

[0075] Described herein are biosensor systems to be used in advanced electrochemical detection methods for proteins with different sizes and affinities, amongst other analytes. The biosensor systems may include a plurality of biosensor probes and an electrode. The biosensor systems are reagentless and require no further reagents for analyte detection other than the sample to be tested.

[0076] The biosensor probes may be molecular pendulum biosensor probes. Each molecular pendulum biosensor probe may comprise a negatively-charged rigid rod structure with a first terminus and a second terminus. Each molecular pendulumbiosensor probe may comprise an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure. Each molecular pendulum biosensor probe may be attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure. Each molecular pendulum biosensor probe may comprise a positively-charged rigid rod structure with a first terminus and a second terminus. Each molecular pendulum biosensor probe may comprise an analyte receptor and a redox reporter attached to the first terminus of the positively-charged rigid rod structure. Each molecular pendulum biosensor probe may be attached to the surface of the electrode at the second terminus of the positively-charged rigid rod structure.

[0077] The biosensor probes may be monolayer transporter biosensor probes. Each monolayer transporter biosensor probe may comprise a negatively-charged rigid rod structure with a first terminus and a second terminus. Each monolayer transporter biosensor probe may comprise an analyte receptor attached to the first terminus of the negatively-charged rigid rod structure. Each molecular pendulum biosensor probe may be attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure. Each monolayer transporter biosensor probe may comprise a positively-charged rigid rod structure with a first terminus and a second terminus. Each monolayer transporter biosensor probe may comprise an analyte receptor attached to the first terminus of the positively-charged rigid rod structure. Each molecular pendulum biosensor probe may be attached to the surface of the electrode at the second terminus of the positively-charged rigid rod structure. The electrode may comprise a chemical monolayer attached to the surface of the electrode. The chemical monolayer may be a self-assembling chemical monolayer. The self-assembling chemical monolayer may be 6- mercaptohexanol (MCH).

[0078] The negatively-charged rigid rod structure of the molecular pendulum biosensor probes or the monolayer transporter biosensor probes may be, but is not limited to, double-stranded DNA (dsDNA). The length of the negatively-charged rigid rod structure may be from about 5 nm to about 20 nm. Preferably, the length may be 8 nm. Where the rigid rod is dsDNA, the length can range from about 8 nucleotides to about 60 nucleotides. Preferably, the length may be 24 nucleotides.

[0079] The analyte receptor attached to the negatively-charged rigid rod structure of the molecular pendulum biosensor probe may be, but is not limited to, an antibody, a nanobody, an antigen, an aptamer, an aptamer fragment, a molecular imprint, a protein receptor, DNA, a microorganism, a protein / enzyme substrate, or any combination thereof. Preferably, the analyte receptor of the molecular pendulum biosensor probe may be an aptamer. The analyte receptor attached to the negatively-charged rigid rod structure of the monolayer transporter biosensor probe may be, but is not limited to, an antibody, an antibody fragment, an aptamer, or a nanobody. Preferably, the analyte receptor of the monolayer transporter biosensor probe may be aptamer.

[0080] The analyte detected by the analyte receptor attached to the molecular pendulum biosensor probe may be a small molecule, a macromolecule, a prokaryotic or eukaryotic cell-derived component (e.g., nucleic acid material), a virus, a bacterium, an antibody, a protein, a cellular extract, or any combination thereof. The analyte detected by the analyte receptor attached to the monolayer transporter biosensor probe may be an electroactive analyte. The electroactive analyte may be an ion, a small molecule, a protein, or other biological entity that is electroactive or comprises electroactive moieties. The analytes detected by the analyte receptors of either the molecular pendulum biosensor probes or monolayer transporter biosensor probes may be important physiological markers of disorders or diseases. The disorder or disease may be, but is not limited to, cancer, cardiovascular health, and inflammation.

[0081] The analyte may be present in a biological sample. The biological sample may be, but is not limited to, interstitial fluid, saliva, sweat, or serum. The biosensor probes may detect the target analyte and collect data in situ, such that the biosensor probe can be used in an environment where the analyte naturally exists. The biosensor probes may be used in living animals or in humans to enable a broad range of personalized health monitoring. The biosensor probes may be reusable and used for one, two, three, four, five, six, seven, eight, nine, ten, or more rounds of analyte detection.

[0082] The biosensor system may comprise a plurality of biosensor probes for detecting a single analyte. The biosensor system may comprise a plurality of biosensor probes fordetecting two or more analytes, wherein subsets of biosensor probes include different analyte receptors and optionally different rigid rod structures attached to the surface of the electrode.

[0083] The biosensor system may comprise a plurality of monolayer transporter biosensor probes for detecting a single analyte. The system may comprise a plurality of monolayer transporter biosensor probes for detecting two or more analytes, wherein subsets of monolayer transporter biosensor probes include different analyte receptors and optionally different negatively-charged rigid rod structures. The system may comprise a plurality of monolayer transporter biosensor probes for detecting two or more analytes, wherein subsets of the monolayer transporter biosensor probes include different analyte receptors and optionally different positively-charged rigid rod structures. The monolayer transporter biosensor probe density on the surface of the electrode may be approximately 4.3 ± 0.74 x 1012probes / cm2or 1 probe per 23 nm2.

[0084] The electrode may be a type of electrode known in the art and compatible with the nature of the sample to be tested. The electrode may allow for detection of the change in electrochemical potential when the analyte is bound by the analyte receptor of the biosensor probe. The electrode may be a nanostructure electrode, a non-nanostructure electrode, a micropatterned electrode, or a wire. Preferably, the electrode may be a gold microelectrode containing additional electrodeposited nanostructures, called gold nanostructured microelectrodes.

[0085] The biosensor probe system can be used in an in situ environment where the analyte naturally exists. The biosensor probe system may be used in living animals or in humans to enable a broad range of personalized health monitoring. The biosensor probe system may be reusable and used for one, two, three, four, five, six, seven, eight, nine, ten, or more rounds of analyte detection. The biosensor probe system requires no further reagents for detection other than the sample to be tested and is thus considered reagentless.

[0086] The biosensor probe system may operate by allowing the biosensor probes to bind to analytes in a sample via the attached analyte receptors. The sample may be a biologicalfluid sample collected from a subject. The biosensor probes may be subjected to an electric current from the electrode, which causes the biosensor probe to be displaced toward the surface of the electrode and through the chemical monolayer. The electric current may be a positive current when the biosensor probes include negatively-charged rigid rod structures. The electric current may be a negatively charged current when the biosensor probes include a positively-charged rigid rod structure. The concentration of the analyte present in the sample may be determined using the method disclosed below.III. Biosensor Detection Device

[0087] Provided herein is a device utilizing the biosensor probes and biosensor system.The system as described above may be incorporated into a device. The device may include a collection component for collecting a desired biological sample from a subject, a detection component attached to the collector component, which includes a plurality of the biosensor probes as described above and an electrode; and an electrochemical measurement component connected to the electrode of the detection component, which detects the changes in the electrical potential induced by activation of the electrode by the bound analyte.

[0088] The collection component and detection component may be designed to allow for contact between the biosensor probes and the collected sample. The collection component may be designed to at least partially encase the detection component. The collection component can be characterized to have a first portion and a second portion, where the first portion contains the collected sample and the second portion encases the detection component.

[0089] The first portion of the collection component may include a nozzle and a sampling reservoir. The nozzle can be patterned with microfluidic channels in order to initiate passive capillary fluid flow once the nozzle comes in contact with the sample.The nozzle can include microfluidic channels which are treated with a hydrophilic coating.

[0090] The collection component may include a microneedle component to collect a biological sample from the subject and deliver the sample to the embedded biosensor in the detection component. The microneedle may be hollow. The microneedle may comprise three inlets that are large enough to prevent needle blockage during sample collection but small enough to induce capillary force to draw the biological sample from the subject. The inlet nozzles may be approximately 300 pm * 400 pm. The inlets may attach to a main channel. The channel may have a diameter of around 300 pm. The electrodes and the embedded biosensor may have access to the main channel through a side channel in the detection component. The biological sample may be an interstitial fluid sample that may be drawn from the skin’s dermis layer. The microneedle may be designed in such a way as to protect the embedded biosensor in the detection component from tissue intervention during microneedle insertion and continuous monitoring. The microneedle may be inserted into the skin of a subject and may induce capillary force upon the biological sample, including, but not limited to, interstitial fluid upon insertion into the skin of the subject.

[0091] As mentioned above, the detection component of the device includes the biosensor system of the present technology and thus includes the biosensor probes connected to a first working electrode. The detection component may include further electrodes, such as two working electrodes (WE), a counter electrode (CE) and a reference electrode (RE). The two additional working electrodes can allow positive and negative controls to be determined within the same sensor from measuring at different working electrodes.

[0092] The electrode of the biosensor system and the two additional working electrodes may be made of gold. The counter electrode can be made of platinum and the reference electrode may be made of silver. However, each electrode can either be solid gold, platinum or silver metal, or a plating / coating on an inexpensive metal. In some embodiments, each electrode can be in the form of a wire.

[0093] The detection component of the device can have a cylindrical form with the electrode (e.g., electrode wires) positioned spaced apart within the sensing componentalong a length thereof. The cylinder forming the detection component can have a length L and a diameter D. The length L and the diameter D of the cylinder can be adjusted. The length L can be from about 700 pm to about 1000 pm (e.g., about 800 pm) and the diameter D can be from about 300 pm to about 500 pm, for instance about 400 pm. These values are provided only as examples and may be adapted depending on the whole design of the device. In the case where the detection component is in the form of a cylinder, the collection component can thus have at least a portion thereof having a cylindrical form to properly encase the detection component.

[0094] The detection component may also be designed such that the electrodes (e.g., electrode wires) are held in a matrix of a non-conductive material. Any non-conductive material commonly used in the field can be used to form the matrix holding the electrodes spaced apart. The non-conductive material can be a silicon resin. The non- conductive material can be a cured polyorganosiloxane, such as cured polydimethylsiloxane (PDMS). In the case where the electrodes are in the form of wires, the matrix can embed the wires which are parallelly positioned along the length of the cylinder, thereby forming a kind of sheath protecting the electrode wires.

[0095] The detection component can be designed such that biosensor probes of the biosensor system are in contact with the sample to be tested and which is collected in the collection component of the device. The biosensor probes are bound to a first end of the electrode wire, and the biosensor probes are exposed to the sample when the first end is in contact with the collected sample. Each electrode wire (biosensor, WE, RE and / or CE) is in contact with the sample, at a first end thereof, when the sample has been collected in the collection component. A second end of the wires can be in contact with a connector that connects the detection component with the electrochemical measurement component. The wires can extend from the extremity of the detection component's matrix opposite to the extremity in contact with the sample, to allow the electronic connection with the connector and thus the electrochemical measurement component.

[0096] The detection component can be connected to the electrochemical measurement component via a connector. The connector can be in the form of a cap, which can connectto the extremity of the detection component opposite the extremity in contact with the sample. The detection component, which can be encased within the second portion of the collection component, can be secured within this second portion of the collection component using a locking system. The collection component and the connector can be provided with complementary locking means to secure the detection component within the device. The locking means can be any known system for fastening the second portion of the collection component with the connector, for instance complementary screw threads, sealed joint, etc.

[0097] The extremity of the detection component opposite the extremity in contact with the sample can also be closed off by an adapter, which can adapt the electrode wire connections to a plug and cable, which can connect to a handheld device. The adapter can have holes to feed through the wire ends to larger metal terminals of the plug and can be fixed electrically using solder cup or crimp connections for instance.

[0098] The electrochemical measurement component may directly display the results of the analyte detection conducted by the detection component. The electrochemical measurement component may be connected to a separate readout component that receives the data from the electrochemical measurement component and displays the results of the analyte detection conducted by the detection component.

[0099] The different elements of the device can be fabricated using 3D printing.

[0100] The device can provide an alternative to PCR-based testing and could accelerate the availability of high-quality diagnostic information.

[0101] The device may be a disposable device that may be disposed of after one, two, three, or more rounds of analyte detection.

[0102] As can be noted, the biosensor and device described herein can be advantageously used for the detection of different types of analytes in a sample. Although examples of analytes that can be detected using the present biosensor probes have been provided herein, the technology is not limited to such analytes and the biosensor can be designedby varying the biosensor probes and analyte receptor to allow detection of a large variety of analytes.IV. Biosensor Detection Method

[0103] The analytes bound to the analyte receptor attached to the biosensor probe may be detected via the electron transfer rate difference between when the analyte is bound by the biosensor and when there is no analyte bound. Once the target analyte is bound to the receptor of the biosensor and in contact with the applied electric field or electric potential, an electrochemical signal can be produced which shows a difference between the electron transfer rate seen when no analyte is bound to the receptor and the electron transfer rate seen when the analyte is bound to the receptor.

[0104] One such method of electron transfer rate that may be used to detect the binding of the analyte is chronoamperometry. Chronoamperometry is a potential-stepping method used to observe the changes in current or electron transfer decay in the presence and absence of the analyte to calculate the concentration of the analyte in the sample.

[0105] For the molecular pendulum biosensor probes, this technique involves measuring an electrochemical signal produced following application of an electric potential to the molecular pendulum biosensor probes, which translates a difference between the electron transfer rate observed when no analyte is bound to the analyte receptor and the electron transfer rate where the analyte is bound to the analyte receptor. The electron transfer, either in an unbound or bound state, can be observed as the redox reporter and the molecular pendulum biosensor probe approach the electrode surface upon application of an electric potential by the electrode. The electron transfer rate may be dependent on the molecular pendulum biosensor probe motion time rate. In other words, the electron transfer rate can be dependent on the time rate at which the molecular pendulum biosensor probes are displaced towards the electrode surface. The electron transfer rate in an unbound state may be dependent on a time rate at which the unbound molecular pendulum biosensor probes are displaced. Similarly, the electron transfer rate in a bound state may be dependent on a time rate at which the bound molecular pendulum biosensor probes are displaced. Equations for the molecular pendulum biosensor probes’ motioncan be developed form first principles, and the characteristic time, T, equivalent to the electron transfer rate, k = 1 / T, can be expressed by taking a first order approximation as: T = 2^7. This equation can allow one to make an informed assessment for molecular pendulum biosensor probe designs based on the analyte detection of interest. To observe analytes at the protein level, it was found that the temporal response could lay in the experimentally accessible microsecond-to-milli second regime when the length of the linker was in the range of 3-10 nm. Longer probes were slower to respond and exhibited weaker responses since the field was decayed when the charged entity was above about 10 nm from the electrode surface. Since T is the first-order response of a time-invariant system, it was equated to the exponential decay of the field-induced electrochemical response of a monolayer of molecular pendulum biosensor probes in the form Aoe~1 / T. By this, the change in electron transfer decay rate between bound and unbound sensors by applying a step voltage to the electrode can be detected. The amplitude of the signal at r, corresponding to the current measurement at that time, can then be determined, making it easy to differentiate signals of bound and unbound sensors.

[0106] For the monolayer transporter biosensor probes, this technique involves stepping of the electric potential of the electrode and monitoring of the resulting current from faradaic process (i.e., either oxidation or reduction) occurring at the electrode as a function of time after applying the potential step. Thus, there would be a change in the current detected based on oxidation of the analyte transported across the chemical monolayer to the surface of the monolayer transporter electrode by the analyte receptor on the monolayer transporters. This change in current would be proportional to the amount of analyte transported across the monolayer, allowing for determination of the concentration of the analyte based on relative changes in the resulting current detected.V. Active Reset Mechanism

[0107] Provided herein is a method of actively resetting biosensor molecules to allow for subsequent rounds of accurate detection. Previous methods of detection failed to include an active reset mechanism and instead relied on passive reset of the biosensor molecules, which involves the spontaneous dissociation of the analyte from the analyte receptorbased on the analyte receptor’s affinity to the receptor. Development of an active reset mechanism allows for the utilization of more sensitive analyte receptors that have a higher affinity for the analyte as opposed to receptors with less affinity but shorter dissociation times, while not remaining dependent on passive dissociation of the complex between the analyte and analyte receptor. Utilization of the active reset mechanism in a biosensor system may allow for more accurate detection of analytes in a subject continuously, as it would reduce the level of false positive readings of bound analyte that was not dissociated from the analyte receptor from earlier rounds of detection. The active reset mechanism may be employed with biosensor molecules other than molecular pendulum biosensors and monolayer transporter biosensors that include redox reporters, including, but not limited to, structure-switching biosensors and aptasensors that include redox reporters.

[0108] The active reset mechanism may utilize the electrode of the biosensor system to drive dissociation of the analyte and the analyte receptor. The electrode may subject the biosensor molecules to rapid oscillation using changes in the electric potential and applied electric field. This rapid oscillation may depend on the voltage and frequency of the electric potential applied to the biosensor molecules. The voltage applied to the biosensor molecules must be lower than the voltage that would result in significant degradation of the biosensor molecules and / or the analyte receptors. The voltage applied to the biosensor molecules must also be lower than the voltage that would result in full oxidation of the bound redox reporter if such a redox reporter is attached to the biosensor molecule. The voltage may be between 0 to 0.36 volts (V). The frequency of the electric potential applied to the biosensor molecules may be between 20 to 95 hertz (Hz). More preferably, the frequency applied to the biosensor molecules may be between 50 and 95 Hz. The active reset mechanism may result in reset of the biosensor system within one minute.

[0109] The active reset mechanism may be utilized one, two, three, four, five, six, seven, eight, nine, ten, or more times without significant degradation of the biosensor molecules. The active reset mechanism may be employed between rounds of detection via the biosensor system in order to ensure more accurate detection of analytes in the sample.VI. Biosensor Applications

[0110] The molecular pendulum system and any other similar biosensor system may be useful in various applications requiring the detection of an analyte in a sample, including a sample of biological fluid. The biosensor systems can be used in disease monitoring. As described above, the biosensor systems may be incorporated into devices for use in continuous and unsupervised monitoring of biomarkers of diseases or disorders. The biosensor system may utilize the active reset mechanism described above to allow for accurate detection of analyte presence continuously in a subject. The biosensor system can be used in both in situ and in vivo detection applications.[0U1] The biosensor system can be incorporated into a wearable device for real-time analyte detection. The biosensor system can be utilized for a variety of applications in personalized health monitoring as well as disease applications. The biosensor system may be configured to detect small molecules, macromolecules, prokaryotic or eukaryotic cell- derived components, including nucleic acids, viruses, bacteria, antibodies, proteins, or cellular extracts in a sample. The biosensor system may be configured to detect various markers of different physiological states, such as stress, allergic responses, infections, inflammation, and cancer presence. The biosensor system may be utilized with various biological fluids, including, but not limited to, blood, saliva, urine, feces, tears, interstitial fluid, and sweat.EXAMPLES

[0112] The present technology is further illustrated by the following Examples, which should not be construed as limiting in any way.Example 1: Active-reset protein sensors enable continuous in vivo monitoring of inflammationIntroduction

[0113] The development of wearable and implantable sensors for health-related biomarkers has immense potential to revolutionize disease monitoring and prevention.The ability to continuously measure protein biomarker concentrations in bodily fluids would enable precise tracking of patients at risk of acute, rapid deterioration as well as those managing chronic diseases. Moreover, the ever-emerging role of inflammation in all aspects of human health indicates that real-time monitoring of this phenomenon could be pivotal in preventing many major health issues, including heart disease, diabetes, and even depression.

[0114] Previously reported body-interfaced (e.g., wearable, implantable) sensors are mainly restricted to monitoring of small molecules, such as electrolytes, drugs, neurotransmitters, hormones, and metabolites. Continuous protein monitoring has proven significantly more difficult for in vivo analysis given that many sensing strategies require reagents or reporter groups that are incompatible with body-localized sensing. While progress is being made in the area of reagentless protein sensing systems, existing systems remain incompatible with dynamic measurements of proteins in vivo. The detection of low levels of proteins present in physiological media requires the use of affinity receptors, such as antibodies and aptamers, that have long dissociation times, thus prohibiting real-time tracking of decreasing protein levels. For example, affinity receptors that exhibit KD values of 1 nM may have dissociation rates of 10'5s'1and complexation half-lives of around 20 hours. Thus, slow dissociation kinetics impede the ability of sensors to dynamically respond to and report on changes in protein levels. Given the clinical potential of protein monitoring for the management and prevention of chronic diseases, it is crucial to develop new methods for protein level determination in vivo that circumvent this challenge. This challenge necessitates the development of sensors capable of selectively capturing a target protein and then promptly releasing it upon detection.

[0115] Although prior studies have explored tracking protein biomarker levels in vivo including in excreted fluids like sweat, real-time monitoring of changes in live animals has remained an unmet need.

[0116] It was reasoned that to wait for complex dissociation to occur passively underlies the slowness of sensor response, but that taking an active approach, where an externalstimulus was applied to force dissociation, could break through this bottleneck. Activereset sensors are reported herein, which are designed to enable dynamic measurements (FIG. 1A). Specifically, development of a sensor reset mechanism is reported that accelerates analyte dissociation via potential-induced oscillation, thus forcing the bound proteins to detach from the receptors due to high shear stress.

[0117] This approach is showcased for continuous monitoring of protein biomarkers in interstitial fluid (ISF). A capillary-based microdevice was developed that delivers dermal ISF to integrated microelectrodes, enabling the electrochemical monitoring of biomarkers via reagentless molecular pendulum bioanalysis (FIG. IB and 1C). The recognition elements are highly specific and stable aptamers that can capture target proteins and enable their detection using hydrodynamic drag-mediated readout. This allows for the in vivo monitoring of dynamic changes in protein concentrations aligning with physiologically relevant timeframes. The combination of features reported successfully enables the real-time measurement of inflammatory cytokine levels in an animal model of diabetes.Biomolecular Analysis Approach

[0118] Applicant reasoned that the wait times for passive complex dissociation to occur underlies the slowness of sensor response; but that an active approach, where an external stimulus was applied to force dissociation, could break through this bottleneck. A reagentless molecular pendulum employing a tethered DNA sensor construct for temporarily Faradaic readout (FIG. 1A) was employed as the starting point for the development of an implantable sensor for continuous protein monitoring (FIG. IB). This construct includes a rigid, electrode-bound double-stranded DNA scaffold with a terminal ferrocene molecule (e- donor) and protein aptamer sequence. Aptamers were selected as receptors in this system owing to their programmable interactions, small physical size, high stability, and long shelf-life, making them ideal candidates for highly sensitive and selective in vivo protein monitoring. Application of +500 mV (vs. Ag) to the electrode causes the negatively-charged DNA construct to fall to the surface in an arc-like motion, where a ferrocene molecule attached to the complex is oxidized by the same potential.Protein detection is achieved through differences in electron transfer rate observed because protein-bound constructs travel to the electrode surface more slowly due to hydrodynamic drag and exhibit delayed oxidation. The pendulum sensing approach has previously been used for reagentless protein monitoring in situ, but given the slow time scale of passive reset, could not resolve changing protein concentrations in real time.

[0119] This sensing approach is reagentless as it does not require the addition of reporter groups or other additives and is therefore ideally suited for the development of implantable protein sensing devices (FIG. 1C). The aptamer portion of the sensor is a modular element that can be substituted to allow the development of different protein analytes (see Table 1 for aptamer sequences). High levels of sensitivity and specificity for a range of analytes can be achieved in simulated interstitial fluid(ISF) using this approach (FIGS. 2-5) that match the requirements of physiological monitoring. However, the high- affinity receptors that are used to achieve the needed levels of sensitivity are associated with long dissociation times, requiring the development of an active regeneration / reset mechanism to ensure that sensors can report on both increases and decreases in concentration.Development of Sensor Reset Approach for Continuous Protein Measurements

[0120] When a protein interacts with the aptamer-based receptor, the rate of electron transfer rate between the sensor-appended redox label and the electrode surface is slowed due to the altered hydrodynamic drag in molecular pendulum methodology. This electron transfer delay manifests as a decrease in current decay in chronoamperometry (FIG. 6A), allowing for the correlation of the change in current at a given timepoint (Al) with protein concentration. The measurement of sensor response when three aptamers with varying dissociation constants (KD, see Table 1 for aptamer sequences and sources) for a model analyte (myeloperoxidase) of 27 nM (low affinity), 2 nM (moderate affinity), and 160 pM (high affinity), illustrates the need for an active reset mechanism (FIG. 6B). When a protein-containing solution is substituted with buffer (passive reset), the signal changes observed among aptamers are 70% for the low affinity aptamer, 30% for the moderateaffinity aptamer, and 5% for the high affinity aptamer after 30 minutes of incubation.Despite the lack of protein target in solution, the sensor with pM affinity to the protein is unable to report on this pronounced concentration change. Given that many protein biomarkers are present in the picomolar range, the use of high affinity receptors is critical, and the low levels of dissociation observed presents a significant problem for development of continuous monitoring capabilities.Table 1. Sequence of Aptamers Used as Biorecognition Elements

[0121] While programming receptors for higher dissociation rates is one method of reducing dissociation times, it also runs the risk of adversely affecting receptor sensitivity and selectivity. Here is presented a more active form of analyte dissociation and sensor reset using an oscillation-based approach. Without wishing to be bound by any particular theory, it was hypothesized that by inducing rapid oscillation of the sensor using changes in potential and applied electric field, that the drag force induced might cause the target protein to rapidly dissociate when concentrations dropped below KD. It was observed thatthe application of + 0.3 V with 95 Hz frequency was able to regenerate and essentially reset the sensor in less than one minute (FIG. 6C). Multiple cycles of reset can be used without significant degradation of the sensor (FIGS. 7 and 8).

[0122] Quartz crystal microbalance (QCM) measurements validate this reset, with Af values post-reset matching PBS-only values, indicating complete analyte displacement (FIG. 61 and FIG. 28). Negative alternating potentials have no significant effect on the bound signal (FIG. 29). The reset methodology is versatile for aptamers with various affinities by adjusting voltage and frequency (FIG. 30), with multiple reset cycles feasible without significant degradation (FIGS. 7, 8, 9, 31, and 32, and 35). FIG. 6D shows that the extent of reset or dissociation can be controlled by varying voltage and frequency. There is also a strong correlation between aptamer affinity and the reset parameters, highlighting the method’s adaptability for different aptamer-target complexes (FIG. 6J).

[0123] Using this sensor reset approach, we challenged the platform to monitor MPO for over 700 minutes, with concentrations from 10 pg / ml to 1 ng / ml (FIG. 6K). These studies indicated a high level of reproducibility as determined through statistical analysis of separate data points collected at the same concentration. This data set also demonstrates the effectiveness of the active reset approach when tracking increasing and decreasing concentrations.

[0124] The universality of the method was also investigated by applying the active-reset methodology to various target molecules of different sizes, sensing approaches, and receptor types (FIGS. 6L and 6M, 33, and 34). The results confirmed that the active-reset methodology is applicable to a wide range of sensors, utilizing either aptamers or antibodies as receptor elements. This approach successfully targeted a variety of analytes, including small molecules like serotonin, peptide-based hormones like insulin, as well as larger proteins such as MPO.

[0125] Next, a range of applied voltages and frequencies were investigated to map the conditions that could be used for sensor reset (FIGS. 6D and 6E). Higher frequencies and voltages facilitated the dissociation of the aptamer-protein complexes and with the exact conditions depending on the affinity of the aptamer for its target protein. Higher affinityaptamers appear to require stronger drag forces, which can be achieved through higher frequency or voltage. Given that ferrocene is fully oxidized at voltages over + 0.375 V vs. Ag, the potential was limited to 0.36 V to avoid damaging the redox label during regeneration (FIG. 6D). Moreover, excessively high frequencies can compromise sensor functionality, and it was observed via square wave voltammetry (SWV) that oscillation frequencies below 100 Hz ensure sensor integrity and safety (FIG. 9). Using this sensor reset approach, the platform was challenged for over 180 minutes, with concentrations from 1 pM to 1 nM (FIG. 6F), where the ability to rapidly detect increases or decreases in the analyte was demonstrated.

[0126] Using molecular dynamics (MD) simulations, potential mechanisms driving complex dissociation and sensor reset were investigated (FIG. 6G), along with monitoring of the structural details of the aptamer-protein interaction and investigating how oscillating potentials can affect complexation (FIG. 6H, FIGS. 10-16). FIGS. 6H(i) and 6H(ii) report on the quantity of non-covalent interactions between the aptamer and protein for scenarios with and without oscillation, with hydrogen bonding being the dominant mode of interaction.

[0127] The simulation data reveals a significant observation: the application of oscillation (voltage 0.3V and frequency l / ( 5ns)) has no discernable impact on the Rg, RMSD, and RMSF values for the Pl aptamer, as depicted in FIGS. 11A-11C. This is in stark contrast to the behavior of the P2 aptamer. The key distinguishing factor is that the Pl aptamer possess a double helical structure, a structural element known for its stability. Consequently, the Pl aptamer appears to demonstrate greater resistance to the perturbations induced by oscillation. While the P2 aptamer experiences notable changes in its dynamics and interactions due to oscillation, the Pl aptamer remains comparatively stable.

[0128] The increase in Rg values of the P2 aptamer, show in FIG. 1 ID, along with the earlier findings of increased flexibility and fluctuation (as indicated by RMSD and RMSF), collectively suggest that the introduction of oscillation has a destabilizing effect on the P2 aptamer and its interactions with the protein. This reduced stability implies agreater propensity for structural fluctuations and dynamic changes within the complex. The introduction of oscillation significantly enhances the mobility and flexibility of the P2 aptamer, indicated by increased RMSD values, as demonstrated in FIG. 1 IE.

[0129] Simultaneously, the elevated RMSD and RMSF values for the protein suggest that its overall structural stability is compromised by the presence of oscillation (FIGS. 12A- 12B). The increase of Rg for the protein indicates a greater spatial distribution or expansion of its mass, signifying increased flexibility or conformational freedom (FIG. 12C). Therefore, the documented increase in hydrogen bonding between the protein / P2 and water (FIG. 6H(v) and 6H(vi)), alongside the dynamic structural changes in both the protein and P2 aptamer in the presence of oscillation, collectively suggest a significant alteration in their mutual interactions. It is reasonable to infer that these changes correspond to a decrease in affinity between the protein and the P2 aptamer, which was further confirmed by RDF computations (FIGS. 13A-13C).

[0130] FIG. 14 provides an insightful depiction of the interaction modes between the P2 aptamer and the protein in the absence of oscillation, utilizing cluster analysis. These figures reveal that, without oscillation, there is a robust network of molecular interactions between the aptamer and the protein. Specifically, the analysis identifies the formation of 12 hydrogen bonds and 2 salt bridges, underscoring the strong non-covalent connections between the two entities. However, the scenario changes markedly when oscillation is introduced. In its presence, the landscape of non-covalent interactions is substantially altered, leading to a notable reduction in these interactions from 12 to 6 (FIGS. 14 and 15). This reduction suggests a decreased affinity and binding strength between the P2 aptamers and the protein under the influence of oscillation.

[0131] FIG. 16 provides snapshots of the gold-aptamer-protein system at various time points both in the absence and presence of oscillation. As seen in the figure, oscillation induces increased mobility and flexibility of the P2 aptamer.

[0132] In order to distinguish the effect of alternating voltage and frequency range on the sensor, MD simulations were also performed for two different systems with high energy (voltage=0.5V and frequency=l / (5 ns)) and low energy (voltage=0.3V andfrequency= 1 / (20 ns)) and compared the various energies of each molecule, number of water h-bond, and RMSD in the P2 strand and protein over 100 ns simulation (FIGS. 35- 37). The P2 and protein energy results show that the oscillation significantly affects the non-covalent electrostatic energy, which is shown to be higher in the high-energy system (solid line) with respect to the low-energy system (dot line).

[0133] Increasing the P2-water and protein-water H-bonds — one of the most important parameters in the structural understanding of biomolecules — indicates the dissociation of the protein-receptor complex. The water h-bonds of P2 and protein for a high-energy system are significantly lower than those for a low-energy system, and the number of h- bond differences between the two systems increases over time (FIG. 36).

[0134] Evaluating RMSD, which shows the stability of the molecules, shows that applying higher energy to the system enhances the destabilization of the protein and P2 with respect to lower energy. As is shown, the system becomes more unstable over time, indicating the dissociation of the target protein (FIG. 37).

[0135] When a receptor dislodges from the target analyte, the value of Solvent Accessible Surface Area (SASA) increases. Evaluating the SASA value for P2, it is clear that the water molecules have more access to the P2 in the higher energy system (FIG. 40A). In addition, the root mean square fluctuation (RMSF) value in P2 shows that applying higher alternating potential increases the fluctuation of P2. The increase in the fluctuation results in an increase in the amount of drag force and inertia force, consequently facilitating the analyte dissociation (FIG. 38B)

[0136] Notably, when the entire 100 ns simulation is taken into consideration, the average count for hydrogen bonds decreases when oscillations are introduced from 21.5 to 16.9 (FIG. 10). Additionally, the Solvent Accessible Surface Area (SASA) of the aptamer increases significantly with external oscillation (FIG. 6H(iii)), indicating increased exposure of the aptamer to water molecules that may outcompete proteins for binding sites. Analyzing the RMSF values in FIG. 6H(iv), it was observed that oscillation predominantly amplifies the mobility and fluctuations within the protein binding domain of the aptamer (residues 55 to 117). FIGS. 6H(v) and 6H(vi) further reveal that oscillationleads to an increase in hydrogen bonding between water molecules and the protein / P2, suggesting that oscillation may enhance protein / P2 solvation and decrease aptamer protein interactions. Taken together, the simulations performed support the idea that rapid oscillation of the protein-bound sensor complex is destabilizing, likely due to drag force and increased collisional frequencies with water molecules. Comparisons of high-energy (higher frequency / voltage) and low-energy (lower frequency / voltage) systems (FIGS. 35- 38) further align with observed experimental behavior.

[0001] To further understand the relation of drag and inertia forces on the sensor reset performance, we oscillated the sensor in a solution with significantly increased viscosity (FIG. 39). The higher viscosity leads to an increase in the drag force and a decrease in the inertial force as the sensor’s angular displacement decreases during oscillation. By decreasing frequency and consequently increasing the inertial force, levels of regeneration increased (FIG. 39). These experiments, combined with the observations of the MD simulations, indicate that the mechanism of sensor reset involves a variety of molecular forces imparted by the potential-induced oscillations.Microdevices for Electrochemical Measurements In Vivo

[0137] In order to test active-reset sensors for the intended application of continuous protein monitoring, an implantable device contacting the dermal interstitial fluid (ISF) was developed for electrochemical testing. ISF has emerged as a key biofluid for biomolecular analysis, owing to its minimally invasive sampling, rapid equilibration with surrounding tissues, and high compositional correlation with blood. However, despite its advantages, ISF has not gained widespread adoption in clinical applications dues to its often-complex extraction methodologies and small volumes available for ex vivo analysis. An alternative approach to ISF monitoring involves direct in situ analysis of interstitial analytes using sensor-integrated microneedles or microdevices that can be adapted to various form factors.

[0138] A microscale enclosure for wire-based electrodes was fabricated that featured three channels for working, reference, and counter electrodes and a main central channel featuring three inlets within the tip of the device that are internally connected for ISFflow (FIG. 17A and 18). ISF enters the microdevice via capillary flow through the inlets owing to its hydrophilicity and the overall design of the channels (FIGS. 18 and 19). Electrodes are embedded within the device, ensuring that the functionalized electrodes are protected during skin insertion (FIGS. 19 and 20).

[0139] The Young’s modulus of the printed needle is about 5GPa, which is sufficient for skin puncture. The ISF inlets were designed to be small enough to ensure sufficient capillary force for the collection of ISF and large enough to prevent blockage during insertion and measurement (FIG. 17B). Fluid movement in the dermis is driven by a descending hydrostatic gradient, ranging from positive capillary pressure (10.5 to 22.5 mmHg relative to atmospheric pressure) to negative pressure in the dermis (-1 to -4 mmHg). To assess that the device could draw ISF from the dermis through capillary force, the worst-case scenario for the fluid flow when inserted into the skin was simulated, assuming the dermis pressure to be -4 mmHg (FIG. 17C). The computational fluid dynamics simulation shows that the ISF fills the device within 5 ms, resulting in rapid immersion of the embedded electrodes in ISF (FIG. 17D). Fluorescent microscopy imaging of the device exposed to fluorescent simulated ISF confirmed complete filling of the inner channels (FIG. 17D, inset). Velocity profiles and pressure changes for each step of the filling process are shown in FIG. 21. ISF flow into the channels must be finely controlled to ensure consistent electrode wetting and repeatable sensing readout. Hence, the top part of the device was designed to expand immediately at 90 degrees, acting as a capillary stop valve (FIG. 22). Biocompatability of the microdevice and the materials used for fabrication were tested, and an insignificant change in cell viability in 48 hours with very low cytotoxicity was observed (FIG. 23).

[0140] Protein concentration within the microdevices undergoes diffusion-based changes, equilibrating over time with exterior ISF (FIG. 17E and 24). The time required for equilibration depends on the magnitude of the concentration change. For instance, a transition from 1,000 pM to 500 pM requires approximately 20 minutes (FIG. 17E).Continuous Protein Monitoring of Inflammatory Markers in Diabetic Rats

[0141] In the diabetic population, the inflammatory cytokines IL-6 and TNF-a are known to be markers of vascular inflammation, a leading cause of morbidity and mortality for these patients. Methods that monitor inflammatory biomarkers would enable management of chronic diseases like diabetes and could be used as a means to control complications. Having a means to monitor inflammatory biomarkers would open new possibilities for the management of chronic diseases like diabetes and could be used as a means to control complications. These two cytokines were therefore used as a test case for continuous protein monitoring in a rat model of diabetes.

[0142] To apply the microdevices for testing for inflammatory cytokines in living animals, sensors were developed for IL-6 and TNF-a (FIG. 2-4). High affinity aptamers were used for both analytes that enabled the realization of pg / mL detection limits, which corresponds to the physiological levels of these proteins. The sensor reset strategy was shown to be effective with both sensors for more than 350 minutes, as shown in FIG. 25A, 25B, and FIG. 42 which were performed in simulated ISF in vitro.

[0143] For the in vivo study, the fabricated microdevice, integrated with electrodes, was administrated on the dorsal skin of the rat. To verify the penetration through the epidermis and insertion into the dermal layer, the puncture site of the microdevice was analyzed, and Haematoxylin and eosin (H&E) staining of the extracted rat skin tissue was performed (FIG. 17F). To evaluate systemic toxicity, we isolated vital organs (heart, liver, lungs, kidneys) and performed H&E staining to assess immune cell infiltration post-administration. No significant differences or histological abnormalities were found between tissues from control rats and those implanted with microdevices (Fig. 17G).

[0144] An assessment was also performed on whether the microdevice or its oscillations triggered localized inflammation by analyzing over 60 genes related to inflammatory pathways, including adhesion molecules, TNF receptors, and prostaglandin metabolism. No significant changes in gene expression were observed in skin tissues from control, pre-oscillation, and post-oscillation groups (FIG. 17H).

[0145] Diabetic rats exhibit higher levels of pro-inflammatory cytokines than healthy rats due to hyperglycemia-induced cytokine release. Fasting leads to downregulation ofcytokines through improved metabolic health, leading to better control of glucose and insulin levels, inducing autophagy, increasing adiponectin and ghrelin release, and decreasing both reactive oxygen species and immune cell activity (FIG. 25C).

[0146] In diabetes, pro-inflammatory cytokines are elevated compared to healthy rats due to hyperglycemia-induced cytokine release. Diabetes is characterized by impaired insulin secretion and sensitivity, which is exacerbated by oxidative stress, endoplasmic reticulum stress, pancreatic amyloid deposition, and ectopic lipid accumulation, all of which can trigger or worsen inflammation . In type 1 diabetes mellitus (T1DM), inflammation in pancreatic islets (insulitis) progressively destroys insulin-producing P cells, resulting in lifelong insulin dependence for affected individuals, often children or adolescents. Fasting can reduce inflammation by downregulating cytokines, improving metabolic health, and enhancing glucose and insulin control. This includes inducing autophagy, increasing adiponectin and ghrelin, and decreasing reactive oxygen species and immune cell activity (FIG. 25C). An animal study was conducted directly observe fasting effects on diabetic rats, using IL-6 and TNF-a as inflammatory markers, and investigate insulin and Lipopolysaccharide (LPS) injection effects (FIG. 25D). Sensor biofouling was assessed with and without BSA and found no significant difference. All in vivo measurements included a negative control to monitor sensor drift and biofouling.

[0147] Using implanted sensors, IL-6 and TNF-a in ISF were measured every 20 minutes in diabetic rats dorsal skin (Fig. 25E and FIG. 43). Blood samples were collected every 40 minutes for ELISA confirmation. Dermal ISF was also extracted every 40 minutes with a hydrogel microneedle patch, and cytokine levels were measured by ELISA (FIG. 44). IL-6 in fasting diabetic rats decreased from -520 pg / ml to -190 pg / ml (FIG. 25E). Insulin-injected rats showed a temporary inflammatory response at 60 minutes due to tissue trauma but then displayed a similar IL-6 decrease as untreated rats. Similar trends were observed for TNF-a (FIG. 25E).

[0148] LPS injection triggers an immune response, causing pro-inflammatory cytokine release. The microdevice’s ability to monitor cytokine changes was tested under fastingand LPS injection conditions (FIG. 25E). Comparing sensors with and without activereset shows that active-reset effectively tracks sharp protein level changes (FIG. 45).

[0149] The cytokine levels measured by the microdevice align with offline ELISA results of ISF, and consistent trends are observed between electrochemical and ELISA measurements for ISF and serum (FIG. 25E). Bland-Altman analysis confirms agreement between in vivo and offline ELISA measurements (FIG. 46).

[0150] The reduction of inflammation in the fasting-only, fasting + insulin, and fasting+LPS rats were observed to have different rates. The rates of cytokine level decrease were measured, and it was found that insulin injection enhances the lowering of cytokine levels, likely due to improved regulation of glucose levels (FIGS.26-27). This insulin-enhancing cytokine reduction was more evident in ISF than in plasma. Insulin accelerated cytokine reduction and LPS exacerbated inflammation (FIGS. 26, 27, and47). These results indicate that active-reset sensors can successfully track inflammatory biomarkers in vivo.

[0151] In order to show the versatility of the active reset methodology, in addition to the real-time monitoring of the MPO, IL-6, and TNF-a using aptamer-based sensors (FIGS. 6K, 25A and 25B), the application of the active-reset methodology across various biomarkers with different sizes and affinities and used several different types of sensor architectures was also investigated to show the generality of the approach, including monolayer transporter biosensors, antibody-based molecular pendulums, and structureswitching EAB biosensors (FIGS. 6L and 6M, 33 and 34).

[0152] To showcase the applicability of the reset methodology across different sensing mechanisms, a common approach known as the structure-switching electrochemical aptamer-based (EAB) sensor was utilized. This method involves a single-stranded aptamer tethered to the electrode surface, with a redox reporting molecule (methylene blue) attached at its distal end. Binding-induced conformational changes in the aptamer alter the rate of electron transfer from the redox reporter, thereby changing the peak current observed during sensor interrogation using square-wave voltammetry. By employing this EAB sensor, we successfully detected and reset the doxorubicin molecule,demonstrating that the active-reset method can be extended to all types of aptasensors. It is noteworthy that for the EAB sensors, active reset was performed with positive potentials, but redox signals were collected in the negative range.

[0153] The active-reset approach's efficacy in detaching antibodies and antigens was investigated. Instead of using an aptamer as a receptor, we conjugated anti-IL-6 and insulin antibodies to the pendulum and monitored the IL-6 and insulin proteins, respectively, in real-time using the active-reset methodology. FIGS. 6M and 33-illustrate the results of real-time detection for the mentioned analytes. It is worth mentioning that the calculated probe density shows that the DNA molecules have sufficient space to oscillate appropriately for all sensing methodologies (FIG. 40).Conclusion

[0154] Herein is described an approach enabling real-time, continuous monitoring of protein biomarkers in ISF. While previous efforts in wearable and implantable sensors have focused primarily on small molecules, this device offers a novel solution for protein biomarker analysis, overcoming the limitations of slow dissociation times associated with affinity receptors. By employing oscillation-based active sensor reset and reagentless molecular pendulum technology, this device facilitates real-time measurements with high sensitivity and specificity. The active sensor reset mechanism facilitates return to the unbound state within one minute, allowing for continuous protein monitoring. The versatility of this platform is demonstrated through using a wide range of sensor architectures for the successful monitoring of various protein and molecule biomarkers with different sizes and affinities, including cytokines, in an animal model of diabetes. This innovative technology opens new avenues for disease management, prevention, and exploration of protein biomarker dynamics in real-time, contributing to improved patient care and outcomes.Materials and Methods

[0155] Microdevice Fabrication. The microdevice fabrication process began with the initial design in CAD software, followed by the generation of Stereolithography (STL)files. These files were then transformed into Direct Laser Writing commands using the Describe software by Nanoscribe GmbH in Karlsruhe, Germany.

[0156] For the substrate material, single crystal Si (100) wagers were used due to their cost-effectiveness and commercial available, highly polished surfaces. These Si wafers were meticulously prepared by cutting them into 25 x 25 mm square slides, subjecting them to Piranha cleaning, and subsequently rinsing them with acetone and isopropyl alcohol (IP A) to ensure a pristine surface. The actual 2-Photon Polymerization (2PP) 3D laser writing process was executed using the Photonic Professional GT system, also from Nanoscribe GmbH, in a Dip-in Laser Lithography (DiLL) configuration with a 10X objective lens.

[0157] To initiate the process, the proprietary IP-Q photoresist was drop-casted onto the substrate, and the microscope objective was carefully immersed into the photoresist. The 2PP writing commenced at a predefined out-of-plane z-position of 1 pm within the substrate. This depth was chosen to ensure secure anchoring and to account for any potential slight tilting of the substrate. Additional writing parameters such as writing speed (40,000 pm / s), laser intensity (35% at the interface, 100% at scaffolds, and 100% at contours), and hatching distance (0.75 pm) were determined through multiple trial runs. The stage was then manipulated using a piezo-motor, and the laser was employed to scan the field of view of the objective in galvo-scan mode. Given that the laser’s field of vision was insufficient to cover the entire structure, the structure was divided into block regions with axes set at x = 999.4 pm, y = 999.4 pm, and z = 300 pm. The stage was used to transition between these regions, enabling separate writing with a 5 pm overlap for effective stitching.

[0158] Following the completion of 2PP writing, the samples underwent a 30-minute immersion in a propylene glycol monomethyl ether acetate (PGMEA) solution, followed by cleaning in IPA baths lasting 10 minutes each. The IPA baths were essential for removing uncured photoresist from an annular cavity in the tip of the microdevice. Subsequently, the printed microdevice was detached from the substrate and immersed once more in PGMEA, followed by IPA baths lasting 1 hour and 30 minutes,respectively, to ensure that the features at the base of the needles were free of uncured resin. Finally, the microdevice was fully cured under UV for 3 minutes.

[0159] Biocompatibility and Vitro Cytotoxicity Assay. The assessment of cytotoxicity for the printed microdevice material involved conducting a 3-(4,5)-dimethylthiazol(-z- yl)-3,5-diphenyltetrazolium bromide (MTT) assay utilizing NIH3T3 fibroblast cells obtained from ATCC. In brief, these cells were seeded into a 96-well plate at a density of 10,000 cells per well. After an initial 24-hour incubation period in 200 pL of Dulbecco’s Modified Eagle Medium (DMEM) supplemented with 10% fetal bovine growth serum (FBS), the test samples were introduced into the respective wells.

[0160] Subsequently, the cells were left to incubate for an additional 24 hours posttreatment. Following this incubation, 100 pL of MTT reagent was added to each of the treated wells, and the cells were incubated for an additional 4 hours. In the subsequent step a solution comprising 10% SDS in 0.01 M HC1, totaling 100 pL, was introduced into each well, and the mixture was incubated for an additional 4 hours.

[0161] Finally, the absorbance of the plate was measured at 550 nm using a BioRad UV- Vis plate reader. To calculate the percentage of cell survival, the following equation was employed and subsequently plotted on a semi-logarithmic scale. sample signal — background signal% Survival = - ; — - - — ■ — - : — - ; X 100 control signal — background signal

[0162] Synthesis of methacrylated hyaluronic acid (Control). Methacrylated hyaluronic acid (MeHA) was prepared using established methods detailed in previous research with some modifications 50). Initially, 0.5 g of hyaluronic acid (HA) was dissolved in 25 m of distilled-deionized (DDI) water. Methacrylic anhydride (MA), 40 pL, was then added dropwise to this solution. The pH of the mixture was adjusted to 9 using 5M NaOH. The reaction mixture was kept at 4 °C and stirred for 24 hours. Following this, 0.5 M NaCl was added, and the product was precipitated using 50 mL of acetone. The precipitate was re-dissolved in DDI water and purified through dialysis against DDI water for 3 days. The final product, obtained as an off-white powder with an 85% yield, was characterized by 'H-NMR. spectroscopy. The 'H-NMR (300 MHz, D2O, 8) readings were: 1.85-1.93[CH2 = C(CH3) CO], 1.95 [NHC0CH3], 5.68 [CH1H2 = C(CH3) CO], and 6.19[CH1H2 = C(CH3) CO],

[0163] Fabrication of crosslinked MeHA microneedle patches. The stainless steel master mold of MN, featuring a 300 pm base diameter, a 5 pm tip radius, and a height of 1000 pm, was sourced from Micropoint Technologies Pte Ltd in Singapore. A negative mold was created using poly dimethylsiloxane (PDMS, 10 mm thick, Dow Corning 184 Sylgard) to precisely replicate the master mold. This process involved pouring PDMS over the microneedle master structure, degassing it in a vacuum oven for 10 minutes, and curing the polymer at 70°C for 1 hour.

[0164] To fabricate the crosslinked MeHA-MN patch, MeHA (50 mg mL 'j and a photoinitiator (Irgacure 2959, 0.5 mg mL ') were dissolved in deionized water. The resulting mixture was poured into the plasma-treated PDMS mold, filling the cavities completely. The PDMS mold was then centrifuged at 4000 rpm for 3 minutes to ensure the material filled the needle voids. An additional solution was added to create a sturdy backing. After drying at room temperature in a fume hood for approximately 12 hours, the MeHA-MN patches were carefully extracted from the mold and trimmed. They were subsequently exposed to UV light (wavelength = 360 nm, intensity = 17.0 mW cm2, model 30, OAI) for 10 minutes. The optimal crosslinking for the MeHA-MN patches was achieved with 10 minutes of UV exposure.

[0165] Biosensor Preparation. A 1 pM solution of a thiolated probe, referred to as Pl, was prepared in 1 x PBS. This solution was combined with 1 mM Tris(2-carboxyethl) phosphine hydrochloride (TCEP) and incubated in darkness at room temperature for one hour. This incubation served to reduce the protective disulfide bond present in the probe.

[0166] The Pl probe is composed of a ferrocene redox reporter at the 3' end and a thiol group at the 5' end. Following the reduction step, the Pl mixture was briefly heated to 55 °C for 5 minutes and then allowed to cool to room temperature.

[0167] In parallel, P2, an aptamer containing a complementary sequence to Pl and a recognition element, was subjected to a heat treatment at 90°C for 3 minutes.Subsequently, it was introduced into the Pl solution and allowed to incubate at room temperature for 15 minutes, facilitating hybridization between the two probes. In the case of antibody-conjugated P2, the P2 solution was not preheated before being mixed with the Pl solution.

[0168] To this mixture, a solution containing MCH was added, achieving a final concentration of 9 pM. Next, the nanostructured gold rod electrodes with a diameter of 100 pm were submerged into this mixture and left to incubate overnight at room temperature in a dark and humidified environment to immobilize the DNA probes onto the electrodes.

[0169] After a 16-hour probe incubation, the electrodes were briefly rinsed with 1 x PBS. They then underwent a second incubation in 1 mM MCH for 1 hour. Following this, the electrodes were rinsed again with 1 x PBS and incubated with 1% BSA for 30 minutes. The resulting functionalized biosensors were thoroughly washed by immersing in 1 x PBS vials for 5 minutes, repeated twice, followed by a final wash in a vial containing O.lx PBS. Finally, the functionalized electrodes were integrated into the microdevice setups, alongside reference (Ag, 100 pm diameter) and counter (Pt, 250 pm diameter) rod electrodes.

[0170] Electrochemical Methods. Electrochemical analyses were conducted using a PGSTAT204 Potentiostat / Galvanostat instrument manufactured by Metrohm-Autolab in the Netherlands. Measurements were performed on rod-shaped electrodes, which consisted of a gold working electrode, a silver reference electrode, and a platinum auxiliary electrode. In an effort to enhance the sensor’s sensitivity, a procedure was implemented involving the growth of a thin layer of gold nanoneedles on the surface of the gold rod electrode. To achieve this, the procedure began by electrochemically cleaning the gold electrodes with a 0.5 M H2SO4 solution, utilizing cyclic voltammetry (CV). This cleaning process covered a potential range of 0 - 1.1 V (versus Ag / AgCl) and was executed at a scan rate of 100 mV / s for 30 cycles.

[0171] Following the thorough cleaning, the cleaned gold electrodes were decorated with gold (Au) nanoneedles. This was achieved through chronoamperometry (CA) in asolution composed of 0.1 mM L-cysteine, 0.1 M H2SO4, and 5 mM HAuCh. During this step, the electrodes were held at a constant potential of 0 mV for a duration of 100 seconds.

[0172] To further refine the sensor’s surface characteristics and nanocluster morphology, a second electrochemical deposition stage was initiated within the same solution. During this step, the potential was maintained at a constant -800 mV for a brief period of 5 seconds. This meticulous process resulted in the formation of a fine coating of gold nanoclusters on the gold electrode’s structure, achieving the desired surface area enhancement and nanocluster configuration conducive to improved sensor sensitivity.

[0173] To capture electrochemical signals, a rapid chronoamperometry (CA) technique was utilized with a potential range spanning from 0 to +500 mV relative to the silver reference electrode for a duration of 100 milliseconds. It is worth noting that the reporter electrochemical results represent the average of at least three repetitions to ensure accuracy and reliability.

[0174] For the purpose of regenerating the sensor, an oscillation procedure was employed characterized by repetitive potential changes (square-shaped alternating potential). This procedure encompassed a frequency range of 1-1000 Hz and a potential range of 0.15- 0.36 V. During this oscillation, the current was monitored as a key parameter for the measurements.

[0175] Biolayer interferometry measurement of protein binding. Biolayer interferometry (BLI) was utilized to verify the binding of rat IL-6 and TNF-a cytokines to their respective aptamer receptors. Streptavidin-coated BLI sensors served as anchors for biotin-labeled molecular pendulum probes. These probes were then tested against rat IL-6 and rat TNF-a in both PBS and simulated interstitial fluid (ISF).

[0176] Quartz crystal microbalance (QCM) measurements. The QCM experiment was conducted using a function generator linked to a QSense electrochemistry module. The setup included a three-electrode system: an Ag / AgCl reference electrode, a platinum counter electrode, and QCM chips with gold coating (0.785 cm2) serving as the workingelectrode. The functionalization of the electrode followed the same procedure used for the electrochemical sensor.

[0177] To investigate the mass transfer dynamics of the functionalized sensor, we monitored the resonant frequencies and dissipation of the functionalized piezoelectric crystal sensors in real-time. The QCM technique captures the resonant frequency shift, which is inversely proportional to mass changes on the sensor surface — a decrease in resonant frequency signifies an increase in mass. Due to the negative frequency shift observed in QCM measurements, a larger frequency difference (ft - fO) indicates a greater mass accumulation, as expressed by the Sauerbrey equation (Eq 1):

[0178] In this equation, Afsauerbrey.n represents the Sauerbrey frequency shift for the nth order overtone, and AM denotes the change in mass. The constant fi is the fundamental frequency of the QCM chip, and Zqis the acoustic impedance of the quartz crystal.

[0179] Animal Study. Approval for all experiments and procedures was granted by the Institutional Animal Care and Use Committee of Northwestern University under Protocol IS00024118. Adult male Sprague Dawley rats (Charles River) were acquired form Harlan Laboratories in Denver, CO. These rates were individually housed in Nalgene Plexiglas cages measuring 45 x 25.2 x 14.7 cm and were kept within the animal facility at Northwestern University. The facility maintained a 12: 12-hour light-dark cycle (with lights on at 07:00 and off at 19:00) and controlled temperature (22 °C) and humidity conditions.

[0180] Upon arrival, the rats were 9 to 10 weeks old and weighed approximately 230 grams. They were allowed to acclimate to the facility for one week before the commencement of the experimental procedures. Throughout the study, the rats had continuous access to food and water, following a diet of 2018 Teklad Global 18% protein (Harlan Laboratories, Denver, CO), except during fasting periods.

[0181] To establish a model of Type-1 Diabetes (T1D), streptozotocin (STZ) was administered at a dose of 65 mg / kg to healthy male Sprague Dawley rats from Charles River. These rats were monitored over a 7-day period, with blood glucose measurements taken every 48 hours using a glucometer. After this period, approximately 70% of the rats developed diabetes, as indicated by blood glucose levels exceeding 300 mg / dL.

[0182] On the day of the experiment, diabetic rats were placed under isoflurane anesthesia (1.5 to 2%), while a circulating heating pad maintained their body temperature around 37 °C.

[0183] For the fasting + insulin protocol, the diabetic rats, after their fasting period, received a subcutaneous injection of an overdose of insulin at t = 20 minutes (3 U / kg).

[0184] For the fasting+LPS injection group, diabetic rats received an intraperitoneal injection of LPS (1 pg / g) at t=220 min to induce an acute phase response and systemic inflammation. For all the animal scenarios blood glucose levels were monitored every 30 minutes via tail pricks using a glucometer during the animal study.

[0185] To prepare the rats for the experiment involving the microdevice (MN) application, the rats had their skin shaved, hair removal cream applied, and was thoroughly dried. The microdevice platforms were then administered to the rats’ dorsal skin by thumb-pressing, approximately 20 minutes before the experiment. The level of pro-inflammatory cytokines were measured electrochemically every 20 minutes, followed by a regeneration process after each measurement, which spanned roughly 320 minutes.

[0186] ELISA measurements. ISF extraction from diabetic rats. To extract ISF, the initial mass of the microneedle patch was recorded before it was inserted into the rat's back skin and secured with Tegaderm tape. After 10 minutes, the MN patches were removed from the rat's skin, transferred into an Eppendorf tube, and weighed. The difference in weight was recorded as the mass of the ISF.

[0187] To assess plasma IL-6 and TNF-a levels in diabetic rats using conventional ELISA, 250 pL of blood was collected from the tail vein of rats in the experimentalgroups every 40 minutes. These samples were promptly centrifuged at 2500 * g for 5 minutes. To prevent proteolysis, 1,000 KIU of an aprotinin / ethylenediaminetetraacetic acid solution per mL of whole blood was added to each tube before transferring blood for glucagon assay. The isolated plasma samples were stored at -20 °C until they were analyzed using ELISA kits to determine cytokine concentrations. The cytokine levels in the ISF were recovered from the MN patches (applied to the skin every 40 min) in an Eppendorf tube fdled with 100 pL of PBS by centrifuging at 10,000 rpm for 5 minutes. The levels of IL-6 and TNF-a were then measured using an ELISA kit.

[0188] Assessing systemic toxicity of microdevice using histology and immunohi stochemi stry , Rat organs (Heart, kidney, lung, spleen, liver, and skin) were isolated surgically after sacrifice and were fixed in 10% neutral buffered formalin followed by embedding in paraffin. Four-micrometer sections were stained with hematoxylin and eosin. The images were viewed on an Olympus microscope.

[0189] All the experimental protocols adhered to ethical and legal guidelines as specified by the Northwestern University Animal Care and Use Committee and were conducted with their approval.

[0190] Analysis of a panel of inflammatory genes using qPCR. To assess localized inflammation, skin tissue was surgically resected after the indicated experimental treatment. The tissue was then snap-frozen with liquid nitrogen and homogenized before RNA extraction. RNA was isolated according to manufacturer protocol (Qiagen: ID74104) and quality was assessed using Nanodrop. Next, the first strand cDNA was synthesized using SuperScript™ IV VILO™ Master Mix Catalog number: 11756500. Afterward, a panel of rat inflammatory genes was interrogated using TaqMan Rat Inflammation Arrays (Therm ofi scher 4414081) to assess differential expression. The qPCR run was performed using Quantstudio 3.

[0191] Statistical Analysis. All experimental data were derived from a minimum of three independent experiments, each including at least three parallel samples per condition. All data were retained after analysis. The diabetic animals were randomly selected for the test. Results are presented as means ± standard deviations (SD). Statistical significancewas evaluated using ANOVA and Student’s t-test, with a p-value of < 0.05 considered significant.

[0192] Molecular Dynamic Simulations and Simulation System Construction. The P2 DNA aptamer, with the sequence of AGCAGCACAGAGGTCAGATGGTCTGGAAACGACGAGGGCCACTGATTAACG TAGTTAATTGGTCTTGTCGAAGTGTAGTGTCTCCGTGGCTTTTCTTATTAGATA CAATAGCTGGTA was constructed based on four sequential steps, as previously reported. First, Mfold was employed to predict the DNA secondary structure based on free energy minimization techniques followed by constructing refined equivalent 3D RNA models using Assemble2 and Chimera. In the Mfold software, it approximates all conceivable secondary structures based on Watson-Crick base pairing, and subsequently, it singles out the most thermodynamically stable structure. To initiate this process, the initial sequences are set as linear at a temperature of 37 °C, with an ionic concentration of 1 M of Na+and 0 M of Mg2+to mimic the physiologically relevant ionic environment. Then, the 3D RNA model was converted into DNA model using VMD. Finally, VMD was used to minimize and optimize the final 3D DNA structure. For a base part of P2 DNA aptamer with the sequence of “TCTTATTAGATACAATAGCTGGTA,” the double helix structure was assigned, and its complementary strand was named Pl. Then, the Pl strand was connected to the surface of the gold slab (111 atoms) structure through a thiol moiety of CeSH. A crystalline slab of atomic gold with the (100) facet exposed, namely Au(100), was constructed as previously reported.

[0193] Molecular Docking and Molecular Dynamics (MD) Simulations. To simulate the impact of oscillation on the structural integrity of the aptamer interacting with human promyeloperoxidase as well as to investigate their binding affinity, the promyeloperoxidase structure was retrieved, identified by its PDB ID 5MFA, from the Protein Data Bank. The pyDockDNA server was used to dock the protein towards the constructed aptamer. Overall, two independent systems (in the absence or the presence of the oscillation) were introduced to the MD studies by Desmond software. Both systems, without or with oscillation, were solvated in explicit TIP3P (three-site transferrable Intermolecular Potential) water model and the OPLS3 (Optimized Potentials for LiquidSimulations version 3) force field parameters. A temperature of 310 K, pH of 7.4, and a pressure of 1 bar with a simulation length of 100 ns were assigned for each MD run. The Particle mesh Ewald method was employed to calculate the long-range electrostatic interactions. The cut-off radius for computing the Coulomb interactions was 9.0 A and a cubic periodic box with periodic boundary conditions was defined for both systems in the solvation step. For neutralizing each system, Na+and Cl’ counter ions were added. The distance of 10.0 A was assigned between the periodic boundary conditions and the closest aptamer / protein atom. For each system, the Martyna-Tuckerman-Klein chain coupling scheme and Nose-Hoover chain coupling scheme were engaged for the pressure and temperature control during MD simulation, respectively. A total of 1,000 frames / MD runs were allocated, and the trajectories were saved at 10 ps intervals for further analysis. BiasingForce or e bias script available in Desmond was employed to induce the oscillation. To understand the effects of oscillation on the structure and stability of aptamer-protein complex, the root mean squared deviation (RMSD), the root mean square fluctuation (RMSF), the radius of gyration (Rg) values, radial distribution function (RDF), and the number of different non-covalent interactions of molecules in two different systems were assessed. RMSD, which quantifies the average movement or displacement of an atom within a defined time interval, is determined using Eq 2:where the averaging is performed over the n pairs of equivalent atoms and di is the distance between the two atoms or particles in the ithpair. The Rg is defined as the distribution of atoms of a molecule around its axis and was computed from Eq 3:where N is the total number of residues, r(i) and reenter are the coordinates of a residue i and the center of mass, respectively.

[0194] RDF was employed to find a particle at distance “r” from another particle during MD simulation. The RDF analysis was performed on the MD trajectory obtained from the simulation of all systems by using the following formula Eq 4:

[0195] For clustering, the frames, trj cluster.py was employed to get the representative structure in aptamer, gold and protein containing systems in the absence and presence of oscillation. Affinity propagation is a clustering algorithm used in the Desmond software package to cluster molecular dynamics trajectory frames according to their similarity. Unlike other clustering algorithms, affinity propagation does not require the number of clusters to be specified in advance. Instead, the algorithm identifies the number of clusters by considering each frame as a potential cluster center and iteratively propagating the affinities (or similarities) between frames until a stable set of clusters is obtained. Affinity propagation has the advantage of being able to find clusters of different sizes and shapes, and it can also identify cluster centers and representative members for each cluster.

[0196] Fluid Dynamic Simulation. To simulate the flow of interstitial fluid (ISF) within the microdevice numerically, COMSOL Multiphysics was employed to solve the two- phase flow, capillary -based ISF movement into microdevice, and diffusion-based protein concentration equilibrium problems within an axisymmetric geometry. For the capillarybased ISF flow, initially, the primary channel of the needle contained air, while the nozzle was filled with ISF. Considering the contact angle of the printed TP-Q material is approximately 60 degrees, ISF adheres to the needle’s inner wall and gradually advances along the interior boundaries of the cylinder (the main channel). The deformation of the ISF surface results in the generation of surface tension at the interface between air and ISF, leading to a pressure discontinuity across this interface. These pressure variations drive the upward movement of both ISF and air.

[0197] In the simulation, the level set method was employed, which automatically establishes the equations governing the advection of the interface. The fluid interface isrepresented by the 0.5 contour of the level set function denoted as” Essentially, this function defines the volume fraction of ISF, with a value of 0 in the air and 1 within ISF. The transport of the fluid interface, which separates the two phases, is determined by the following equation Eq 4:in which e denotes the thickness of the interface, and y parameter determines the amount of reinitialization.

[0198] The Navier-Stokes equations are used to describe the movement of mass and momentum in fluids with constant density. To address capillary effects, the integrated surface tension was integrated into the model. As a result, the Navier-Stokes equations are adjusted as follows:Here, the variables are defined as follows: p represents the density (in kg / m3), p stands for dynamic viscosity (in Ns / m2), M signifies velocity (in m / s),p represents pressure (in Pa), and g denotes the gravitational acceleration vector (in m / s2). The Multiphysics coupling feature defines the density and viscosity as: p = pair+ P / SF ~ Pair)*and P = Pair "b (p[SF Pair) P-

[0199] The surface tension force, denoted by Fst, operates at the interface between air and ISF, and its computation is described by the formula Fst= aSKn. In this equation, n corresponds to the interface normal, computed as|. The symbol cr represents the surface tension coefficient (in N / m), K stands for curvature, computed as —V ■ n, and d is a Dirac delta function that is nonzero solely at the fluid interface. A set of partial differential equations under specific boundary conditions, encompassing hydrostatic pressure at the nozzle inlet, wall boundary conditions along the needle, atmospheric pressure at the outlet, and symmetry considerations for the needle’s central axis were solved numerically.

[0200] In the diffusion-based protein migration, the equation for conservation of proteins, - I- u. VC = DBV. (VC), in which DB is Brownian diffusion coefficient, is solved dt considering C = constant for inlet and dC / dn = 0 for needle walls and outlet. In this regard, the fluid velocity is zero.

[0201] To arrive at a solution, COMSOL employs a finite element method. In summary, the governing partial differential equation are transformed into their weak forms and subsequently solved numerically using the Galerkin finite element method. In this simulation, a nonuniform mesh is employed. Notably, mesh points were clustered in close proximity to the walls to enhance solution accuracy and capture gradients near these surfaces.

[0202] The iteration process concludes when the changes in dependent variables between two successive iterations satisfy the following established criterion: < 10-8.EQUIVALENTS

[0203] The present technology is not to be limited in terms of the particular embodiments described in this application, which are intended as single illustrations of individual aspects of the present technology. Many modifications and variations of this present technology can be made without departing from its spirit and scope, as will be apparent to those skilled in the art. Functionally equivalent methods and apparatuses within the scope of the present technology, in addition to those enumerated herein, will be apparent to those skilled in the art from the foregoing descriptions. Such modifications and variations are intended to fall within the scope of the present technology. It is to be understood that this present technology is not limited to particular methods, reagents, compounds, compositions or biological systems, which can, of course, vary. 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.

[0204] All patents, patent applications, provisional applications, and publications referred to or cited herein are incorporated by reference in their entirety, including all figures andtables, to the extent that are not inconsistent with the explicit teachings of this specification.

Claims

CLAIMSWhat is claimed is:

1. A method of actively resetting one or more biosensor probes in a biosensor system comprising rapidly oscillating the biosensor probes by inducing a change in electric current and / or applied electric field to the biosensor probes, wherein the biosensor probe comprises an analyte receptor that can bind to the analyte and wherein rapid oscillation of the biosensor probe results in the release of an analyte bound by the biosensor probe.

2. The method of claim 1, wherein the change in electric current or electric field is a change of up to 0.36 V.

3. The method of claim 1 or 2, wherein the frequency of rapid oscillation is between 50 and 95 Hz.

4. The method of any one of claims 1-3, wherein the biosensor system comprises:(a) an electrode; and(b) a plurality of biosensor probes, each biosensor probe comprising a negatively- charged rigid rod structure with a first terminus and a second terminus and the analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure.

5. The method of claim 4, wherein the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA).

6. The method of claim 4 or 5, wherein the electric current and / or applied electric field is applied to the biosensor probes via the electrode.

7. The method of any one of claims 4-6, wherein the surface of the electrode is further bound by a chemical monolayer.

8. The method of claim 7, wherein the analyte is selected from a small molecule, a protein, or other biological entity.

9. A method of detecting an analyte using a biosensor system, comprising allowing a biosensor probe to contact a biological sample that may contain the analyte, wherein the biosensor probe is configured to rapidly oscillate in response to a change in electriccurrent and / or applied electric field, and after a first detecting step is performed, applying a change in electric current or electric field to the biosensor probe to cause rapid oscillation before performing subsequent detection steps with the biosensor probe.

10. The method of claim 9, wherein the change in electric current or electric field is a change of up to 0.36 V.

11. The method of claim 9 or 10, wherein the frequency of rapid oscillation is between 50 and 95 Hz.

12. The method of any one of claims 9-11, wherein the biosensor system comprises:(a) an electrode; and(b) a plurality of biosensor probes, each biosensor probe comprising a negatively- charged rigid rod structure with a first terminus and a second terminus and an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure.

13. The method of claim 12, wherein the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA).

14. The method of claim 12 or 13, wherein the electric current and / or applied electric field is applied to the biosensor probes via the electrode.

15. The method of any one of claims 12-14, wherein the surface of the electrode is further bound by a chemical monolayer.

16. The method of claim 15, wherein the analyte is selected from a small molecule, a protein, or other biological entity.

17. The method of any one of claims 9-16, wherein the human biological sample comprises an interstitial fluid sample, a sweat sample, saliva sample, or serum sample.

18. A method of determining a concentration of an analyte, the method comprising:(a) applying a biological sample to a biosensor system, wherein the biosensor system comprises:(i) an electrode, and(ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a secondterminus and an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure, wherein each biosensor probe is attached to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure, wherein the analyte present in the biological sample binds to the analyte receptors of the plurality of the biosensor probes;(b) applying an electric current to the biological sample and the plurality of biosensor probes via the electrode, wherein the biosensor probes and analyte bound by the analyte receptors attached to the biosensor probes are displaced toward the surface of the electrode;(c) measuring the chronoamperometric response of the electric current by translating the difference between an unbound electron transfer rate and a bound electron transfer rate;(d) calculating the concentration of the analyte in the biological sample based on a chronoamperometric response based on difference between the electrochemical signal produced by biosensor probes have bound the analyte and by biosensor probes have not bound the analyte; and(e) actively resetting the biosensor probes by inducing rapid oscillation of the biosensor probes via the electrode using changes in electric current and / or applied electric field.

19. The method of claim 18, further comprising repeating steps (a) through (f) at least once.

20. The method of claim 18 or 19, wherein the change in electric current or electric field applied to the biosensor probes to induce rapid oscillation is a change up to 0.36 V.

21. The method of claim any one of claims 18-20, wherein the frequency of rapid oscillation is between 50 and 95 Hz.

22. The method of any one of claims 18-21, wherein the negatively-charged rigid rod structure comprises double-stranded DNA (dsDNA).

23. The method of claim 22, wherein upon application of the electric current, the electrode induces an electron transfer reaction with the redox reporter as the biosensor probes are displaced toward the surface of the electrode.

24. The method of claim 23, wherein the chronoamperometric response is dependent on a time rate at which the biosensor probes are displaced.

25. The method of claim 24, wherein the unbound electron transfer rate is dependent on a time rate at which biosensor probes that have not bound an analyte at the analyte receptor are displaced.

26. The method of claim 24 or 25, wherein the bound electron transfer rate is dependent on a time rate at which biosensor probes that have bound an analyte at the analyte receptor are displaced.

27. The method of any one of claims 23-26, wherein, upon application of the electric current, the redox reporter approaches the electrode and the electron transfer is based on a redox reaction or electron tunneling current.

28. The method of claim 27, wherein the analyte is selected from a small molecule, a protein, or other biological entity.

29. The method of any one of claims 18-27, wherein the biosensor systems further comprise a chemical monolayer bound to the surface of the electrode.

30. The device of claim 29, wherein the chemical monolayer comprises a self-assembling chemical monolayer.

31. The device of claim 30, wherein the self-assembling chemical monolayer comprises 6- mercaptohexanol (MCH).

32. The method of any one of claims 29-31, wherein, upon application of the electric current, the biosensor probes, redox reporters attached to the biosensor probes, and the analyte bound by the analyte receptors attached to the biosensor probes are displaced through the chemical monolayer and toward the surface of the electrode.

33. The method of claim 32, wherein the redox reporter attached to the biosensor probe and displaced toward the surface of the electrode through the chemical monolayer is oxidized by the electrode.

34. The method of claim 33, wherein the chronoamperometric response is dependent upon the change in electrochemical signal caused by oxidation of the redox reporter displaced toward the surface of the electrode through the chemical monolayer.

35. The method of any one of claims 18-34, wherein the biological sample is from a human subject.

36. The method of claim 35, wherein the human biological sample comprises an interstitial fluid sample, a sweat sample, saliva sample, or serum sample.

37. The method of any one of claims 18-36, wherein the method is used to continuously monitor the concentration of the analyte in a subject.

38. A device for continuously detecting an analyte in a biological sample from a subject, the device comprising:(a) a collector component for collecting the biological sample, wherein the collector component comprises a microneedle having three inlet channels that connect to a main channel in the interior of the microneedle;(b) a detection component attached to the collector component comprising:(i) an electrode, and(ii) a plurality of biosensor probes, each biosensor probe comprising a negatively-charged rigid rod structure with a first terminus and a second terminus, wherein each biosensor probe comprises an analyte receptor and a redox reporter attached to the first terminus of the negatively-charged rigid rod structure and wherein each biosensor probe is bound to the surface of the electrode at the second terminus of the negatively-charged rigid rod structure;(c) an electrochemical measurement component connected to the electrode of the detection component, wherein the electrochemical measurement component detects changes in the electric current at the surface of the electrode.

39. The device of claim 38, wherein the main channel of the collector component introduces the biological sample to the electrode and biosensor probes bound to the surface of the electrode via a side channel connecting the collector component and the detection component.

40. The device of claim 38 or 39, wherein the microneedle of the collector component induces capillary force upon the biological sample from the subject upon insertion of the microneedle into the skin of the subject.

41. The device of any one of claims 38-40, wherein the size of the inlet channels of the microneedle are 300 pm x 400 pm.

42. The device of any one of claims 38-41, wherein the diameter of the main channel of the microneedle is 300 pm.

43. The device of any one of claims 38-42, wherein the electrochemical measurement component comprises a display that provides the results of the analyte detection.

44. The device of any one of claims 38-43, wherein the electrochemical measurement component is connected to a separate component that displays the results of the analyte detection.

45. The device of any one of claims 38-44, wherein the negatively-charged rigid rod structure comprises dsDNA.

46. The device of claim any one of claims 38-45, wherein the analyte receptor comprises an antibody, an antibody fragment, a nanobody, or an aptamer.

47. The device of any one of claims 38-46, wherein the analyte is selected from a small molecule, a protein, or other biological entity.

48. The device of any one of claims 38-47, wherein the detection component further comprises a chemical monolayer bound to the surface of the electrode.

49. The device of claim 48, wherein the chemical monolayer comprises a self-assembling chemical monolayer.

50. The device of claim 49, wherein the self-assembling chemical monolayer comprises 6- mercaptohexanol (MCH).

51. The device of any one of claims 38-50, wherein the subject is a human.

52. The device of any one of claims 38-51, wherein the biological sample is an interstitial fluid sample.