Rapid multi-analyte sensor devices, systems, and methods

Abstract

Disclosed are rapid multi-analyte sensor devices, systems and methods. In some aspects, a sensor device for rapid detection of two or more analytes include a substrate; a first analyte sensor comprising a first working electrode disposed on the substrate and a first functionalization layer disposed on the first working electrode, wherein the first functionalization layer includes a catalyst to facilitate an enzymatic electrochemical reaction to detect a first analyte at the first analyte sensor; and a second analyte sensor comprising a second working electrode disposed on the substrate and a second functionalization layer disposed on the second working electrode, wherein the second functionalization layer includes an aptamer receptor to facilitate an aptamer affinity bioassay reaction to detect a second analyte at the second analyte sensor.

Description

PCT Application Attorney Docket No.: 009062.8573.WG00RAPID MULTI-ANALYTE SENSOR DEVICES, SYSTEMS, AND METHODS CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This patent document claims priority to and benefits of U.S. Provisional Patent Application No. 63 / 747,852 entitled “RAPID MULTI-ANALYTE SENSOR DEVICES, SYSTEMS, AND METHODS” filed on January 21, 2025. The entire content of the aforementioned patent application is incorporated by reference as part of the disclosure of this patent document.TECHNICAL FIELD

[0002] This patent document relates to devices, systems, and processes that use biosensing for detection of multiple analytes.BACKGROUND

[0003] Biosensors can provide real-time detection of physiological substances and processes in living things. A biosensor is an analytical tool that can detect a chemical, substance, or organism using a biologically sensitive component coupled with a transducing element to convert a detection event into a signal for processing and / or display. Biosensors can use biological materials as the biologically sensitive component, e.g., such as biomolecules including enzymes, antibodies, nucleic acids, etc., as well as living cells. For example, molecular biosensors can be configured to use specific chemical properties or molecular recognition mechanisms to identify target agents, which can be useful in diagnosis and treatments for various health care applications.SUMMARY

[0004] Disclosed are rapid multi-analyte sensor devices, systems and methods. In some aspects, the disclosed technology provides a reagent-less, multiplexed bio-sensor chip for rapid, precise, and real-time / continuous multi-analyte monitoring of analytes such as protein, peptide, hormone, and / or metabolite markers, including insulin and glucose. The detection can be made from a small drop of a biofluid, such as blood or saliva, and the sensor device can be configured in a deployable and wearable device format. In some aspects, the disclosed multi-analyte sensorPCT Application Attorney Docket No.: 009062.8573.WG00devices, systems, and methods integrate different assays (e.g., enzymatic and aptamer assays) on a single chip for simultaneous detection of metabolites and protein / peptides.

[0005] In some implementations, for example, the disclosed sensor devices, systems, and techniques provide a reagent-less multiplexed biosensor chip to carry out a coupled biocatalytic and aptamer bioassay for rapid and precise on-site monitoring of protein / peptide hormone and metabolite biomarkers, e.g., particularly for simultaneous and ultra-rapid insulin and glucose sensing. The disclosed devices, systems, and techniques integrate aptamer-based insulin sensing and enzymatic glucose detection onto a single microchip platform, which in some embodiments includes a four-electrode sensor chip. In such embodiments, for example, the insulin sensor contingent leverages a redox indicator-modified aptamer for voltammetric detection, while the glucose sensor contingent uses an enzymatic amperometric approach. The disclosed devices, systems, and techniques enable simultaneous, reagent-less detection of millimolar glucose and picomolar insulin concentrations from single microliter droplets of serum or saliva within two minutes (e.g., one minute or less detection of glucose from the microsample and two minutes or less detection of insulin from the same microsample), and can do so without any apparent crosstalk. Example embodiments of the multiplexed biosensor chip can achieve rapid, cost-effective, and reliable detection of various protein / peptide / small molecule analytes, which are not limited to diabetes and include protein / peptide / small molecule analytes for measuring inflammation, cancerous, and Parkinson disease biomarkers.

[0006] In implementations, for example, the disclosed multi-analyte sensor devices and techniques provide a cutting-edge tool for quick and precise health monitoring that fits on a tiny chip. In some embodiments of the multi-analyte sensor in accordance with the present technology, two advanced sensing techniques are combined and integrated with the chip design to detect insulin and glucose levels from just a small drop of blood or saliva in under two minutes — without extra chemicals or complicated steps needed for the simultaneous, rapid detection. In such implementations, the sensor chip uses specialized molecular tools to measure insulin and glucose levels accurately without interference. For example, insulin is detected using a DNA aptamer molecule that changes its electrical signal when insulin is present, while glucose is measured through a chemical reaction that generates an electrical signal. Beyond diabetes, the disclosed technology can also be adapted to detect signs of other conditions like sepsis, inflammation, cancer, or Parkinson’s disease. The disclosed technology provides a fast,PCT Application Attorney Docket No.: 009062.8573.WG00affordable, and reliable way to monitor health using only a tiny sample in remote areas and home self-testing applications.

[0007] In some aspects, a sensor device for rapid detection of two or more analytes includes a substrate; a first analyte sensor comprising a first working electrode disposed on the substrate and a first functionalization layer disposed on the first working electrode, wherein the first functionalization layer includes an enzyme configured to facilitate a first electrochemical reaction enabling detection of a first analyte at the first analyte sensor; a second analyte sensor comprising a second working electrode disposed on the substrate and a second functionalization layer disposed on the second working electrode, wherein the second functionalization layer includes a redox indicator-modified aptamer receptor configured to facilitate a second electrochemical reaction enabling detection of a second analyte at the second analyte sensor: a counter electrode configured to complete an electrical circuit with respect to one or both of the first working electrode and the second working electrode during an electrochemical detection technique; and a reference electrode configured to provide a fixed electrical potential for reference to measure an electrical potential at one or both of the first working electrode and the second working electrode during the electrochemical detection technique.

[0008] In some aspects, a method for rapid monitoring of two or more analytes in a biofluid sample includes providing a reagent-less sensor device that detects a first analyte by an enzymatic electrochemical reaction measured by a first electrochemical detection technique and detects a second analyte by an aptamer-based electrochemical reaction measured by a second electrochemical detection technique.

[0009] The subject matter described in this patent document can be implemented in specific ways that provide one or more of the following features.BRIEF DESCRIPTION OF THE DRAWINGS

[0010] FIG. 1A shows a diagram illustrating an example embodiment of a reagent-less dual glucose-insulin biosensor chip in accordance with example embodiments of the disclosed multianalyte sensor devices in accordance with the present technology.

[0011] FIG. IB shows a schematic of the enzymatic and redox cycling process involved in an example glucose (G) sensor (WEI) in accordance with the present technology.

[0012] FIG. 1C shows a schematic of insulin (I) recognition involving a guanine-richPCT Application Attorney Docket No.: 009062.8573.WG00aptamer in an example I sensor (WE2) and insulin-binding induced changes in the signaling aptamer probe in accordance with the present technology.

[0013] FIG. 2A shows a schematic of an example embodiment of a scalable, low-cost dual G-I chip fabrication process in accordance with the present technology.

[0014] FIGS. 2B and 2C show schematics of an example embodiment of a dual G-I senor fabrication protocol in accordance with the present technology.

[0015] FIG. 3A and 3B show an example timeline and operation of reagent-less simultaneous G-I measurements in accordance with the present technology.

[0016] FIG. 4A shows a schematic of an example single sensor platform along with signal transduction reaction for the sensing of glucose.

[0017] FIG. 4B shows a data plot depicting cyclic voltammetry curves of an example G sensor (at 100 mV / s), performed in the absence and in the presence of 1 mM glucose.

[0018] FIGS. 4C and 4D show data plots depicting the chronoamperometric response of an example G sensor at +0.3 V, examining the effect of the enzymatic reaction time, from 0 (i) to 90 s (xii) in 10 s increment, for detecting 10 mM G in PBS.

[0019] FIGS. 4E and 4F show data plots depicting the chronoamperometric response of an example G sensor to increasing G concentrations, from 0 (i) to 15 mM (vii) in 2.5 mM increments.

[0020] FIG. 5A shows a schematic of an example single sensor platform along with signal transduction reaction for the sensing of insulin.

[0021] FIG. 5B shows a data plot depicting a cyclic voltammogram obtained in 0.05 M, H2SO4 solution (performed between 0 and 1.8 V at 100 mV / s). to increase the microscopic surface area of the electrode and determine the electroactive surface area of the gold electrode.

[0022] FIG. 5C shows a data plot depicting a cyclic voltammogram performed between 0 and -0.5 V to quantify the number of aptamers modified with the redox reporter on the electrode, integrating the area under the peak associated with methylene blue reduction.

[0023] FIG. 5D shows data plots illustrating the exemplary results of the study of the square wave voltammetry (SWV) frequency for the E-AB sensor for insulin, (a) Profile of current changes as a function of SWV frequency for the E-AB sensor for insulin, before (black line) and after addition of 100 nM insulin (red line), (b) SWV current responses of sensor I at different frequencies: black curve (no insulin) and red curve (100 nM insulin) at: (i) 5 Hz; (ii) 20 Hz; (iii)PCT Application Attorney Docket No.: 009062.8573.WG0050 Hz; (iv) 100 Hz; (v) 200 Hz; (vi) 300 Hz; (vii) 400 Hz; (viii) 500 Hz.

[0024] FIG. 5E shows data plots depicting SWV voltammetric current responses of an example I sensor towards 100 nM I at every 30 s from 0 to 180 s.

[0025] FIG. 5F shows data plots depicting SWV voltametric current change profiles (black line: blank, green line: 10 nM insulin, red line: 100 nM insulin).

[0026] FIGS. 5G and 5H show data plots depicting the voltammetric current response of an example E-AB sensor to increasing I concentrations from 0 to 50 nM in PBS with 2 min incubation.

[0027] FIG. 6A shows a schematic depiction of selectivity and influence of electroactive interfering species on an example G-I chip.

[0028] FIGS. 6B and 6C show data plots depicting the amperometric response of ascorbic acid (100 pM AA) and uric acid (100 pM UA) in 1 mM G for an example dual G-I senor.

[0029] FIG. 6D shows data plots depicting the SWV current responses with 100 pM of AA, UA, and H2O2 and 100 nM of glucagon, BSA in 50 nM I for an example dual G-I sensor.

[0030] FIG. 7A shows data plots depicting amperometric and voltammetric responses of example G-I sensor chips for increasing G concentrations and in the presence of a fixed I level: (Chip A) 2.5 mM G / 10 nM I, (Chip B) 5 mM G / 10 nM I, and (Chip C) 10 mM G / 10 nM I.

[0031] FIG. 7B shows data plots depicting amperometric and voltammetric responses of example G-I sensor chips for increasing I concentrations in the presence of a fixed G level: (Chip A) 10 mM G / 5 nM I, (Chip B) 10 mM G / 100 nM I, and (Chip C) 10 mM G / 300 nM I.

[0032] FIGS. 8 A and 8B show data plots depicting the amperometric response of an example G-I chip for increasing concentration of glucose from 0 (i) to 15 mM (vi) in 3 mM increments in undiluted human serum.

[0033] FIG. 8C shows a data plot depicting the voltammetric response of an example G-I chip for increasing insulin concentration from 0 to 125 nM (using 25 nM increments) in undiluted human serum.

[0034] FIGS. 9 A and 9B show data plots depicting the amperometric response of an example G-I chip for increasing glucose concentration from 0 (i) to 7 mM (vii) in 1 mM increments in undiluted human saliva.

[0035] FIG. 9C shows a data plot depicting the voltammetric response of an example G-I chip to increasing insulin concentration from 0 to 125 nM (using 25 nM increments) in undilutedPCT Application Attorney Docket No.: 009062.8573.WG00human saliva.

[0036] FIG. 10 shows data plots depicting the performance of example G-I chips in G and I spiked human saliva samples of three different subjects and the current changes calculated from the amperometric and voltammetric curves.DETAILED DESCRIPTION

[0037] Disclosed are rapid multi-analyte sensor devices, systems and methods. In some aspects, the disclosed technology provides a reagent-less, multiplexed bio-sensor chip for rapid, precise, and real-time / continuous multi-analyte monitoring of analytes such as protein, peptide, hormone, and / or metabolite markers, including insulin and glucose. The detection can be made from a biofluid such as a small drop of blood or saliva, and the sensor device can be configured in a deployable and wearable device format. In some aspects, the disclosed multi-analyte sensor devices, systems, and methods integrate different assays (e.g., enzymatic and aptamer assays) on a single chip for simultaneous detection of metabolites and protein / peptides.

[0038] Unlike the conventional analyte sensor devices, the disclosed sensor devices and techniques provide a combination of enzymatic and aptamer-based dual analyte rapid sensing of glucose and insulin on a small microchip footprint. Previous multi-analyte sensor devices have described using coupling enzymatic and affinity assay such as immunoassay for glucose and insulin sensing, but it requires long incubation time for detection about >20 minutes including multiple washing steps. Example embodiments of the disclosed rapid detection of dual or multianalyte diabetes biomarker technique can facilitate dramatically faster results with less complex sensor preparation with respect to existing centralized and point-of-care bioassays.

[0039] Example embodiments of the disclosed multi-analyte sensor device and technique provide the first demonstration of reagent-less protein analyte detection along with metabolite in a single form factor without requiring any additional treatment or reagent for electrochemical transduction. For example, the reagent-less, wash-free and rapid microchip detection of electrochemical aptamer-based insulin (about 2 minutes) was demonstrated in human serum and saliva. Moreover, the potential interference of protein and other electrochemically active species, e.g., such as ascorbic acid, uric acid, acetaminophen with the disclosed multi-analyte sensor and technique was studied extensively in experimental implementations, which demonstrated low or no influence in the electrochemical analysis. Notably, the disclosed multi-PCT Application Attorney Docket No.: 009062.8573.WG00analyte sensors, systems, and techniques can extend to a large variety of multi-analytes, e.g., such as aptamer-based assays for two peptide hormones (e.g., insulin and glucagon) and two metabolites (e.g., glucose and ketone) for advanced diabetes monitoring.

[0040] Some existing multi-analyte sensor devices use coupling of enzymatic and affinity assay, such as immunoassay for glucose and insulin. However, the immunoassay for insulin requires two antibodies (e.g.. capture and signal-tagged detection antibodies) and a reagent (e.g., 3,3',5,5'-Tetramethylbenzidine (TMB)) for electrochemical transduction, which makes this bioassay less simple, miniaturize, and user friendly.

[0041] The disclosed technology provides a reagent-less system that couples biocatalytic (e.g., enzymatic) and aptamer-based affinity bioassays on a single chip. This innovative platform can simultaneously measure glucose and insulin in biofluids without requiring additional washing steps or external reagents within 2 minutes using a single microliter sample droplet. For example, the measurements of two or more analytes using the disclosed platform can obtain measurements in a minimal amount of time, e.g., less than 150 seconds, thereby being virtually simultaneous. In some example, the sensor device includes a metal-sputtered electrochemical chip, a glucose oxidase- and mediator-modified enzymatic sensor for glucose detection, and a redox indicator (methylene blue)-modified, guanine-rich (G-rich) thiolated aptamer receptor for insulin detection.

[0042] The disclosed sensor devices, systems, and techniques can address critical challenges related to minimizing crosstalk and cross-reactivity between adjacent biocatalytic and aptamer affinity sensors on the same chip, despite the integration of distinct surface chemistries for the enzymatic and aptamer sensors.

[0043] The disclosed sensor devices, systems, and techniques can employ distinct detection principles and electrochemical transduction mechanisms. For example, (i) an aptamer sensor works under a direct electron transfer between a sensor surface and the redox indicator on the aptamer, detecting target binding-induced conformational changes; and (ii) an enzymatic sensor works under mediated electron transfer between a sensor and enzyme during biochemical reactions.

[0044] The disclosed sensor devices, systems, and techniques can work largely in physiological concentrations of picomolar / nanomolar concentrations of protein, peptide analytes, and millimolar concentrations of metabolites.PCT Application Attorney Docket No.: 009062.8573.WG00

[0045] The disclosed sensor devices, systems, and techniques can support regeneration for subsequent sensing cycles due to the reversible binding of aptamers to their target analytes and the direct biochemical reactions of enzymes with their substrates.

[0046] The disclosed sensor devices, systems, and techniques can be further extended to detect more than two analytes, such as two electrochemical aptamer-based sensors for two peptide hormones (insulin and glucagon) and two enzymatic sensors for two metabolites (glucose and ketone), or larger numbers of target biomarkers.

[0047] Beyond diabetes monitoring, the disclosed sensor devices, systems, and techniques can detect various protein, peptide, and small-molecule biomarkers, including those related to inflammation, cancer, and neurodegenerative diseases such as Parkinson’s disease.

[0048] The disclosed sensor devices, systems, and techniques are versatile and applicable for use in hospitals, clinics, home self-testing, and remote diagnostic settings, allowing biomarker detection across different biofluids.

[0049] Example embodiments of the sensor of the disclosed sensor devices, systems, and techniques can include the following. For example, the sensor electrodes can be fabricated using thin-film metal deposition (e.g., sputtering) for a compact, low-cost chip capable of performing simultaneous bioassays with minimal (e.g., microliter) sample volumes. The sensor can be for: metabolites (e.g., glucose, ketone, and / or lactate), small molecules and ions (e.g., cortisol, uric acid, vitamin C, acetaminophen, levodopa, sodium, potassium, chloride, calcium, ammonium, alcohol, and / or caffeine), and proteins and peptides (e.g., insulin, glucagon, cytokines, and / or cancerous biomarkers). In some embodiments, for example, the sensor can include at least four electrodes comprising two working electrodes, one reference electrode, and one counter electrode. The electrodes may include the following components: aptamers, enzymes, mediators, polymers, surfactants, crosslinkers, thiol monolayers, stabilizers, and / or peptides. In some examples, each electrode may include conductive current collectors made of gold, silver, copper, or platinum. In example embodiments of the sensor using an aptamer, the aptamer may include deoxyribonucleic acid (DNA), ribonucleic acid (RNA), and / or peptide nucleic acid (PNA) oligonucleotides. In example embodiments of the sensor using an enzyme, the enzyme may include glucose oxidase, P-hydroxybutyrate dehydrogenase, lactate oxidase, uricase, tyrosinase, alcohol oxidase, glucose dehydrogenase, alcohol dehydrogenase, and / or lactate dehydrogenase. In example embodiments of the sensor using a mediator, the mediator may include ferrocene,PCT Application Attorney Docket No.: 009062.8573.WO00quinones, ferricyanide, tetrathiafulvalene, polyoxometalates, nicotinamide adenine dinucleotide (NADH), and / or Prussian blue. The sensor output can be measured using electrochemical techniques including but not limited to voltammetry and amperometry and processed with appropriate algorithms to get an estimate of the blood glucose concentration trend. The sensor device may be wearable format either on the skin (e.g., a fluidic system relying on a moisture-wicking material such as paper or nitrocellulose membrane) or under the skin (e.g., microneedles) for detecting a wide range of analytes by coupling biocatalytic and bioaffinity assays in one form factor. Example implementations of the disclosed sensor devices, systems, and techniques can measure the target analytes in a biofluid. For example, the biofluid can be blood, serum, saliva, sweat, urine, or interstitial fluid.Example Embodiments and Implementations Demonstrating Simultaneous and Rapid Detection of Glucose and Insulin: Single Chip Enzymatic and Aptamer-based Assays

[0050] Diabetes management demands precise monitoring of key biomarkers, particularly insulin (I) and glucose (G). In example implementations discussed below, an exemplary bioelectronic chip device in accordance with the disclosed technology is presented that enables the rapid, simultaneous detection of I and G in biofluids, e.g., within 2 minutes. This example dual biosensor chip integrates aptamer-based insulin sensing with enzymatic glucose detection on a single platform, employing a four-electrode sensor chip. For the example embodiment disclosed below, the insulin voltammetric sensor employs a G-quadraplex methylene-blue-modified aptamer, while the amperometric biocatalytic glucose sensor utilizes a second-generation mediator-based approach. Simultaneous reagent-less sensing of I and G has been achieved by addressing key challenges. These include combining different surface chemistries, assay formats and detection principles at closely-spaced working electrodes and the substantially different concentration levels of the I and G targets. An attractive analytical performance, with no apparent crosstalk, is demonstrated for the simultaneous detection of millimolar G concentrations and picomolar I concentrations in single microliter serum or saliva sample droplets. This exemplary dual biosensor offers rapid, cost-effective, and reliable monitoring, addressing the unmet need for integrated multiplexed diabetes biomarker detection in decentralized settings. Such integration of enzymatic and aptamer-based bioassays could greatly expand the scope of decentralized testing in healthcare, beyond diabetes care.

[0051] Managing diabetes is a critical global health challenge affecting millions worldwide.PCT Application Attorney Docket No.: 009062.8573.WO00Significant efforts have been made over the past three decades to advance diabetes care towards tighter glycemic control. Glucose (G) has been the most momentous biomarker in the management of type 1 (T1D) and type 2 (T2D) diabetes. Besides glucose. Insulin (I) is essential for regulating glucose metabolism and maintaining normal blood glucose levels. Early detection of insulin resistance, which can occur up to a decade before clinically diagnosed T2D, underscores the need for combined G and I analyses in prediabetic individuals.

[0052] Current automated insulin delivery (AID) systems and artificial pancreas (AP) platforms are commonly limited to a single (G) biomarker input and fail to consider changes in other diabetes biomarkers influenced by diet, exercise, and other activities. Optimal I dosing in AID and AP systems should rely on real-time data from multiple biomarkers to enhance personalized treatment. Point-of-care (POC) diagnostics can provide significant insight into an individual’s overall homeostatic condition in decentralized settings. Accordingly, there are urgent needs for multiplexed POC detection of key diabetes biomarkers, for facilitating fast, accurate, and personalized treatment without long delays associated with centralized lab analyses. In particular, a simple and rapid simultaneous on-site detection of glucose and insulin via a single sample droplet is paramount for effective diabetic care.

[0053] Insulin is typically measured in serum or plasma samples in centralized laboratories using gold-standard affinity assays, such as enzyme-linked immunosorbent assay (ELISA) methods, which significantly increase the sample-to-answer time and overall assay costs. Such traditional insulin assays thus fall short of supporting immediate corrective action, particularly timely interventions to tackle unforeseen events. To address the growing demand for simpler and faster on-site I measurements, substantial efforts are being devoted to developing decentralized I immunoassays, replacing conventional ELISA methods. Considering the importance of I monitoring alongside G, recent efforts have led to the development of a dualanalyte enzymatic / immunoassay sensing chip for detecting G and I. However, fabricating these electrochemical immunosensors relies on lengthy (e.g., -20-30 min) incubation times along with binding of an enzyme-tagged detection antibody and the corresponding washing steps and detection reactions that greatly increase complexity of such POC assays. Additionally, such assays are constrained by the limited availability and high costs of high-quality insulin antibodies. Moreover, the requirement of a long incubation (e.g., -20 min) in I immunoassays hinders their use for rapid on-site frequent insulin measurements.PCT Application Attorney Docket No.: 009062.8573.WO00

[0054] The ability to track closely temporal I concentration profiles would provide effective guidance for tailoring I dosing during diabetes management. Economically, reagent-free sensors are beneficial as they reduce long-term operating costs by eliminating the need for reagents and consumables. Various synthetic I receptors have been proposed recently for achieving highly selective and reagent-less I measurements, including insulin- specific aptamers biomimetic molecularly imprinted polymers (MIP), and oranti-insulin single chain variable fragment (scFv). Among these, aptamers have gained growing popularity due to their ability to bind reversibly a broad range of target analytes towards reagent-less biosensing. Electrochemical aptamer-based (E-AB) sensors have been introduced recently towards real-time wearable sensing applications of diverse biomarkers.

[0055] In this example study, rapid, simultaneous, and reagent-less detection of I and G is demonstrated within 2 minutes on a single sensing chip device, e.g., integrating rapid aptamerbased I detection alongside parallel biocatalytic G measurements (FIG. 1A-1C). This disposable bioelectronic device thus offers rapid and simple on-site I and G monitoring by integrating E-AB-based I detection and enzymatic G detection on a single platform. The E-AB biosensing approach can be easily transferred to such rapid decentralized in-vitro measurements, e.g., home self-testing.

[0056] FIG. 1A shows a diagram illustrating an example embodiment of a reagent-less dual glucose-insulin biosensor chip in accordance with example embodiments of the disclosed multianalyte sensor devices in accordance with the present technology. The diagram shows schematics of a sensor chip 100 on a substrate 102 (e.g., a polyethylene terephthalate glycol substrate) with two gold (Au) working electrodes (WEs) 104 and 106, for the glucose (G) and insulin (I) sensors, respectively. The sensor chip 100 also includes an Ag / AgCl reference electrode (RE) 108 and an Au counter electrode (CE) 110. The G sensor also includes a layer of immobilized glucose oxidase enzyme (GOx) 112, while the I sensor includes a layer of insulin aptamer bioreceptors 114, which provide localized detection of G and I on the single sensor chip 100.

[0057] FIG. IB shows illustrations of enzymatic and redox cycling process involved in the G sensor. For example, G detection is performed amperometrically at +0.3 V on the Au WEI 104 through a tetrathiafulvalene (TTF)-mediated enzymatic (GOx) oxidation of glucose in the GOx layer 112.PCT Application Attorney Docket No.: 009062.8573.WO00

[0058] FIG. 1C shows illustrations of an insulin recognition involving a guanine-rich aptamer layer 114 on the I sensor WE2 106 and insulin-binding induced changes in the signaling aptamer probe, resulting in a decrease in measured voltammetric peak currents.

[0059] FIG. 2A shows schematic of an example embodiment of a scalable, low-cost dual G-I chip fabrication process in accordance with the present technology. At step 202, a polyimide layer is laminated on a polyethylene terephthalate glycol (PETG) substrate 102. At step 204, a pattern is cut on the laminated film / PETG. At step 206, the pattern is pilled to selectively expose portions of the PETG substrate for sputtering. At step 208, Cr / Au / Ag sputtering is performed to produce WEI 104, WE2 106, RE 108, and CE 110.

[0060] FIG. 2B shows a simplified illustration of fabrication of the G-I sensing chip array on a substrate, e.g., a polyethylene terephthalate glycol (PETG) substrate, following sputtering step 208. At step 210, etching is performed on the WEs 104 and 106 and the CE 110. At step 212, the RE 108 is oxidized

[0061] FIG. 2C shows a schematic of further details of an example embodiment of a dual G-I senor fabrication protocol in accordance with the present technology. At step 209, a substrate with a plurality of sputtered electrodes is diced into individual G-I sensing chips. At step 210A, the Ag electrode layer is etched by nitric acid (HNOs). At step 212A, the RE 108 is chloridated to Ag / AgCl. At step 214, TTF is drop cast onto the WEI 104. At step 216, GOx is drop cast onto the TTF in three layers. At step 218. chitosan and nafion are loaded onto the GOx layer. At step 220, insulin aptamers are immobilized onto the WE2 106. At step 222, 6-mercapto-l -hexanol (MCH) is immobilized with the insulin aptamers, to yield the dual G-I sensor chip 100.

[0062] FIG. 3A shows illustrations of a timeline for reagent-less simultaneous G-I measurements: glucose is measured amperometrically in 90 s, including a 60 s incubation period 302 in which an enzymatic reaction takes place, followed by a 30 s G detection period 304, followed by an additional 30 s of incubation 306 for insulin-aptamer binding to occur, and 15 s for square wave voltammetry (SWV) insulin measurement 308. FIG. 3A also depicts additional details of the G and I sensors and corresponding reactions. For example, the G sensor may include a WEI 104, TTF and GOx layer 112, and a capping layer 312. The I sensor may include a WE2 106 and a layer 114 of insulin aptamers modified with methylene blue.

[0063] FIG. 3B shows an illustration of the exemplary glucose-insulin biosensor chip utilized in an example implementation for rapidly detecting glucose and insulin levels (e.g., < 1 -minutePCT Application Attorney Docket No.: 009062.8573.WO00detection of glucose and < 2-minute detection of insulin) from a microliter sample (e.g., serum or saliva).

[0064] As illustrated in FIGS. 1A-1C and 3A-3B, the rational chip design and its judicious optimization have enabled reliable simultaneous measurements of G and I in a single sample droplet (e.g.. serum or saliva) without crosstalk between the neighboring enzymatic (G) and aptamer (I) sensors, despite the different surface chemistries and detection principles at these adjacent electrodes, and the large difference in the I and G concentrations. The dual sensing approach significantly simplifies the operation by eliminating the need for lengthy incubation and washing steps for removing unbound signaling antibodies, which greatly enhances the speed compared to existing antibody-based affinity I assays. Such greatly reduced I assay time reflects the rapid and selective binding and instantaneous conformational-induced signaling inherent to the E-AB assay. The G-I bioelectronic device is designed as a four-electrode sensor chip 100, incorporating a combined Ag / AgCl reference electrode (RE) 108, two Au working electrodes (WE) 104 and 106, and one Au counter electrode (CE) 110 (FIG. 1A). The device was diced on polyethylene terephthalate glycol (PETG), and sputtering was used to produce thin Au and Ag films (FIGS. 2A-2C) for the WEs 104 and 106, RE 108, and CE 110. The G sensor WE 104 employed a second-generation enzymatic amperometric biosensor using a biocatalytic reagent layer 112 of the electron-carrying mediator (tetrathiafulvalene, TTF) and the glucose oxidase (GOx) enzyme on the Au WEI surface (FIGS. 1A-1B and 3A). The I sensor utilizes a methylene-blue (MB) modified guanine-rich (G-rich) thiolated aptamer receptor layer 114, coimmobilized with 6-mercapto-l -hexanol (MCH) on Au WE2 106, that exhibits insulin binding-induced conformation change that modulates the distance between an Au electrode surface and the MB redox indicator (FIGS. 1A, 1C, and 3A). Simultaneous G-I detection can thus be realized within 2 minutes using the same 30 pL sample droplet, with G measured amperometrically within the initial 90s, while the square wave voltammetry (SWV) detection of I requires an additional 30s to complete the G-I assay (FIG. 3A). The resulting biochip approach thus allows rapid simultaneous on-site measurements of insulin and glucose, and hence holds considerable promise for improved management of diabetes. Detailed optimization and characterization of the G-I sensing device, including dynamic range, selectivity, stability, and sensor crosstalk using biological fluids will be discussed in the following sections.PCT Application Attorney Docket No.: 009062.8573.WO00

[0065] Example Methods of Example Implementations

[0066] Materials and Apparatus. Insulin, glucose oxidase (GOx, from Aspergillus Niger, Type X-S (EC 1.1.3.4)), bovin serum albumin (BSA), chitosan (medium molecular weight), nafion (perfluorinated resin solution), tetrathiafulvalene (TTF, 97%), FeCh, 6-mercapto-l-hexanol (MCH), tris(2-carboxyethyl) phosphine hydrochloride (TCEP), 2-amino-2-(hydroxymethyl)-l,3-propanediol (Tris base), sodium chloride (NaCl), potassium chloride (KC1), magnesium chloride (MgCh), 10% sodium dodecyl sulfate (SDS), glucagon, hydrogen peroxide, uric acid and 1-ascorbic acid were purchased from Sigma- Aldrich. The insulin aptamer probe (5’-HO-(CH2)6-S-S-(CH2)6-AAAAGGTGGTGGGGGGGGTTGGTAGGGTGTCTTCT-MB-3’) (SEQ ID NO: 1) synthesized by Integrated DNA Technologies (IDT) Inc. (Coralville, IA) was used as received. Phosphate buffered saline (PBS, 4 mM, pH 7.4) was acquired from Gibco. HPLC grade 2-propanol and acetone, and hydrochloric acid (HC1) and nitric acid (HNO3) were purchased from Fisher Chemical. Human serum from undefined male was acquired from Bio reclamation IVT and used without any dilution. All electrochemical measurements were performed at room temperature using CH instruments (CHI1230A), an EmStat3 Blue potentiostat (PalmSens, The Netherlands) controlled by PSTrace software, version 5.9.

[0067] Chips Fabrication. The disposable single-use metal electrode arrays with two Au WEs 104 and 106 (15 mm2area) for glucose and insulin detection, a Au CE 110 and a Ag / AgCl RE 108 (FIG. 1A) were designed as below. For this process, a 1.0-mm-thick PETG (glycol-modified polyethylene terephthalate) sheet (Small Parts Inc.) was used as the substrate 102. In step 204, a Cricut cutting machine was used to generate the electrode pattern design on protective laminated cover of the PETG substrate 102. In step 208. metal electrodes were obtained by sequential deposition of Cr (300 W for 6 min) and Au (100 W for 15 min) and Ag (100 W for 15 min) layers in direct current (DC) mode for Cr and Au and in radio frequency (RF) mode for Ag (FIG. 2A). The process deployed an Ar gas pressure of 2.4 mTorr, using a Denton Discovery 635 sputter system. In step 210A the Ag layer on the WEs 104 and 106 and CE 110 was removed by reacting 5 pL of 6 M HNO3 solution for 5 min, and in step 212A the Ag surface of the RE 108 was subsequently modified with 5 pL of FeCh to obtain the Ag / AgCl-based layer for 2 min. The electrodes were cleaned before modifications by immersing them in solutions of isopropyl alcohol and water, each for 10 minutes (FIG. 2C).

[0068] Initially, WEI 104 was modified for the fabrication of the glucose biosensor, asPCT Application Attorney Docket No.: 009062.8573.WG00shown in FIG. 2C. Tn step 214, 2 pL of TTF was deposited on the electrode surface. Next, in step 216, 6 pL of a homogeneous GOx(BSA) was applied. After GOx dropcasting, in step 218, 2 pL Chitosan mixture was applied. Finally, WEI 104 was modified with 2 pL of Nafion and allowed to dry. Once the modification of WEI 104 for glucose was completed, WE2 106 was modified for insulin detection in step 220 by applying 15 pL of an aptamer solution for insulin at a concentration of 0.5 pM to the electrode surface. The multiplexed chips for glucose and insulin were then incubated overnight at about 4°C. After immobilization of the aptamer, only the surface of WE2 106 was gently washed by applying 10 pL of ultrapure water twice. Next, in step 222, the surface of WE2 106 was passivated by deposition of 10 pL of 30 mM MCH for 5 hours at 4°C. Finally, 5 pL of 10% SDS was applied to WE2 106 for 15 min. The multiplexed chips for glucose and insulin detection were then ready for analysis.

[0069] For dual analyte detection, 30 pL of G-I sample was dropped on the G-I chip (FIG.3A-3B). After 60 s incubation, the amperometric response was recorded on G sensor at +0.3 V (vs Ag / AgCl) for 30 s. Following G measurement, an additional 30 s incubation (total of 2 minutes) was required for the affinity binding of insulin with MB-tagged aptamer on I sensor. Then, SWV voltammetric response of MB signal was recorded on I sensor using frequency of 100 Hz with an amplitude of 25 mV. The same procedure was performed both for human serum and saliva measurements.

[0070] Example Results of Example Implementations

[0071] Integrating enzymatic and aptamer assays onto a single reagent-less chip for concurrent G and I sensing requires careful attention to major fabrication and operational challenges. These challenges include: i) the distinct sensor modification and immobilization chemistry at the neighboring enzymatic and aptamer sensors; ii) the differing assay formats and signal transduction mechanisms (enzymatic / amperometric for G and aptamer / voltammetric for I); iii) the vast difference in the physiological concentrations of I (pM) and G (mM); iv) the necessity for a compact, low-cost chip capable of performing simultaneous bioassays with minimal (microliter) sample volumes; v) minimizing cross-talk and cross-reactivity between the adjacent biocatalytic and bioaffinity sensors in the chip; and vi) the need for short analysis times in rapid testing to provide timely information for ultra-low target (I) concentrations. The fabrication and working principle of the dual G-I sensor chip are illustrated in FIGS. 2A-2C, and 3A-3B, respectively. The three-layer metal-sputtered (Cr-Au-Ag) bioelectronic G-I chip 100PCT Application Attorney Docket No.: 009062.8573.WO00includes two Au working electrodes, 104 and 106, for detecting G and I analytes, combined with an Ag / AgCl RE 108 and an Au CE 110. The glucose and insulin assays rely on specific biocatalytic and bioaffinity layers with distinct surface immobilization strategies on the two adjacent Au electrodes 104 and 106. These strategies involve precise spatial separation of the electrodes and the use of ultra small sample volumes. The fabrication of the G sensor involved a layer-by-layer modification of the WEI 104 Au surface with the electron-transfer mediator TTF and GOx enzyme to form a layer 112, confined within a capping layer of chitosan and a sulfonated tetrafluoroethylene polymer (e.g., Nafion) film (e.g., embodiment of capping layer 312) for highly selective glucose detection. In the presence of target G, the enzymatic redox reaction involves the Au electrode 104, the electron-carrying TTF mediator, and the GOx enzyme, which amplifies the oxidation currents (FIG. IB). The I sensor was fabricated by immobilizing an insulin-specific guanine-rich aptamer onto the WE2 106 Au surface through the formation of a self-assembled monolayer 114 containing the thiolated aptamer and MCH. The insulin binding (recognition) leads to conformational change of the aptamer that increases the distance of the MB tag from the surface, and results in smaller MB reduction currents compared to the unbound state (FIG. 1C). As illustrated in FIG. 3A, simultaneous detection of G and I with a single sample droplet was achieved in 2 minutes. G was quantified amperometrically at +0.3 V on WEI 104 through TTF-mediated enzymatic oxidation in layer 112 within the initial 90 s, including a 60 s incubation for the enzymatic reaction. During G measurement, the I target in the same sample droplet was bound to the aptamer layer 114 on WE2 106. Following the G measurement, an additional 30 s incubation (totaling 2 minutes) was required for continuing affinity binding of insulin towards the wash-free ultrasensitive voltammetric I detection based on changes in the reduction current of the MB tag.

[0072] G and I Sensor Optimization and In-Vitro Performance. In order to achieve sensitive G and I detection, the experimental variables in biocatalytic and affinity reactions and of the corresponding electrochemical transductions were thoroughly examined. The individual G and I bioassays (FIG. 4A and FIG. 5A) were optimized and characterized first.

[0073] FIGS. 4A-4F show schematics and data plots depicting optimization of the individual glucose (G) sensors.

[0074] FIG. 4A shows a schematic of a single (G) sensor platform and the biocatalytic reaction involved in the G detection using a single analyte chip. This G sensor chip 400 operatedPCT Application Attorney Docket No.: 009062.8573.WG00as a two-electrode system, which included a substrate 402 with an Au WE 404 modified with a TTF-GOx layer 412, and an Ag / AgCl combined reference / counter electrodes (RE / CE) 408.

[0075] FIG. 4B shows a data plot depicting cyclic voltammetry curves of the G sensor (at 100 mV / s), performed in the absence (black curve, 420) and in the presence of 1 mM glucose (red curve, 422). The electrochemical behavior of G sensor was examined in PBS solution by running cyclic voltammogram at the TTF-GOx modified electrode. As expected, in the presence of G, enzymatic transduction amplifies the oxidation currents of TTF, while in the absence of G, TTF oxidation currents were remarkably low.

[0076] FIGS. 4C and 4D show data plots depicting chronoamperometric response of the G sensor at + 0.3 V, examining the effect of the enzymatic reaction time, from 0 (i) to 90 s (xii) in 10 s increment, for detecting 10 mM G in PBS, to determine the incubation time that yields the maximum G current response. An example preferred waiting time was found to be 60 s although higher current magnitudes are observed until 60 s.

[0077] FIGS. 4E and 4F show data plots depicting chronoamperometric response of the G sensor at + 0.3 V, examining the response to increasing G concentrations, from 0 (i) to 15 mM (vii) in 2.5 mM increments, to evaluate the linearity and detection limit of the G sensor. Well-defined currents signals were observed for each increment (FIG. 4E); the relative standard deviation (RSD) of < 2.6% obtained for each G concentration demonstrates the consistency of the G sensor fabrication. The calculated limit of detection (EOD) from the constructed calibration plots in PBS (FIG. 4F) was found to be 47 pM, indicating high sensitivity for G monitoring in both serum and saliva (see Table 1).

[0078] Table 1 shows analytical characteristics of the independent G sensor chips for the detection of different G concentrations in PBS buffer using 1 min incubation. The Eimit of Detection (EOD) values were calculated using the 3Sb / m criterion, where Sb is the standard deviation for blank responses and m is the slope value of the calibration plot.Table 1PCT Application Attorney Docket No.: 009062.8573.WO00

[0079] FIG. 5A shows a schematic of a single (I) sensor platform. Similar systematic optimization was conducted in parallel for the I voltammetric sensor. The I sensor chip operated as a three-electrode system including an Au WE 506, Au CE 510 and Ag / AgCl RE 508. The insulin-specific aptamer used in the layer 514 has a high inclination for G-quadraplex formation. The aptamer packing heterogeneity and interstrand interactions are key for sensor performance and affinity binding with the target I. Especially at higher aptamer density, the G-rich aptamer favors forming interstrand G-quadraplex conformations over intrastrand G-quadraplex, which can affect insulin binding.

[0080] FIG. 5B shows a data plot of a cyclic voltammogram obtained in 0.05 M, H2SO4 solution (performed between 0 and 1.8 V at 100 mV / s), to increase the microscopic surface area of the electrode and determine the electroactive surface area of the gold electrode.

[0081] FIG. 5C shows a data plot of a cyclic voltammogram performed between 0 and -0.5 V to quantify the number of aptamers modified with the redox reporter on the electrode, integrating the area under the peak associated with methylene blue reduction. The aptamer-probe density was optimized, based on the data shown in FIGS. 5B-5C, and the calculated surface density of the aptamer utilized for all experiments was 2.26xl012molecules / cm2.

[0082] In the example implementations, the SWV technique was deployed for insulin monitoring, observing the electron transfer rate of redox mediator MB. SWV is a highly attractive electrochemical technique for aptamer-based sensors, providing high sensitivity and tunable parameters for optimal sensor performance.

[0083] FIG. 5D shows data plots illustrating the exemplary results of the study of the SWV frequency for the E-AB sensor for Insulin, (a) Profile of current changes as a function of SWV frequency for the E-AB sensor for insulin, before (black line, 520) and after addition of 100 nM insulin (red line, 522). (b) SWV current responses of sensor I at different frequencies: black curve (no insulin) and red curve (100 nM insulin) at: (i) 5 Hz; (ii) 20 Hz; (iii) 50 Hz; (iv) 100 Hz; (v) 200 Hz; (vi) 300 Hz; (vii) 400 Hz; (viii) 500 Hz. The SWV frequency parameter was optimized by interrogating frequencies from 5-500 Hz to obtain high current change, with 100 Hz offering the best response because of its maximum signal inhibition with respect to thePCT Application Attorney Docket No.: 009062.8573.WO00baseline.

[0084] FIG. 5E shows a data plot depicting SWV voltammetric current responses of I sensor towards 100 nM I at every 30 s from 0 s (data set 530) to 180 s (data set 536).

[0085] FIG. 5F shows a data plot depicting a corresponding current changes profile (black line 540: blank, green line 542:10 nM insulin, red line 544:100 nM insulin).

[0086] Following the optimization of SWV operation parameters, the aptamer response rate of the developed I sensor was tested with 10 and 100 nM I in every 30 s. I binding with aptamer was rapid and the MB reduction currents were saturated in 90 s (data set 533, FIG. 5F). An example preferred waiting time for the I sensor was found to be 120 s for measuring both high and low concentrations (FIG. 5F).

[0087] FIG. 5G shows a data plot depicting voltammetric current response of the E-AB sensor to increasing I concentrations from 0 nM (data set 550) to 50 nM (data set 554) in PBS with 2 min incubation.

[0088] FIG. 5H shows the calibration plot corresponding to FIG. 5G. The mean and the error bars represent the standard deviation of the measurements performed with three different chips.

[0089] In addition to the tunability of SWV and insulin response with the aptamer, the analytical performance of the I sensors was evaluated with different I increments from 0 to 50 nM, showing increasing diminutions of the SWV MB reduction currents, consistent with E-AB operations (FIG. 5G). To ensure that these current diminutions are not due to interfering analytes in PBS, blank additions (data set 560) followed by insulin additions (data set 562) were performed (FIG. 5H). The overall RSD values of < 4.5% for each I concentration indicate that the aptamer I sensor is highly reproducible. The measured current difference (Ai), before and after insulin binding, was proportional to the insulin increments, whereas negligible signal changes were obtained with blank additions (FIG. 5H). The LOD of the I sensor from the calibration plot in PBS was found to be 800 pM (see Table 2), meeting the need for ultrasensitive I detection. Table 3 compares the developed E-AB I sensor with other aptamer-based sensors for insulin detection, showing that the developed sensor can detect insulin with a low LOD.

[0090] Table 2 shows analytical characteristics of the independent I sensor chips for the detection of different concentrations of insulin standards in PBS buffer solutions using 2-min incubation.PCT Application Attorney Docket No.: 009062.8573.WO00Table 2

[0091] Table 3 shows comparison of insulin aptamer sensors.Table 3

[0092] Interference and Cross Talk Evaluation of the G-I Chip. To achieve ultrasensitive accurate detection with the dual G-I chip, it is crucial to evaluate the performance of the multiplexed sensor in the presence of interfering species that could compromise its selectivity for glucose and insulin. The presence of electroactive species, such as ascorbic acid (AA) and uricPCT Application Attorney Docket No.: 009062.8573.WG00acid (UA), in biological samples can pose a significant challenge for amperometric biosensors. For the G sensor, AA and UA were thus evaluated as potential interferents, considering the biocatalytic oxidation of TTF at +0.3V.

[0093] FIGS. 6B-6D shows data plots depicting interference and crosstalk evaluation of the example embodiment of the G-I chip. Amperometric measurements were thus conducted in the presence of 100 pM AA (dataset 610), 100 pM UA (dataset 611), and 5 mM G (dataset 612) at +0.3 V.

[0094] FIG. 6A shows a schematic of an example G-I chip illustrating selectivity and influence of electroactive interfering species of G-I chip.

[0095] FIG. 6B shows data plots depicting the amperometric response of ascorbic acid (100 pM AA, dataset 610) and uric acid (100 pM UA, dataset 611) in 1 mM G sensing.

[0096] FIG. 6C shows data plots depicting a histogram corresponding to the data depicted in FIG. 6B with three independent chips. The measured currents in the presence of interfering species are negligible compared to the G response, reflecting the highly selective G detection.

[0097] A similar selectivity evaluation was conducted for the E-AB I sensor. Bovine serum albumin (BSA), hydrogen peroxide, and glucagon were selected as potential interfering species, in addition to UA and AA. Glucagon, being a peptide hormone like I, could potentially cause a non-specific response with aptamer, thereby reducing selectivity of the sensor.

[0098] FIG.6D shows data plots depicting a SWV current responses with a blank (dataset 620), 100 pM of AA and UA (dataset 621), H2O2 (dataset 624), 100 nM of glucagon (dataset 623), BSA (dataset 622), and 50 nM I (dataset 625) sensing. The potentially interfering compounds displayed a negligible change in the MB reduction signal compared to the well-defined I response, indicating that the developed E-AB sensor is highly selective towards I.

[0099] After carefully evaluating the selectivity of the G and I sensors, it was important to examine the potential crosstalk between the G and I sensors to ensure accurate and simultaneous detection of the two target analytes. A rigorous cross-reactivity evaluation was conducted during sample incubation on each G-I dual chip, analyzing the response of both sensors in the presence of mixtures containing a fixed concentration of one analyte and varying concentrations of the other.

[0100] FIG. 7A displays the amperometric signals for G (blue, 710) and voltammetric signals for I (red, 712) on three different chips exposed to different G / I mixtures with varyingPCT Application Attorney Docket No.: 009062.8573.WO00concentrations. Specifically, the response of the three different G-T chips (Chip A, B, and C) was evaluated by gradually increasing the G concentration (2.5 mM, 5 mM, and 10 mM in Chips A, B, and C, respectively), while maintaining a fixed (10 nM) I level. This protocol produced a consistent I response (plots 730, 732, 734) and increasing G signals (plots 720, 722, 724), illustrating that the G quantitation is not affected by the presence of I and that the measurement of I is not affected by increasing G concentrations.

[0101] FIG. 7B presents the results of an experiment involving a G concentration of 10 mM (dataset 740) and increasing I concentrations (dataset 742) (5 nM, 100 nM, and 300 nM in Chips A, B, and C, respectively). The G amperomteric response is not affected by the rising I levels (plots 750, 752, 754) while the SWV I response is not influenced by the presence of G signal and remained distinctive (plots 760, 762, 764). In all cases, the signals obtained were clear and well-defined where the changing G or I concentrations did not affect the signals of the other target analyte, confirming a negligible crosstalk between the two sensors. These findings reinforce the specificity of the dual G-I chip for simultaneous measurements of G and I, validating the optimized analysis protocol. The results also highlight the robustness and reliability of the dualanalyte device for advanced diagnostic applications in clinical settings.

[0102] Glucose and Insulin Monitoring with Real Biofluids. A great utility of the exemplary dual G-I chip configuration is its ability to rapidly screen these key diabetes biomarkers using a single droplet of biofluid. The applicability of this dual sensor for on-site analysis in human serum and saliva was evaluated and demonstrated.

[0103] FIGS. 8A-8C and 9A-9C shows data plots depicting the electrochemical sensing performance of the example G-I chip of the example embodiment of the multi-analyte sensor device in serum and saliva samples.

[0104] Firstly, the analytical performance of G-I chip was assessed in undiluted serum spiked with increasing concentrations of G (3 mM increments ranging from 0 to 15 mM) and I (25 nM increments ranging from 0 to 125 nM) (FIGS. 8A and 8C). The G sensor demonstrated well-defined current signals to these 3 mM G additions, that lead to a well-defined calibration plot (FIG. 8B). Similarly, well-defined SWV current changes of the MB tag are observed for the 25 nM I additions to the serum sample, leading to a defined I calibration plot (FIG.8C). The calculated LOD for these serum experiments were 227 pM for G and 13.9 nM for I, based on the corresponding calibration plots (Table 4).PCT Application Attorney Docket No.: 009062.8573.WG00

[0105] FIG. 8 A shows a data plot depicting the amperometric response of the G-I chip for increasing concentration of glucose in undiluted serum from 0 (i) to 15 mM (vi) in 3 mM increments.

[0106] FIG. 8B shows a data plot depicting a calibration curve for the chronoamperometric response of the G sensor in the G-I chip, obtained for increasing G concentrations from 0 to 15 mM in 5 mM increments in undiluted human serum.

[0107] FIG. 8C shows a data plot depicting the voltammetric response of the G-I chip for increasing insulin concentration from 0 to 125 nM (using 25 nM increments) in undiluted human serum.

[0108] Table 4 shows analytical characteristics of the developed G-I dual sensor chip for the detection of different concentrations of glucose and insulin standards spiked in serum samples. The LOD value was calculated using the 3Sb / m criterion, where Sb is the standard deviation for blank responses and m is the slope value of the calibration plot.Table 4

[0109] A similar set of experiments was subsequently performed for the G-I sensor in human saliva (FIGS. 9A-9C). Saliva samples were collected using a salivette tube and directly applied to the G-I chip without further dilution. A 30 pL drop of saliva was spiked to capture the physiological salivary G range with three increasing G concentrations of 1 mM. Well-defined amperometric currents were observed with each increment, closely matching the concentration profiles observed in PBS. The consistency and viscosity of centrifuged saliva were similar to PBS, allowing comparable diffusion G profiles to the sensor surface. Calibration plots for the I sensor in saliva were generated with increasing concentrations of insulin (25 nM increments), resulting in a well-defined SWV I calibration plot (FIG. 9C). The insulin E-AB sensor exhibited reproducible performance in saliva, similar to those observed in buffer and serum, demonstratingPCT Application Attorney Docket No.: 009062.8573.WG00the sensing capabilities and robustness of the dual chip across various biofluids for practical applications. The LOD for the G-I chip in saliva were 245 pM for G and in the 1.9 nM for I (Table 5).

[0110] FIG.9A shows a data plot depicting the amperometric response of the G-I chip for increasing glucose concentration from 0 (i) to 7 mM (vii) in 1 mM increments.

[0111] FIG. 9B shows a data plot depicting a calibration curve for the chronoamperometric response of the G sensor in the G-I chip, obtained for increasing G concentrations from 0 to 7 mM in 1 mM increments in human saliva.

[0112] FIG. 9C shows a data plot depicting the voltammetric response of the G-I chip to increasing insulin concentration from 0 to 125 nM (using 25 nM increments) in undiluted human saliva.

[0113] Table 5 shows analytical characteristics of the developed G-I dual sensor chip for the detection of different concentrations of glucose and insulin standards spiked in saliva samples. The LOD value was calculated using the 3Sb / m criterion, where Sb is the standard deviation for blank responses and m is the slope value of the calibration plot.Table 5

[0114] To further validate the exemplary G-I sensor in real biofluids, saliva samples were collected from three different subjects. All subjects showed similar trends of G and I results.

[0115] FIG. 10 shows a data plot depicting the performance of the G-I chip in G and I spiked human saliva samples of three different subjects and the current changes calculated from the amperometric (G, bottom and left axes, datasets 1002, 1004, 1006) and voltammetric (I, top and right axes, datasets 1008, 1010, 1012) curves. The G y-axis (left axis) is reported as the change in current magnitude due to the existing G concentration (bottom axis) in the subjects’ saliva samples. Although the sensitivity of the G-I chip was lower in blood serum and saliva comparedPCT Application Attorney Docket No.: 009062.8573.WO00to buffer solution, most likely due to potential matrix and fouling effects, the chip exhibits a promising biosensing performance. The biochip is expected to readily quantify G and I in additional various biofluids, such as interstitial fluid and liquid biopsies.

[0116] In summary of the example implementations discussed above, a new biochip strategy has been demonstrated for coupling aptamer and enzymatic assays and its applicability to simultaneous measurements of insulin and glucose within 2 minutes using a single microliter sample droplet. By eliminating traditional time-consuming I immunoassays and adopting a reagent-less rapid aptamer-based SWV I detection, this work takes a significant step toward the rapid, concurrent POC detection of key diabetes biomarkers. The attractive analytical performance of the resulting dual G-I chip has been realized through careful attention to key challenges of combining different immobilization chemistries and detection / transduction principles (for the corresponding biocatalytic and affinity reactions) toward judicious integration of such two fundamentally different assays onto a single device and avoiding sensor crosstalk between the closely spaced sensing electrodes, considering also the tremendous different concentrations of the target I and G analytes (pM vs. mM). Future development of such multiplexed biomarker assays may focus on further enhancing the sensitivity of the I aptamer E-AB sensor towards the physiological range of insulin in diverse body fluids via highly selective aptamer from the systematic evolution of ligands by exponential enrichment (SELEX) process, introducing multiple redox label placement on the surface-tethered aptamer and using a nanostructured gold transducer surface to improve the electron transfer rate kinetics. Future efforts may also involve further miniaturization and integration of the supporting electronic interface and a large-scale clinical validation. The new aptamer / enzyme-based biochip approach can be extended to the simultaneous on-site detection of additional diabetes biomarkers towards enhancing personalized diabetes management. Such integration of enzymatic and aptamer-based bioassays could greatly expand the scope and horizons of decentralized diagnostics in healthcare, beyond diabetes care, considering the broad range of biomarkers that can be detected by such coupling.

[0117] Real Saliva Collection. Saliva samples from authors were collected using Sarstedt Salivettes (Sarstedt, Numbrecht, Germany). Swabs were chewed for 2 min, after which they were inserted into the upper cavity of the Salivette and centrifuged immediately for 2 min at 1000 g to obtain saliva available for laboratory use.PCT Application Attorney Docket No.: 009062.8573.WG00

[0118] Glucose Sensor Fabrication. The glucose sensor was initially developed by deposition of a TTF mediator layer on a gold electrode (Au WE), followed by application of a solution including GOx prepared in BSA, chitosan, and nation. The 50 mM TTF solution was prepared by combing TTF with a nine parts ethanol to one part acetone mixture (9:1). The chitosan solution was prepared in 0.1 M acetic acid and stirred magnetically overnight to obtain a homogeneous solution. GOx (60 U) was dissolved in a PBS solution containing 10 mg / mE BSA. The GOx(BSA) / Chitosan mixture (3:1 V / V) was applied to the TTF-modified electrode and allowed to dry at room temperature. Finally, two 1 pF layers of a nation (2%) solution were added to the electrode surface and allowed to dry for 4 hours at 4 °C.

[0119] Insulin Sensor Fabrication. The fabrication of the insulin sensor includes several steps. Before immobilizing the probe on the working electrode (WE), thiol reduction was performed with 10 mM TCEP for 1 h at room temperature (dark environment) to allow covalent binding to the electrode surface. In detail, 1.5 pL of a 100 pM insulin aptamer solution was mixed with 2 pL of 10 mM TCEP. Next, the solution was diluted with 50 mM Tris buffer containing 10 mM KC1, 100 mM NaCl and 50 mM MgCh, pH 8. The 1 pM solution of the insulin aptamer was heated to about 90 °C for 5 min and then gradually cooled to about 4 °C. This annealing is beneficiary for the formation of intrastrand G-quadruplex structural motifs and that the long passivation times help with forming a stable self-assembled monolayer. Next, the solution of the insulin aptamer probe was further diluted to 0.5 pM with Tris buffer, and 15 pL of the probe was placed on the gold surface of the working electrode. Subsequently, the electrodes were incubated overnight at ~4 °C. After immobilization of the aptamer, the electrodes were thoroughly washed with ultrapure water and passivated with 15 pF of 30 mM MCH for 5 h at ~4 °C.

[0120] Determination of Aptamer Packing Density. The performance of an aptamer-based sensor (E-AB) is strongly influenced by the packing density of aptamers on the electrode, i.e. the number of aptamer molecules per unit area. This density is controlled by varying the concentration of the thiolated oligonucleotide during sensor manufacture and is measured to ensure reproducible performance.

[0121] The first step in determining the packing density of aptamers is to calculate the electroactive surface area of the gold working electrode. To do this, a cyclic voltammogram was performed between 0 and 1.8 V at 0.1 V / s using a 0.05 M sulphuric acid solution (see FIGS. 5BPCT Application Attorney Docket No.: 009062.8573.WO00and 5C). By integrating the area under the peak associated with the reduction of the gold oxide and dividing the result by 400 pC / cm2, which represents the charge density for the formation of a complete oxygen monolayer chemisorbed on the gold, the electroactive surface area of the electrode was determined to be 0.17 cm2.

[0122] To calculate the density of aptamer probes on the electrode surface, expressed as T, the cyclic voltammetry peaks were integrated (using slow scan rates such as 20, 50 and 100 mV / s), using an insulin aptamer concentration of 500 nM, according to the following equation:T = Q / nFA (1)

[0123] In this equation, Q represents the total charge consumed and produced during the scan of the cyclic voltammetry (expressed in mC), n is the number of electrons transferred in the reaction (for the MB it is equal to 2), F is the Faraday constant (expressed in C / mol) and A represents the area of the electroactive surface of this sensor (in cm2), as calculated above. The value of T obtained is represented by the average value of the voltammograms obtained using the three different scanning speeds mentioned above. This protocol produces packing densities of 2.26x1012molecules / cm2.Examples

[0124] In some embodiments in accordance with the present technology (example 1), a sensor device for rapid and simultaneous detection of two or more analytes includes a substrate; a first analyte sensor comprising a first working electrode disposed on the substrate and a first functionalization layer disposed on the first working electrode, wherein the first functionalization layer includes a catalyst to facilitate an enzymatic electrochemical reaction to detect a first analyte at the first analyte sensor; and a second analyte sensor comprising a second working electrode disposed on the substrate and a second functionalization layer disposed on the second working electrode, wherein the second functionalization layer includes an aptamer receptor to facilitate an aptamer affinity bioassay reaction to detect a second analyte at the second analyte sensor. In some embodiments of the sensor device of example 1, the sensor device may optionally include a counter electrode to provide a current path with respect to one or both of the first working electrode and the second working electrode during an electrochemical detection technique; and a reference electrode to provide a fixed electrical potential for reference to measure an electrical potential at one or both of the first working electrode and the second working electrode during an electrochemical detection technique.PCT Application Attorney Docket No.: 009062.8573.WG00

[0125] Example 2 includes the sensor device of example 1 or any of examples 1-19, wherein the catalyst of the first functionalization layer includes glucose oxidase (GOx), and wherein the GOx is immobilized inside a permeable polymer film.

[0126] Example 3 includes the sensor device of example 2 or any of examples 1-19, wherein the permeable polymer film comprises a tetrathiafulvalene (TTF) mediator layer formed on the first working electrode and a capping layer.

[0127] Example 4 includes the sensor device of example 3 or any of examples 1-19, wherein the capping layer comprises chitosan and a sulfonated tetrafluoroethylene polymer.

[0128] Example 5 includes the sensor device of example 1 or any of examples 1-19, wherein the aptamer receptor comprises at least one of a deoxyribonucleic acid (DNA) oligonucleotide, a ribonucleic acid (RNA) oligonucleotide, or a peptide nucleic acid (PNA) oligonucleotide.

[0129] Example 6 includes the sensor device of example 5 or any of examples 1-19, wherein the aptamer receptor includes 5’-HO-(CH2)6-S-S-(CH2)6- AAAAGGTGGTGGGGGGGGTTGGTAGGGTGTCTTCT-MB-3’ (SEQ ID NO: 1).

[0130] Example 7 includes the sensor device of example 5 or any of examples 1-19, wherein the aptamer receptor is modified with a redox indicator that includes methylene blue.

[0131] Example 8 includes the sensor device of example 1 or any of examples 1-19, wherein the first working electrode and the second working electrode include gold.

[0132] Example 9 includes the sensor device of example 1 or any of examples 1-19, wherein the counter electrode includes silver / silver chloride (Ag / AgCl).

[0133] Example 10 includes the sensor device of example 1 or any of examples 1-19, wherein the counter electrode is configured to surround at least 50% of a periphery of the first working electrode and surround at least 20% of a periphery of the second working electrode; and wherein the reference electrode is configured to surround at least 30% of the periphery of the second working electrode.

[0134] Example 11 includes the sensor device of example 1 or any of examples 1-19, wherein the substrate includes a plastic.

[0135] Example 12 includes the sensor device of example 11 or any of examples 1-19, wherein the plastic includes one or more of polyethylene terephthalate glycol (PETG), polyethylene terephthalate (PET), polyethylene naphthalate (PEN), or polyimide (PI).

[0136] Example 13 includes the sensor device of example 1 or any of examples 1-19,PCT Application Attorney Docket No.: 009062.8573.WG00wherein the first analyte includes glucose and the second analyte includes insulin.

[0137] Example 14 includes the sensor device of example 1 or any of examples 1-19, wherein the first analyte and the second analyte include an analyte selected from a group consisting of glucose, ketone, lactate, cortisol, uric acid, vitamin C, acetaminophen, levodopa, sodium ion, potassium ion, chlorine ion, calcium ion, ammonium, alcohol, caffeine, insulin, glucagon, cytokines, and a cancerous biomarker.

[0138] Example 15 includes the sensor device of example 1 or any of examples 1-19, further comprising a third analyte sensor comprising a third working electrode disposed on the substrate and a third functionalization layer disposed on the third working electrode, wherein the third functionalization layer includes a second catalyst to facilitate a second enzymatic electrochemical reaction to detect a third analyte at the third analyte sensor; and a fourth analyte sensor comprising a fourth working electrode disposed on the substrate and a fourth functionalization layer disposed on the fourth working electrode, wherein the fourth functionalization layer includes a second aptamer receptor to facilitate a second affinity bioassay reaction to detect a fourth analyte at the fourth analyte sensor.

[0139] Example 16 includes the sensor device of example 15 or any of examples 1-19, wherein the third analyte includes ketone and the fourth analyte includes insulin.

[0140] Example 17 includes the sensor device of example 1 or any of examples 1-19, comprising an electronics unit including an electric circuit coupled to a data processing unit, wherein, when the sensor device is electrically coupled to the electronics unit, the device is operable to detect the first analyte and the second analyte and determine a concentration of the first analyte and of the second analyte.

[0141] Example 18 includes the sensor device of example 1 or any of examples 1-19, wherein the electrochemical detection technique includes amperometry and voltammetry simultaneously.

[0142] Example 19 includes the sensor device of example 1 or any of examples 1-19, wherein sensor device is operable to detect the first analyte and the second analyte in 2 minutes or less.

[0143] In some embodiments in accordance with the present technology (example 20), a method for rapid, simultaneous monitoring of two or more analytes in a biofluid sample includes providing a reagent-less sensor device that detects a first analyte by an enzymaticPCT Application Attorney Docket No.: 009062.8573.WG00electrochemical detection technique and detects a second analyte by an aptamer-based affinity bioassay technique.

[0144] Example 21 includes the method of example 20 or any of examples 20-26, wherein the method is able to simultaneously measure the first analyte and the second analyte in the biofluid sample without requiring additional washing steps or external reagents.

[0145] Example 22 includes the method of example 20 or any of examples 20-26, wherein the method is able to simultaneously measure the first analyte and the second analyte in the biofluid sample within 2 minutes or less.

[0146] Example 23 includes the method of example 22 or any of examples 20-26, wherein the first analyte is glucose and the second analyte is insulin, and wherein the method is able to simultaneously measure the glucose and the insulin in the biofluid sample within 2 minutes or less.

[0147] Example 24 includes the method of example 23 or any of examples 20-26, wherein the method is able to measure the glucose in the biofluid sample within 1 minute or less.

[0148] Example 25 includes the method of example 20 or any of examples 20-26, wherein the biofluid sample includes a single microliter sample droplet of whole blood, serum, saliva, sweat, urine, or interstitial fluid.

[0149] Example 26 includes the method of any of examples 20-25, wherein the reagent-less sensor device includes the sensor device of any of examples 1-19.

[0150] In some embodiments in accordance with the present technology (example Bl), a sensor device for rapid detection of two or more analytes includes a substrate; a first analyte sensor comprising a first working electrode disposed on the substrate and a first functionalization layer disposed on the first working electrode, wherein the first functionalization layer includes an enzyme configured to facilitate a first electrochemical reaction enabling detection of a first analyte at the first analyte sensor; a second analyte sensor comprising a second working electrode disposed on the substrate and a second functionalization layer disposed on the second working electrode, wherein the second functionalization layer includes a redox indicator-modified aptamer receptor configured to facilitate a second electrochemical reaction enabling detection of a second analyte at the second analyte sensor; a counter electrode configured to complete an electrical circuit with respect to one or both of the first working electrode and the second working electrode during an electrochemical detection technique; and a referencePCT Application Attorney Docket No.: 009062.8573.WG00electrode configured to provide a fixed electrical potential for reference to measure an electrical potential at one or both of the first working electrode and the second working electrode during the electrochemical detection technique.

[0151] Example B2 includes the sensor device of example Bl or any of examples B1-B21, wherein the first working electrode, the second working electrode, the counter electrode, and the reference electrode are arranged such that the counter electrode is adjacent to at least a portion of both the first working electrode and the second working electrode, and such that the reference electrode is positioned adjacent to the second working electrode.

[0152] Example B3 includes the sensor device of example B2 or any of examples B1-B21, wherein the counter electrode is adjacent to least 50% of a periphery of the first working electrode and at least 20% of a periphery of the second working electrode; and wherein the reference electrode is adjacent to at least 30% of the periphery of the second working electrode

[0153] Example B4 includes the sensor device of example Bl, B2, or B3 or any of examples B1-B21, wherein the first working electrode, the second working electrode, the counter electrode, and the reference electrode are configured such that at least a portion of a surface of each of the electrodes can be simultaneously covered by a 30 microliter drop of a biofluid.

[0154] Example B5 includes the sensor device of example Bl or any of examples B1-B21, wherein the first analyte includes glucose and the second analyte includes insulin.

[0155] Example B6 includes the sensor device of example Bl or any of examples B1-B21, wherein the first analyte includes the enzyme of the first functionalization layer includes glucose oxidase (GOx).

[0156] Example B7 includes the sensor device of example B6 or any of examples B1-B21, wherein the first functionalization layer further comprises tetrathiafulvalene (TTF) mediator layer formed on the first working electrode.

[0157] Example B8 includes the sensor device of example B7 or any of examples B1-B21, wherein the first electrochemical reaction comprises glucose oxidation and TTF reduction.

[0158] Example B9 includes the sensor device of example Bl or any of examples B1-B21, wherein the first analyte sensor further comprises a capping layer disposed on the first functionalization layer.

[0159] Example B10 includes the sensor device of example B9 or any of examples B1-B21, wherein the capping layer comprises chitosan and a sulfonated tetrafluoroethylene polymer.PCT Application Attorney Docket No.: 009062.8573.WG00

[0160] Example B11 includes the sensor device of example Bl or any of examples B1-B21, wherein the redox indicator-modified aptamer receptor comprises at least one of a deoxyribonucleic acid (DNA) oligonucleotide, a ribonucleic acid (RNA) oligonucleotide, or a peptide nucleic acid (PNA) oligonucleotide.

[0161] Example B12 includes the sensor device of example Bl 1 or any of examples B1-B21, wherein a redox indicator of the redox indicator-modified aptamer receptor includes methylene blue.

[0162] Example B13 includes the sensor device of example B12 or any of examples B1-B21, wherein the second electrochemical reaction comprises an electron transfer from methylene blue to the second working electrode.

[0163] Example B14 includes the sensor device of example B12 or any of examples B1-B21, wherein the redox indicator-modified aptamer receptor comprises a molecule represented by a chemical sequence HO-(CHI)6-S-S-(CH2)6- AAAAGGTGGTGGGGGGGGTTGGTAGGGTGTCTTCT-MB (SEQ ID NO: 1), wherein A represents adenine, G represents guanine, T represents thymine, C represents cytosine, and MB represents methylene blue.

[0164] Example B15 includes the sensor device of example Bl or any of examples B1-B21, wherein the first working electrode and the second working electrode include gold.

[0165] Example B16 includes the sensor device of example Bl or any of examples B1-B21, wherein the counter electrode includes silver / silver chloride (Ag / AgCl).

[0166] Example B17 includes the sensor device of example Bl or any of examples B1-B21, wherein the substrate includes a plastic.

[0167] Example B18 includes the sensor device of example B17 or any of examples B1-B21, wherein the plastic includes one or more of polyethylene terephthalate glycol (PETG), polyethylene terephthalate (PET), polyethylene naphthalate (PEN), or polyimide (PI).

[0168] Example B19 includes the sensor device of example Bl or any of examples B1-B21, wherein the first analyte and the second analyte include an analyte selected from a group consisting of glucose, ketone, lactate, cortisol, uric acid, vitamin C, acetaminophen, levodopa, sodium ion, potassium ion, chlorine ion, calcium ion, ammonium, alcohol, caffeine, insulin, glucagon, cytokines, and a cancerous biomarker.

[0169] Example B20 includes the sensor device of example Bl or any of examples B1-B21,PCT Application Attorney Docket No.: 009062.8573.WG00wherein the electrochemical detection technique includes at least one of amperometry and voltammetry.

[0170] Example B21 includes the sensor device of example Bl or any of examples B1-B20, wherein sensor device is operable to detect the first analyte and the second analyte in 2 minutes or less.

[0171] In some embodiments in accordance with the present technology (example B22), a method for rapid monitoring of two or more analytes in a biofluid sample includes providing a reagent-less sensor device that detects a first analyte by an enzymatic electrochemical reaction measured by a first electrochemical detection technique and detects a second analyte by an aptamer-based electrochemical reaction measured by a second electrochemical detection technique.

[0172] Example B23 includes the method of example B22 or any of examples B22-B28, wherein the method is configured to detect the first analyte and the second analyte in the biofluid sample without requiring washing steps or external reagents.

[0173] Example B24 includes the method of example B22 or any of examples B22-B28, wherein the first electrochemical detection technique comprises amperometry and the second electrochemical detection technique comprises voltammetry.

[0174] Example B25 includes the method of example B22 or any of examples B22-B28, wherein the method is configured to detect the first analyte and the second analyte in the biofluid sample within 2 minutes or less.

[0175] Example B26 includes the method of example B22 or any of examples B22-B28, wherein the first analyte is glucose and the second analyte is insulin, and wherein the method is able to detect the glucose and the insulin in the biofluid sample within 2 minutes or less.

[0176] Example B27 includes the method of example B22 or any of examples B22-B28, wherein the biofluid sample includes a single sample droplet of whole blood, serum, saliva, sweat, urine, or interstitial fluid.

[0177] Example B28 includes the method of any of examples B22-B28, wherein the reagentless sensor device includes the sensor device of any of examples B1-B21.Conclusion

[0178] Implementations of the subject matter and the functional operations described in this patent document can be implemented in various systems, digital electronic circuitry, or inPCT Application Attorney Docket No.: 009062.8573.WO00computer software, firmware, or hardware, including the structures disclosed in this specification and their structural equivalents, or in combinations of one or more of them. Implementations of the subject matter described in this specification can be implemented as one or more computer program products, i.e., one or more modules of computer program instructions encoded on a tangible and non-transitory computer readable medium for execution by, or to control the operation of. data processing apparatus. The computer readable medium can be a machine-readable storage device, a machine-readable storage substrate, a memory device, a composition of matter effecting a machine-readable propagated signal, or a combination of one or more of them. The term “data processing unit” or “data processing apparatus” encompasses all apparatus, devices, and machines for processing data, including by way of example a programmable processor, a computer, or multiple processors or computers. The apparatus can include, in addition to hardware, code that creates an execution environment for the computer program in question, e.g., code that constitutes processor firmware, a protocol stack, a database management system, an operating system, or a combination of one or more of them.

[0179] A computer program (also known as a program, software, software application, script, or code) can be written in any form of programming language, including compiled or interpreted languages, and it can be deployed in any form, including as a stand-alone program or as a module, component, subroutine, or other unit suitable for use in a computing environment. A computer program does not necessarily correspond to a file in a file system. A program can be stored in a portion of a file that holds other programs or data (e.g., one or more scripts stored in a markup language document), in a single file dedicated to the program in question, or in multiple coordinated files (e.g., files that store one or more modules, sub programs, or portions of code). A computer program can be deployed to be executed on one computer or on multiple computers that are located at one site or distributed across multiple sites and interconnected by a communication network.

[0180] The processes and logic flows described in this specification can be performed by one or more programmable processors executing one or more computer programs to perform functions by operating on input data and generating output. The processes and logic flows can also be performed by, and apparatus can also be implemented as, special purpose logic circuitry, e.g., an FPGA (field programmable gate array) or an ASIC (application specific integrated circuit).PCT Application Attorney Docket No.: 009062.8573.WO00

[0181] Processors suitable for the execution of a computer program include, by way of example, both general and special purpose microprocessors, and any one or more processors of any kind of digital computer. Generally, a processor will receive instructions and data from a read only memory or a random access memory or both. The essential elements of a computer are a processor for performing instructions and one or more memory devices for storing instructions and data. Generally, a computer will also include, or be operatively coupled to receive data from or transfer data to, or both, one or more mass storage devices for storing data, e.g., magnetic, magneto optical disks, or optical disks. However, a computer need not have such devices. Computer readable media suitable for storing computer program instructions and data include all forms of nonvolatile memory, media and memory devices, including by way of example semiconductor memory devices, e.g., EPROM, EEPROM, and flash memory devices. The processor and the memory can be supplemented by, or incorporated in, special purpose logic circuitry.

[0182] While this patent document contains many specifics, these should not be construed as limitations on the scope of any invention or of what may be claimed, but rather as descriptions of features that may be specific to particular embodiments of particular inventions. Certain features that are described in this patent document in the context of separate embodiments can also be implemented in combination in a single embodiment. Conversely, various features that are described in the context of a single embodiment can also be implemented in multiple embodiments separately or in any suitable subcombination. Moreover, although features may be described above as acting in certain combinations and even initially claimed as such, one or more features from a claimed combination can in some cases be excised from the combination, and the claimed combination may be directed to a subcombination or variation of a subcombination.

[0183] Similarly, while operations are depicted in the drawings in a particular order, this should not be understood as requiring that such operations be performed in the particular order shown or in sequential order, or that all illustrated operations be performed, to achieve desirable results. Moreover, the separation of various system components in the embodiments described in this patent document should not be understood as requiring such separation in all embodiments.

[0184] Only a few implementations and examples are described and other implementations, enhancements and variations can be made based on what is described and illustrated in thisPCT Application Attorney Docket No.: 009062.8573.WG00 patent document.

Claims

PCT Application Attorney Docket No.: 009062.8573.WO00CLAIMSWhat is claimed is:

1. A sensor device for rapid detection of two or more analytes, comprising:a substrate;a first analyte sensor comprising a first working electrode disposed on the substrate and a first functionalization layer disposed on the first working electrode, wherein the first functionalization layer includes an enzyme configured to facilitate a first electrochemical reaction enabling detection of a first analyte at the first analyte sensor;a second analyte sensor comprising a second working electrode disposed on the substrate and a second functionalization layer disposed on the second working electrode, wherein the second functionalization layer includes a redox indicator-modified aptamer receptor configured to facilitate a second electrochemical reaction enabling detection of a second analyte at the second analyte sensor;a counter electrode configured to complete an electrical circuit with respect to one or both of the first working electrode and the second working electrode during an electrochemical detection technique; anda reference electrode configured to provide a fixed electrical potential for reference to measure an electrical potential at one or both of the first working electrode and the second working electrode during the electrochemical detection technique.

2. The sensor device of claim 1, wherein the first working electrode, the second working electrode, the counter electrode, and the reference electrode are arranged such that the counter electrode is adjacent to at least a portion of both the first working electrode and the second working electrode, and such that the reference electrode is positioned adjacent to the second working electrode.

3. The sensor device of claim 2, wherein the counter electrode is adjacent to least 50% of a periphery of the first working electrode and at least 20% of a periphery of the second working electrode; and wherein the reference electrode is adjacent to at least 30% of the periphery of the second working electrode.PCT Application Attorney Docket No.: 009062.8573.WG004. The sensor device of any of claims 1 -3, wherein the first working electrode, the second working electrode, the counter electrode, and the reference electrode are configured such that at least a portion of a surface of each of the electrodes can be simultaneously covered by a 30 microliter drop of a biofluid.

5. The sensor device of claim 1, wherein the first analyte includes glucose and the second analyte includes insulin.

6. The sensor device of claim 1, wherein the first analyte includes the enzyme of the first functionalization layer includes glucose oxidase (GOx).

7. The sensor device of claim 6, wherein the first functionalization layer further comprises tetrathiafulvalene (TTF) mediator layer formed on the first working electrode.

8. The sensor device of claim 7, wherein the first electrochemical reaction comprises glucose oxidation and TTF reduction.

9. The sensor device of claim 1, wherein the first analyte sensor further comprises a capping layer disposed on the first functionalization layer.

10. The sensor device of claim 9, wherein the capping layer comprises chitosan and a sulfonated tetrafluoroethylene polymer.

11. The sensor device of claim 1, wherein the redox indicator-modified aptamer receptor comprises at least one of a deoxyribonucleic acid (DNA) oligonucleotide, a ribonucleic acid (RNA) oligonucleotide, or a peptide nucleic acid (PNA) oligonucleotide.

12. The sensor device of claim 11, wherein a redox indicator of the redox indicator-modified aptamer receptor includes methylene blue.

13. The sensor device of claim 12, wherein the second electrochemical reaction comprises an electron transfer from methylene blue to the second working electrode.

14. The sensor device of claim 12, wherein the redox indicator-modified aptamer receptor comprises a molecule represented by a chemical sequence HO-(CH2)6-S-S-(CH2)6-PCT Application Attorney Docket No.: 009062.8573.WO00AAAAGGTGGTGGGGGGGGTTGGTAGGGTGTCTTCT-MB (SEQ ID NO: 1), wherein A represents adenine, G represents guanine, T represents thymine, C represents cytosine, and MB represents methylene blue.

15. The sensor device of claim 1, wherein the first working electrode and the second working electrode include gold.

16. The sensor device of claim 1, wherein the counter electrode includes silver / silver chloride (Ag / AgCl).

17. The sensor device of claim 1, wherein the substrate includes a plastic.

18. The sensor device of claim 17, wherein the plastic includes one or more of polyethylene terephthalate glycol (PETG), polyethylene terephthalate (PET), polyethylene naphthalate (PEN), or polyimide (PI).

19. The sensor device of claim 1, wherein the first analyte and the second analyte include an analyte selected from a group consisting of glucose, ketone, lactate, cortisol, uric acid, vitamin C, acetaminophen, levodopa, sodium ion, potassium ion, chlorine ion, calcium ion, ammonium, alcohol, caffeine, insulin, glucagon, cytokines, and a cancerous biomarker.

20. The sensor device of claim 1, wherein the electrochemical detection technique includes at least one of amperometry and voltammetry.

21. The sensor device of claim 1, wherein sensor device is operable to detect the first analyte and the second analyte in 2 minutes or less.

22. A method for rapid monitoring of two or more analytes in a biofluid sample, comprising:providing a reagent-less sensor device that detects a first analyte by an enzymatic electrochemical reaction measured by a first electrochemical detection technique and detects a second analyte by an aptamer-based electrochemical reaction measured by a second electrochemical detection technique.

23. The method of claim 22, wherein the method is configured to detect the first analyte and the second analyte in the biofluid sample without requiring washing steps or external reagents.PCT Application Attorney Docket No.: 009062.8573.WO0024. The method of claim 22, wherein the first electrochemical detection technique comprises amperometry and the second electrochemical detection technique comprises voltammetry.

25. The method of claim 22, wherein the method is configured to detect the first analyte and the second analyte in the biofluid sample within 2 minutes or less.

26. The method of claim 22, wherein the first analyte is glucose and the second analyte is insulin, and wherein the method is able to detect the glucose and the insulin in the biofluid sample within 2 minutes or less.

27. The method of claim 22, wherein the biofluid sample includes a single sample droplet of whole blood, serum, saliva, sweat, urine, or interstitial fluid.

28. The method of any of claims 22-27, wherein the reagent-less sensor device includes the sensor device of any of claims 1-21.

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