WO2023178277A1 - Closed bipolar electrode based biosensors - Google Patents

Abstract

Devices and methods are disclosed for using a closed bipolar electrode system to detect and/or quantify the amount of an analyte.

Description

CLOSED BIPOLAR ELECTRODE BASED BIOSENSORS

CROSS-REFERENCE TO RELATED APPLICATION

[0001] This application claims the benefit of and priority to U.S. Provisional Application No. 63/269,591, filed on March 18, 2022, the entire contents of which are incorporated herein by reference.

BACKGROUND

[0002] Ion selective electrodes using potentiometric detection are among the most common types of electrochemical sensors across academia and industry. Ion selective electrodes convert the change in activity of one or more ions into a measurable potential (Zdrachek, 2019; Zdrachek, 2021). When used in a potentiometric system, ion selective electrodes are calibrated against a known reference electrode and used in a state where negligible current flows between the 2 electrodes. The most commonly used reference electrode is Ag/AgCl, an ion selective electrode, due to its ease of fabrication, low cost, and simple miniaturization (Sophocleuous, 2017). This electrode may be simply classified as a pseudo (without supporting electrolyte such as 3M KC1) or single junction electrode where the Ag/AgCl wire is inserted into a supporting electrolyte such as KC1 and is partitioned from the environment using a porous glass frit (Bard, 2001; Sopstad, 2018). Ag/AgCl electrodes have a potential of 0.222V with respect to the standard hydrogen electrode, which is produced by inserting a platinum wire or mesh into a compartment that has either sealed or continuous flow of hydrogen gas. The standard hydrogen electrode (SHE) is set to a value of “0.00V” and is treated as a universal baseline to which all reference electrodes are measured (Hu, 2016; Jerkiewicz, 2020).

[0003] There are several limitations to using conventional ion selective electrodes in this type of scheme. The reference electrode is often exposed to the same medium as the sample which can promote error when measuring analytes in a complex sample matrix (interstitial fluid, whole blood, saliva etc.). Generally, this error manifests itself through interference of the working and reference electrodes by exogenous/endogenous competing ions and/or chemicals (Hu, 2016; Sophocleous, 2017). Each ion has a maximum allowed voltage variability that meets the percent error target can be calculated using the Nemst equation (Burtis, 1994; Kennedy, 1995). The maximum variability puts a stringent requirement on the reference/working electrode, especially in the case of continuous monitoring. Additionally, when measuring against a reference electrode, factors such as pH, temperature, or ion concentration can drastically impact the reliability and stability of the electrode.

[0004] Accordingly, there is a need for electrodes which solve the above problems.

BRIEF SUMMARY

[0005] Devices and methods are provided for using a closed bipolar electrode (CBPE) system to detect and/or quantify the amount of an analyte.

[0006] An embodiment is a closed bipolar electrode system comprising:

[0007] a) a sample compartment;

[0008] b) a measurement compartment; and

[0009] c) a bipolar electrode having a first end and a second end,

[0010] wherein the first end of the bipolar electrode is connected to the sample compartment and the second end of the bipolar electrode is connected to the measurement compartment; and

[0011] wherein at least one of the first end and the second end, independently, have a biosensing molecule or a redox mediator immobilized thereto.

[0012] Another embodiment is a method of detecting or quantifying an analyte, comprising:

[0013] a) providing any closed bipolar electrode system described herein;

[0014] b) introducing an analyte into the sample compartment;

[0015] c) applying a sweeping potential to the system; and

[0016] d) measuring the change in open-circuit potential (OCP) at the bipolar electrode in the measurement compartment.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING(S) [0017] Having thus described the invention in general terms, reference will now be made to the accompanying drawings, which are not necessarily drawn to scale.

[0018] Figure 1 illustrates an exemplary scheme of closed biopolar electrode (CBPE) with sweeping voltage across the system. The first potentiostat is leveraged between working electrode, counter electrode, and reference electrode 1, which sweeps a voltage over the entire system. A second potentiostat is then used to monitor the OCP over the bipolar electrode which is electrodeposited with a reversible mediator.

[0019] Figures 2A-2E illustrate results of lactate-detection experiments. Figure 2A shows example measurements of OCP using working electrode 2, as the voltage is swept across the bipolar electrode. Figure 2B shows slope versus voltage with increasing lactate over the applied voltage. Figure 2C shows a calibration curve of current and lactate measured at -0.08V, the peak of the measured oxidation for arPES conjugated Aerococcus viridans derived lactate oxidase ( vLOx). Figure 2D shows current measured over Working electrode 1 (from Fig. 1) upon addition of lactate measured at -0.17V of cyclic voltametric sweep. Figure 2E shows the impact of increasing the sweeping scan rate over the bipolar electrode at 25mM lactate.

[0020] Figures 3A-3E illustrate additional data from lactate tests. Figure 3A shows the correlation and response of OCP over 2 triangle waves. Figure 3B shows calibration curves for both conventional AOCP scheme and bipolar AOCP scheme measured from l-25mM Alactate using an electrode with a diameter of 3mm. Figures 3C and 3D show calibration curves for bipolar AOCP scheme measured from l-25mM Alactate, using a 100 pm diameter circular electrode and a 25 pm diameter circular electrode, respectively. Figure 3E shows a comparison of limits of detection for both the bipolar OCP sensor and the conventional OCP sensor.

[0021] Figure 4 illustrates a calibration curve for a thrombin-sensitive electrode.

[0022] Figures 5A-5B illustrate the stability of the methylene blue (MB) modified aptamer complex of the thrombin-sensitive electrode. Figure 5A shows the stability after 30 minutes of incubation after measurement. Figure 5B shows the stability after several continuous cycles of sweeping voltage over the bipolar system.

[0023] Figure 6 illustrates a scheme of closed bi-polar electrode for detection of chlorine ions. The ratio of Prussian blue to Prussian white will be dictated by the concentration of [CT] bound with Ag(s) which can be detected us-ing the WE2 with open circuit potential (depicted OCP in figure).

[0024] Figure 7 illustrates a logarithmic fit for the first decade of [CT] concentration, giving a slope 25mV per decade of [CT] concentration.

[0025] Figure 8A shows microelectrode response for three electrodes having different surface areas, with [CT] shown with log axis. Figure 8B shows a comparison of AOCP versus A[CT] concentration. Each point represents n=3 replicates. Each electrode was compared against one another using an ANOVA showing no statistically significant difference among the 3 electrodes (p-value: 0.87). An ANOVA was also tested against concentration for each electrode showing statistical significance across each value (p-value = 0.00001). [0026] Figures 9A-9B illustrate a current response on WEi with [Cl’] addition, in log and linear scales, respectively. Specifically, the current is measured over WEi when KC1 is added to the measurement compartment. The response curve plateaus at a much lower concentration then the OCP measurement. This cunent is over the working electrode due to the shift in potential caused by KC1 addition. Upon shifting, current will flow to reach initial voltage and steady state.

[0027] Figure 10 illustrates the impact of shifting the applied voltage across WEI verse RE1 on a 10pm gold electrode.

[0028] Figure 11A demonstrates the stability of a bipolar [CT] electrode over 8 days. Averages and standard deviations were calculated by taking 10-minute period from 3 time points for each day. Figure 11B shows an example time curve for day 1 of the bipolar reference electrode.

[0029] Figure 12 illustrates activation of the nonspecific trans-nuclease catalytic activity of Casl2a.

[0030] Figure 13 illustrates an exemplary layout of the closed bipolar electrode (CBPE) sensor employing CRISPR-Casl2a as an amplification probe for target DNA detection. The left compartment will contain target DNA and Casl2a + gRNA. Upon the interaction of Casl2a and target DNA, the protein’s non-specific nuclease activity will be triggered, cleaving the MB-modified ssDNA immobilized to electrode 2. This will shift electrode 2’s junction potential, causing a change in the ratio of Prussian blue/Prussian white electrodeposited on electrode 4.

[0031] Figure 14 illustrates a schematic demonstrating the relationship between the two junction potentials. Altering one end of the CBPE electrode will facilitate changes on the other end of the electrode, independent of the electrode surface.

[0032] Figure 15 illustrates an exemplary layout of custom SPE CBPE electrode for DNA detection. Light green insulation separates the sample and measurement compartment, and the 2 Ag/AgCl electrodes (denoted by grey circle) hold a constant potential across the system.

DETAILED DESCRIPTION

[0033] The presently disclosed subject matter will now be described more fully hereinafter. However, many modifications and other embodiments of the presently disclosed subject matter set forth herein will come to mind to one skilled in the art to which the presently disclosed subject matter pertains having the benefit of the teachings presented in the foregoing descriptions. Therefore, it is to be understood that the presently disclosed subject matter is not to be limited to the specific embodiments disclosed and that modifications and other embodiments are intended to be included within the scope of the appended claims. In other words, the subject matter described herein covers all alternatives, modifications, and equivalents. In the event that one or more of the incorporated literature, patents, and similar materials differs from or contradicts this application, including but not limited to defined terms, term usage, described techniques, or the like, this application controls. Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary' skill in this field. All publications, patent applications, patents, and other references mentioned herein are incorporated by reference in their entirety.

[0034] The term “subject’' refers to a mammal (e.g., a human) in need of detection or quantification of an analyte, or target substance. The subject may include dogs, cats, pigs, cows, sheep, goats, horses, rats, mice, non-human mammals, and humans. The term “subject” does not necessarily exclude an individual that is healthy in all respects and does not have or show signs of elevated or lowered target substance. In embodiments, the sample is a biological sample.

[0035] Compositions or methods “comprising” or “including” one or more recited elements may include other elements not specifically recited. For example, a composition that “comprises” or “includes” a protein may contain the protein alone or in combination wi th other ingredients.

[0036] Designation of a range of values includes all integers, tenths, and hundredths wi thin or defining the range, and all subranges defined by integers within the range.

[0037] The term “zo vzvo” refers to natural environments (e.g., a cell or organism or body) and to processes or reactions that occur within a natural environment.

[0038] Unless otherwise apparent from the context, the term “about” encompasses values within a standard margin of error of measurement (e.g., SEM) of a stated value or variations ± 0.5%, 1%, 5%, or 10% from a specified value.

[0039] The singular forms of the articles “a,” “an,” and “the” include plural references unless the context clearly dictates otherwise.

[0040] The term “open circuit potential” refers to the potential established between the w orking electrode and the environment, with respect to a reference electrode.

[0041] The term “time dependent change” may refer to dOCP/dt or to dOCP/dVt. [0042] Statistically significant means p <0.05.

[0043] An embodiment is a closed bipolar electrode system comprising:

[0044] a) a sample compartment;

[0045] b) a measurement compartment; and

[0046] c) a bipolar electrode having a first end and a second end,

[0047] wherein the first end of the bipolar electrode is connected to the sample compartment and the second end of the bipolar electrode is connected to the measurement compartment; and

[0048] wherein at least one of the first end and the second end, independently, have a biosensing molecule or a redox mediator immobilized thereto.

[0049] In an embodiment, the first end has a biosensing molecule immobilized thereto, and the second end has a redox mediator immobilized thereto.

[0050] In an embodiment, the biosensing molecule is selected from the group consisting of redox probe-modified redox enzymes, direct electron transfer-type enzymes, a reversible oxidoreductase, a redox probe, an oligonucleotide, an aptamer, a binding protein, or a conjugation of two or more thereof.

[0051] In an embodiment, the redox probe-modified redox enzyme is a lactate oxidase; lactate dehydrogenase; glucose oxidase; glucose dehydrogenase; amino acid oxidase; glycerol-3-phosphate oxidase; glycerol phosphate-3-dehydrogenase; D-amino acid oxidase; fructosyl amino acid oxidase; fructosyl peptide oxidase; peroxidase; cholesterol oxidase; fructose dehydrogenase; beta-hydoxybutyrate dehydrogenase, alcohol oxidase, alcohol dehydrogenase, fructose dehydrogenase, diaphorase, glycerol dehydrogenase, an enzyme comprising flavin adenin dinucleotide (FAD) or flavin mononucleotide (FMN) as a cofactor; a fusion enzyme comprising FAD or FMN redox enzyme without a heme domain and/or without a heme subunit, and a heme protein; or an electron mediator-modified redox enzyme, with or without heme domain or heme subunit.

[0052] In an embodiment, the enzyme is a lactate oxidase; lactate dehydrogenase; glucose oxidase; glucose dehydrogenase; amino acid oxidase; glycerol-3 -phosphate oxidase; glycerol phosphate-3-dehydrogenase; D-amino acid oxidase; fructosyl amino acid oxidase; fructosyl peptide oxidase; peroxidase; cholesterol oxidase; fructose dehydrogenase; cellobiose dehydrogenase; uricase; alcohol oxidase; alcohol dehydrogenase; galactose oxidase; galactose dehydrogenase; pyranose oxidase; pyranose dehydrogenase; glucose-3-dehydrogenase; diaphorase; thyrosinase; 3-hydroxybutyrate dehydrogenase; amine oxidase; monoamine oxidase; polyamine oxidase; dopamine [3- monooxygenase; 4,5-DOPA dioxygenase extradiol; glutamate oxidase; sarcosine oxidase; an enzyme comprising flavin adenin dinucleotide (FAD) or flavin mononucleotide (FMN) as a cofactor; a fusion enzyme comprising FAD or FMN redox enzyme without a heme domain and/or without a heme subunit, and a heme protein; or an electron mediator- modified redox enzyme, with or without heme domain or heme subunit.

[0053] In an embodiment, the enzyme is selected from the group consisting of oxidases, dehydrogenases, monooxigenases and dioxygenases.

[0054] In an embodiment, the direct electron transfer-type enzyme is a glucose dehydrogenase, lactate dehydrogenase, fructose dehydrogenase, or cellobiose dehydrogenase. In an embodiment, the aptamer is a anti-thrombin aptamer, anti-insulin aptamer, anti-glucagon aptamer, anti-incretin aptamer, anti-bevacizumab aptamer, anti- cortizol aptamer, anti-C reactive protein (CRP) aptamer, anti-neuro peptide Y (NPY) aptamer, anti- Cry j 2 aptamer, anti-synuclein aptamer, anti thyroglobulin aptamer or anti- vascular endothelial growth factor (VEGF) aptamer.

[0055] In an embodiment, the biosensing molecule is an oxidoreductase. In an embodiment, the oxidoreductase can directly transfer electrons to the bipolar electrode surface without the need of an external cofactor. In an embodiment, the oxidoreductase is conjugated with a reversible electron mediator capable of transferring electrons to the electrode surface. In an embodiment, the oxidoreductase is immobilized using a thiol-based self-assembly monolayer onto a gold bipolar electrode surface. In an embodiment, the oxidoreductase is immobilized using vapor deposition of glutaraldehyde onto a gold bipolar electrode surface.

[0056] In an embodiment, the biosensing molecule is an oligonucleotide conjugated with a reversible electron mediator. In an embodiment, the oligonucleotide is an aptamer. In an embodiment, the aptamer is conjugated with amine-reactive phenazine ethosulfate or amine-reactive methylene blue, conjugated to a gold bipolar electrode surface. In an embodiment, the aptamer is immobilized using a thiol-modified oligonucleotide and/or thiol-based self-assembly monolayer onto a gold bipolar electrode surface.

[0057] In an embodiment, the biosensing molecule is a binding protein conjugated with a reversible electron mediator. In an embodiment, the reversible electron mediator is potassium ferricyanide, ferrocene, an osmium derivative, or phenazine methosulfate. In an embodiment, the binding protein is conjugated with amine-reactive phenazine ethosulfate or amine-reactive methylene blue conjugated to a gold bipolar electrode surface. In an embodiment, the binding protein is immobilized on the electrode using a thiol-based self- assembly monolayer onto a gold bipolar electrode surface. In embodiments, the binding protein is immobilized by cross-linking, encapsulating into a macromolecular matrix, coating with a dialysis membrane, optical cross-linking polymer, electroconductive polymer, oxidation-reduction polymer, and any combination thereof.

[0058] In an embodiment, a reversible mediator is immobilized to the second end of the bipolar electrode. In an embodiment, the reversible mediator has been immobilized via electrodeposition. In an embodiment, the reversible mediator has been immobilized via electrodeposition on gold or platinum. In an embodiment, the reversible mediator is selected from the group consisting of Prussian blue, thionine, osmium, or methylene blue. In an embodiment, the reversible mediator is in a hydrogel.

[0059] In an embodiment, the system further comprises a first potentiostat and a second potentiostat. In an embodiment, the first potentiostat comprises a first working electrode and a first reference electrode in the measurement compartment. In an embodiment, the second potentiostat comprises a second working electrode in the measurement compartment, a counter electrode in the sample compartment, and a second reference electrode in either the measurement compartment or the sample compartment.

[0060] In an embodiment, the bipolar electrode comprises gold, platinum, carbon, palladium, or any combination thereof, coated by an insulator material. In an embodiment, the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area and the second surface area are the same. In an embodiment, the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area is greater than the second surface area. In an embodiment, the bipolar electrode compnses a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area is less than the second surface area.

[0061] In an embodiment, the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the ratio of the first surface area to the second surface area ranges from about 1: 1000 to about 1000: 1. In an embodiment, the ratio is from about 1 :800 to about 800:1, about 1:400 to about 400:1, about 1 :200 to about 200: 1, about 1:100 to about 100:1, about 1 :50 to about 50: 1, about 1:20 to about 20:1, about 1: 10 to about 10:1, about 1:5 to about 5:1, or about 1:2 to about 2: 1. In an embodiment, the ratio is from about 1: 1000 to about 1:800, from about 1: 1000 to about 1 :400, from about 1 :1000 to about 1:200, from about 1 : 1000 to about 1 :100, from about 1: 1000 to about 1 :50, from about 1: 1000 to about 1 :20, from about 1: 1000 to about 1: 10; from about 1 : 1000 to about 1:5, from about 1: 1000 to about 1 :2, or from about 1: 1000 to about 1 :1. In an embodiment, the ratio is from about 1: 1 to about 1000: 1, from about 1: 1 to about 800:1, from about 1: 1 to about 400: 1, from about 1 :1 to about 200: 1, from about 1: 1 to about 100: 1, from about 1 : 1 to about 50: 1, from about 1: 1 to about 20: 1, from about 1: 1 to about 10: 1, from about 1 : 1 to about 5: 1, or from about 1: 1 to about 2: 1. In an embodiment, the ratio is about 1: 1000, about 1 :800, about 1 :400, about 1:200, about 1 :100, about 1 :50, about 1:20, about 1: 10, about 1 :5, about 1:2, about 1 :1, about 2: 1, about 5: 1, about 10: 1, about 20: 1, about 50: 1, about 100: 1, about 200: 1, about 400:1, about 800:1, or about 1000: 1.

[0062] In an embodiment, the first surface area is from about 10 pm² to about 7.14 nun². In an embodiment, the first surface area is from about 10 jam² to about 100 pm², from about 100 pm² to about 1000 pm², from about 1000 pm² to about 0.1 mm², from about 0.1 mm² to about 1.0 mm², or from about 0.1 mm² to about 7.14 mm². In an embodiment, the first surface area is from about 10 jam² to about 100 jam², from about 10 jam² to about 1000 pm², from about 10 jam² to about 0.1 nun², or from about 10 jam² to about 1.0 mm². In an embodiment, the first surface area is from about 100 jam² to about 7.14 nun², from about 1000 jam² to about 7.14 mm², from about 0.1 mm² to about 7.14 mm², or from about 1.0 mm² to about 7.14 mm².

[0063] In an embodiment, the second surface area is from about 10 pm² to about 7.14 mm². In an embodiment, the second siarface area is from about 10 pm² to about 100 jam², from about 100 pm² to about 1000 pm², from about 1000 pm² to about 0. 1 nun², from about 0.1 mm² to about 1.0 mm², or from about 0.1 mm² to about 7.14 mm². In an embodiment, the second surface area is from about 10 pm² to about 100 pm², from about 10 jam² to about 1000 pm², from about 10 pm² to about 0.1 mm², or from about 10 pm² to about 1.0 nun². In an embodiment, the second surface area is from about 100 jam² to about 7.14 mm², from about 1000 pm² to about 7.14 mm², from about 0.1 mm² to about 7.14 mm², or from about 1.0 mm² to about 7. 14 mm².

[0064] In an embodiment, the bipolar electrode is a single electrode. In an embodiment, the bipolar electrode is an array comprising two or more electrodes. In an embodiment, each of said two or more electrodes comprises an identical biosensing molecule immobilized thereto. In an embodiment, said two or more electrodes comprise more than one different biosensing molecule immobilized thereto. [0065] In an embodiment, the sample compartment comprises a buffer. In an embodiment, the measurement compartment comprises a buffer. In an embodiment, the buffer is a potassium phosphate buffer, a Tris (hydroxymethyl) aminomethane (THAM) hydrochloride (TRIS HC1) buffer, phosphate-buffered saline, HEPES, or MOPS. In an embodiment, the buffer is a potassium phosphate buffer. In an embodiment, the buffer concentration is from about 10 mM to about 1.0 M.

[0066] An embodiment is a wearable device comprising any closed bipolar electrode system described herein. Another embodiment is a test strip comprising any closed bipolar electrode system described herein.

[0067] An embodiment is a method of detecting or quantifying an analyte, comprising: [0068] a) providing any closed bipolar electrode system described herein;

[0069] b) introducing an analyte into the sample compartment;

[0070] c) applying a sweeping potential to the system; and

[0071] d) measuring the change in open-circuit potential (OCP) at the bipolar electrode in the measurement compartment.

[0072] In an embodiment, the change is measured using AOCP. In an embodiment, the change is measured using dOCP/dT. In an embodiment, the change is measured using dOCP/dVt.

[0073] In an embodiment, the sweeping potential is a square wave. In an embodiment, the sweeping potential is a cyclic potential.

[0074] In an embodiment, the scan rate of the potential ranges from about 10 mV/sec to about 200 mV/sec. In an embodiment, the scan rate ranges from about 20 mV/sec to about 150 mV/sec, from about 40 mV/sec to about 120 mV/sec, or from about 80 to about 100 mV/sec. In an embodiment, the scan rate ranges from about 20 mV/sec to about 150 mV/sec, from about 20 mV/sec to about 120 mV/sec, from about 20 mV/sec to about 100 mV/sec, from about 20 mV/sec to about 80 mV/sec, or from about 20 mV/sec to about 40 mV/sec, In an embodiment, the scan rate ranges from about 40 mV/sec to about 200 mV/sec, from about 80 mV/sec to about 200 mV/sec, from about 100 mV/sec to about 200 mV/sec, from about 120 mV/sec to about 200 mV/sec, or from about 150 mV/sec to about 200 mV/sec. As used herein, “scan rate” refers to the rate at which the applied potential changes.

[0075] In an embodiment, the potential ranges from about -1.0 V to about 1.0 V. In an embodiment, the potential ranges from about -0.8 V to about 0.8 V, from about -0.6 V to about 0.6 V, from about -0.4 V to about 0.4 V, or from about -0.2 V to about 0.2 V. [0076] In an embodiment, the potential is applied via a potentiostat.

[0077] In an embodiment, a single cycle takes from about 2 to about 200 seconds. In an embodiment, a single cycle takes from about 5 to about 150 seconds, from about 10 to about 120 seconds, from about 20 to about 80 seconds, or from about 40 to about 60 seconds. In an embodiment, a single cycle takes about 2 to about 10 seconds, about 10 to about 20 seconds, about 20 to about 30 seconds, about 30 to about 40 seconds, about 40 to about 50 seconds, about 50 to about 60 seconds, about 60 to about 70 seconds, about 70 to about 80 seconds, about 80 to about 90 seconds, about 90 to about 100 seconds, about 100 to about 110 seconds, about 110 to about 120 seconds, about 120 to about 130 seconds, about 130 to about 140 seconds, about 140 to about 150 seconds, about 150 to about 160 seconds, about 160 to about 170 seconds, about 170 to about 180 seconds, about 180 to about 190 seconds, or about 190 to about 200 seconds. As used herein, a “cycle” refers to the time for a complete forward and reverse scan.

[0078] In an embodiment, the analyte, or target substance, is selected from the group consisting of an enzyme, a sugar, a nucleic acid, an amino acid, an ion, a heavy metal, a bacteria, a virus, and a metabolite. In embodiments, the analyte, or target substance, is selected from the group consisting of D-serine, lactate, glucose, glycated proteins, glycated amino acid, hydrogen peroxide, cholesterol, glycerol, glycerol-3-phosphate, fructose, urate, ethanol, galactose, 1,5-anhydro-D-glucitol, NAD(P)H, dopamine, 3-hydroxy butyrate, Levodopa (L-DOPA), L-glutamate, L-glutamine, sarcosine, creatine, and creatinine. In an embodiment, the analyte is a bacteria or a virus. In an embodiment, the analyte is a nucleic acid from a bacteria or virus. In an embodiment, the bacteria is selected from the group consisting of Mycobacterium tuberculosis, Bordetella pertussis, Helicobacter pylori and Staphylococcus aureus. In an embodiment, the virus is selected from the group consisting of HIV, HPV, a coronavirus (such as SARS, SARS-CoV-2, orMERS), influenza viruses, varicella, a mumps virus, a herpes virus, a norovirus, a rotavirus, the Zika virus, the Dengu virus, and the Ebola virus.

[0079] In an embodiment, the nucleic acid is single stranded. In an embodiment, the nucleic acid is double stranded. In an embodiment, the nucleic acid can be detected or quantified without first being amplified. In an embodiment, the nucleic acid is an oligonucleotide.

[0080] DISCUSSION AND EXAMPLES

[0081] In recent years, there has been increasing interest in the creation of novel detection schemes to bring forth innovative solutions for many of the traditional challenges that inhibit biosensor development. A prominent example of this includes the realization of direct electron transfer glucose dehydrogenase, which eliminates the oxygen requirement for glucose sensing in addition to lowering the required overpotential, minimizing the impact of common interferents (Sode, 2017; Ferri, 2011; Horaguchi, 2014). Other examples include using new detection apparatuses/ analytical techniques such as open circuit potential (OOP) for measuring the change in an electrode’s junction potential. This technique has several benefits in that it lowers the power requirement and eliminates many issues caused by large overpotential application, which is often needed to oxidize/reduce an enzyme’s cofactor (Lee, 2018).

[0082] More recently, the use of bipolar electrochemistry has gained excitement for several benefits compared to traditional electrochemical detection schemes. Bipolar electrochemistry is the field of electrochemistry where a single electrode is held under a potential, driving anodic and cathodic reactions on either end. This method has been used to measure various ions, and even enzymes by transducing the potential change at one of the electrode to either an electrochemical luminescence change or colorimetric change on the other end (Jansod, 2020; Jansod, 2018; Fosdick, 2013; Wang, 2019; Karimian, 2019). This approach to detection offers several advantageous compared to traditional detection layouts. First is the ability of separating the sample compartment from the measurement compartment which mitigates common interference found when measuring the target analyte in a complex sample. In addition, the measurement cell’s buffer can then be controlled, independently of the solution’s characteristics. Further, since the signal transduction across the bipolar electrode is based on the Nemst equation, the signal is independent to electrode surface area allowing the signal to be transduced over a range of electrode sizes. This has allowed the system to measure various analytes ranging from glucose to calcium to DNA based on intercalation reaction with a complementary strand (Zhang, 2019; Wang, 2020).

[0083] A novel approach to using a closed bipolar electrode (CBPE) is presented, where instead of applying a constant potential as traditional methods propose, a sweeping voltage over the entire system is used. This enables the ability to both measure OCP changes, as well as measuring current at specific potentials where the electron mediator oxidizes, reducing the bipolar surface.

[0084] Figure 1 shows a non-limiting example of a CBPE system. Two potentiostats are employed, one to apply the sweeping potential across the system, while the second monitors OCP of the bipolar junction electrodeposited with Prussian blue. The sample compartment is where analyte is added, which will interact with the analyte sensitive bipolar electrode. This shift injunction potential will be measured across the conductive connection within the bipolar electrode using working and counter electrode 2. The electrodeposited Prussian blue on the bipolar electrode acts as the reference electrode 2 for working electrode 2 which measured the change in open circuit potential. Note the first potentiostat measures current based on the sweeping potential, and the second potentiostat records the changes in open circuit potential.

[0085] The function of this sensor can be mapped closely to the Nemst equation. To use the mechanism in Example 1, below, as a non-limiting example, upon interaction with lactate, lactate dehydrogenase is reduced, altering the potential of the enzyme and of the electrode’s junction. To balance this shift and maintain equilibrium, the ratio of Prussian blue to Prussian white will have to change by conservation of charge. This difference can be measured using OCP. One challenge when using OCP to monitor enzyme cofactor reduction is the need to “reset” the enzyme through application of an oxidative potential. Lee et al. demonstrated continuous monitoring of glucose using a novel algorithm: applying an oxidation potential, followed by OCP measurement (Lee, 2018). To mimic this effect, a sweeping potential is imposed over the system. When the junction is swept, the enzyme coated electrode will experience oxidative and reductive potentials, allowing the reset of the enzyme. This gives way to dual detection using OCP when the potential is outside of the oxidation range, and detection using current when the potential reaches the oxidation bias. By allowing the enzyme to re-oxidize, changes can be continuously measured in both OCP and current.

[0086] A key advantage of using a CBPE is that it affords separation of the measurement and sample compartments. This is very beneficial for avoiding common electrochemical interferents which may convolute the OCP signal upon interaction with either the reference or working electrode. Prussian Blue (PB), for instance, can be impacted by changes in pH or potassium ions (Haghighi, 2004; Ozeko, 1987).

[0087] In embodiments, the working electrode, system, or a compartment described herein is miniature, e. , less than about 100pm. In one embodiment, the working electrode, system, or compartment is less than about 2mm, less than about 1mm, or less than about 0.5mm in diameter. In embodiments, the working electrode, system, or compartment fits in a cell in vivo. In embodiments, the working electrode, system, or compartment is less than about 100pm, less than about 50pm, less than about 25 pm, less than about 10 pm, or less than about 1 pm in diameter. In embodiments, the working electrode, system, or compartment is about 1pm to about 100pm, about 10pm to about 100pm, about 25pm to about 100pm, about 50pm to about 100pm, about 75pm to about 100pm, about 1pm to about 75pm, about 1pm to about 50pm, about 1pm to about 25pm, or about 1pm to about 10pm in diameter.

[0088] The disclosed subject matter is further described in the following non-limiting Examples. It should be understood that these Examples, while indicating preferred embodiments of the invention, are given by way of illustration only.

[0089] Example 1 - Lactate biosensing

[0090] All chemicals were purchased from Sigma Aldrich (St. Louis, MO, USA) unless specified otherwise. Gold disk electrode (2mm and micro electrodes) were purchased from CH Instruments (Texas, USA).

[0091] Electrode preparation:

[0092] Gold disk electrode and micro electrodes were cleaned by polishing with alumina oxide pads of 0.3, and 0.05pm in subsequent order. Then each electrode was soaked in piranha acid for 20 minutes and thoroughly rinsed with MQ water. To ensure cleanliness, cyclic voltammetry was performed on each electrode in 50mM KOH until a consistent response curve was seen. Electrodes were rinsed in stored under MQ until use.

[0093] Prussian blue deposition:

[0094] Deposition of Prussian blue was carried out using ImL volume of 2.5mM ferric chloride (FeC13), 2.5mM ferricyanide, in a supporting electrolyte of 0. IM hydrochloric acid and 0.1M KC1 solution. Deposition was performed using a Biologic VMP3 Potentiostat (Biologic, France), with Ag/AgCl as the reference electrode, platinum wire as the counter electrode, and the gold disk electrode as the working electrode. To deposit the Prussian blue an applied bias of 0.4V was applied to the working electrode for 3 cycles of 4 minutes while monitoring the current. After deposition, the electrode was vigorously rinsed in MQ water and activated using cyclic voltammetry.

[0095] Prussian blue stabilization:

[0096] Cyclic voltammetry was run using 5 mL of 0.1M KC1 and 0.1M HC1, at a scan rate of 50mVsec-l, from -0.1V up to 0.5V. This was done for 25 cycles displaying a stable oxidation/reduction peak. Each electrode was rinsed with MQ and stored dry at 4°C until further use.

[0097] Prussian blue removal: [0098] The Prussian blue film was removed off the gold electrodes by soaking in 50mM KOH, then rinsed abundantly by MQ water. Electrodes were stored in MQ water until further use.

[0099] F abri cation of lactate sensitive electrode

[00100] The lactate sensitive electrode employed Aerococcus viridans LOx (AvLOx) modified with amine-reactive phenazine ethosulfate (PES) as previously described (Hiraka, 2020). Specifically, the electrode is configured with engineered LOx with Ala96Leu and Asn212Lys substitutions, which repressed its oxidase activity by keeping catalytic activity for the oxidation of L-lactate (reductive half reaction), with an additional Lys residue modified by PES. 16.1mg/mL enzyme was mixed with 2% mesoporous carbon, then drop casted on a 2mm gold disk electrode. After allowing the electrode to incubate until dry, polyethylene glycol) diglycidil ether (PEGDE) and polyethemmine (PEI) were combined at 200mg/mL and 0. Img/mL respectively and added to the electrode and allowed to incubate until dry. A schematic of the closed bipolar sensor system used in this experiment is depicted in Figure 1.

[00101] Electrochemical Testing Parameters

[00102] Performance for all electrochemical tests were done in lOOmM potassium phosphate buffer, using a platinum wire counter electrode and a single junction Ag/AgCl reference electrode unless specifically stated otherwise. All tests were performed using lOmL electrochemical cells and were covered with aluminum foil when needed to prevent Prussian Blue degradation caused by light. The sweeping potential across the CBPE ranged from 0.5 to -0.5V and has a scan rate of lOOmV/sec (unless specifically denoted otherwise). To compare the fidelity of the data, a 3 electrode OCP experiment was performed under the same conditions, except for having no sweeping potential and configured in a conventionial monopolar potentiometric scheme.

[00103] RESULTS

[00104] Current toward L-Lactate

[00105] In order to oxidize the arPES modified enzy me, between 0.0V - -0.2V must be applied (Hatada, 2018). To achieve this, the voltage was swept from -0.5V up to 0.5V using CV. Based on the scheme of the CBPE 3 different peaks would be expected, first between 0.2V and 0.4V due to the oxidation and reduction of Prussian blue/white, and then a second oxidation peak which correlates to the oxidation of arPES conjugated to AvLOx. With increasing concentration of lactate, there should be greater current at all three peaks due to having more reduced arPES. This phenomenon is shown by taking the slope (mA/mM) versus the sweeping voltage (V).

[00106] Figure 2A shows the change in OCP measured between the electrodeposited bipolar electrode and working electrode 2. The triangle waveform is due to the impact of sweeping the voltage across working electrode 1. The sensor is composed of 2 mm diameter gold electrode as a working electrode.

[00107] Figures 2B-2C show the relationship of current measured across working electrode 1, with respect to the changing lactate concentration. Using the same lactate sensitive electrode described in Figures 2A and 3B, below, it is clearly shown that current will flow⁷ across the bipolar electrode under certain voltage conditions. This allows the detection of lactate to occur simultaneously using OCP, AOCP, dOCP/dT as seen in Figure 3B, or with current as seen in Figures 2B-2C.

[00108] Figure 2B shows the slope (mA/mM) response over the entire voltage range swept (x-axis). The peak between -0.2V and -0.05V is due to oxidation of arPES conjugated to LOx, while the oxidation and reduction peaks between 0. IV - 0.4V are due to conversion of Prussian Blue/White.

[00109] Figure 2C is an example of a calibration curve between current and lactate concentrations measured when the sweeping voltage was at -0.08V, the oxidation peak found of the conjugated arPES and AvLOx. This is only one example of a selected voltage; other potentials could be used based on the mediator chosen to conjugate to the system. The measured slope was -0.13pA/mM lactate which was much lower than previous traditional amperometry detection schemes. At the addition of lactate, there is a clear increase in the baseline of measured OCP. This is due to the fraction of arPES conjugated AvLOx that is reduced upon binding with lactate. Figure 2D shows a similar calibration curve, but measured when the sweeping voltage was at -0.17V.

[00110] This decrease in measured current could be because the current measured is over working electrode 1 as shown in Figure 1. This electrode can only drive current through oxidation or reduction of water or oxygen, since there are no catalytic units attached. This may limit the ability to measure current over the bipolar scheme.

[00111] Figure 2E shows the impact of altering the scan rate (sweep rate) across the system using working electrode 1. The scan rate was changed from 20mV/sec up to lOOmV/sec using the lactate sensitive electrode described above, while lactate was held constant at 20mM concentration in the solution. As the scan rate slows, the change in OCP measured across the bipolar electrode becomes closer in magnitude to the applied sweeping potential. This allows the ability to tune or impact the range of measure OCP across the bipolar electrode based on the applied scan rate being swept over the entire system. As the scan rate increases, there is less change in magnitude of the OCP between the two peaks. This might be due to not allowing long enough time periods for lactate to interact with arPES-4 vLOx before the enzyme is re-oxidized.

[00112] OCP detection of L-Lactate

[00113] To measure the lactate using OCP, a similar approach was used as previously described for current. The slope of OCP versus lactate (mV/mM) over 1 entire CV sweep was plotted, as well as the correlation. This then gave a “peak” point where the slope was close to Nemstian for a 2 -valence electron transfer event (29.5mV/mM for Prussian Blue to Prussian White (Bakker; Zdrachek, 2019), and where the correlation gave a strong calibration curve. To validate using a bipolar scheme for OCP detection of the enzyme, these results were compared to a conventional OCP sensing scheme employing 2 electrodes. Both sets of arPES - 4 vLOx electrodes were developed in the same manner described above. The conventional potentiometric set up used an Ag/AgCl reference and counter electrode, while the bipolar scheme followed that described in Figure 1.

[00114] Figure 3A shows the slope of the OCP (mV/mM) and correlation over a 100 second time segment, showing 2 clear repetitive peaks at 20mV/ mM slope with a correlation of 0.9 RSQ. This subsequent voltage was then measured over each sweep, for each additional concentration of lactate. To help account for junction potential differences between the conventional OCP sensor, and the bipolar, the AOCP versus Alactate over the entire range (l-25mM) was analyzed.

[00115] Figure 3B shows the calibration curve for the average of 3 electrodes measured 3 times each. The slope and correlation were 36.4 mV/mM, and 0.98 for the bipolar electrode, and 21.5 mV/mM and 0.98 for the conventional electrode scheme. Figure 3B shows that the response measured in the bipolar scheme was like that of the conventional potentiometric design but had improved sensitivity at the lower end of lactate concentration. AOCP is one option which can be used to measure the lactate interaction with arPES modified double mutant LOx. Any other point along this triangle wave may also be used to develop a calibration curve, or to normalize the sensor to electroactive interference. Another option for measuring the lactate concentration is using dOCP/dT, as described in PCT International Patent Application No. PCT/US2021/062189).

[00116] Overlaid on Figure 3B is the detection of lactate using the same analyte sensitive electrode configuration, but in the conventional potentiometric scheme demonstrating the signal fidelity is preserved across the bipolar electrode. Figures 3C and 3D show a calibration curve for a bipolar AOCP scheme measured from l-25mM Alactate, using a 100 pm diameter electrode and a 25 pm diameter electrode, respectively.

[00117] Discussion

[00118] This Example aimed to measure a change injunction potential upon the reduction of arPES conjugated to A vLOx. This was demonstrated effectively using OCP in a closed bipolar scheme in addition to a conventional OCP sensor. To facilitate reversibility which would enable continuous detection, the enzyme must be periodically oxidized. This was achieved by sweeping a potential over the bipolar electrode from 0.5V to -0.5V. Limits of detection of 0.076 mM and 0.33 mM were calculated for both the sweeping bipolar OCP sensor and conventional OCP sensor, respectively. Use of a CBPE scheme improved the limit of detection by nearly 3 -fold when compared to the conventional scheme (Figure 3E). This improvement comes from the “resetting” of the enzyme, allowing OCP measurement to be relative to anew baseline.

[00119] This system has several benefits compared to previous potentiometric and bipolar enzyme-based sensors. Using arPES modified AvLOx facilitates direct detection of changes in lactate concentration as compared to previous methods, which measure changes indirectly using an enzymatic byproduct such as gluconic or lactic acid through electrochemical luminescence (Wang, 2019; EBmann, 2015; Xiao, 2017; Xu, 2017). These indirect approaches are limited in several important facets; first, tooling capable of measuring luminescence and skilled lab personnel are required for operation. Moreover, interference may compromise the signal when translated in vivo resulting from the impact of byproducts such as hydrogen peroxide (Hoffman, 2018; Joseph, 2018). Other methods rely on high applied potential biases, reaching nearly 9V across the system which can cause heat, or electroactive interference over the working electrode (Zhang, 2014). Potentiometric methods have a direct advantage compared to optical in being independent to surface area, making this an ideal platform form for miniaturization and translation (Smith, 2019; Liao, 2007). However, one challenge with conventional potentiometric sensors is reference electrode drift, or interference by competing ions such as Cl; Ag⁺, and OH' which may alter the signal measured (Guth 2009; Hu, 2016). Separation of the measurement and sample compartments drastically mitigates such interference because the reference electrode’s environment can be appropriately controlled (Walker, 2021).

[00120] Experiment 2 - Thrombin-specific electrode [00121] This experiment used the same scheme shown in as in Experiment 1, but with a thrombin-specific electrode. Methylene blue (arMB) was conjugated to a thrombin-specific aptamer. The MB aptamer was then immobilized onto a 3mm gold disk electrode using thiol-gold interaction. Upon the addition of thrombin and interaction with the MB conjugated aptamer, the junction potential shift; this shift was then measured across the bipolar electrode with electrodeposited Prussian blue.

[00122] Figure 4 is an example calibration curve for thrombin detection added to the sample compartment. The sample was allowed to sit for 30 minutes to ensure complete reaction occurs.

[00123] Figures 5A-5B show the stability of the MB aptamer complex over the bipolar electrode. Figure 5A shows the stability after 30 minutes of incubation after measurement. Figure 5B shows the stability after several continuous cycles of sweeping voltage over the bipolar system.

[00124] Experiment 3 - Chloride sensing

[00125] Ion selective electrodes using potentiometric detection are among the most common types of electrochemical sensors across academia and industry. Ion selective electrodes convert the change in activity of one or more ions into a measurable potential (Zdrachek, 2019; Zdrachek, 2021). When used in a potentiometric system, ion selective electrodes are calibrated against a known reference electrode and used in a state where negligible current flows between the 2 electrodes. The most commonly used reference electrode is Ag/AgCl, an ion selective electrode, due to its ease of fabrication, low cost, and simple miniaturization (Sophocleuous, 2017). This electrode may be simply classified as a pseudo (without supporting electrolyte such as 3M KC1) or single junction electrode where the Ag/AgCl wire is inserted into a supporting electrolyte such as KC1 and is partitioned from the environment using a porous glass frit (Bard, 2001; Sopstad, 2018). Ag/AgCl electrodes have a potential of 0.222V with respect to the standard hydrogen electrode, which is produced by inserting a platinum wire or mesh into a compartment that has either sealed or continuous flow of hydrogen gas. The standard hydrogen electrode (SHE) is set to a value of “0.00V” and is treated as a universal baseline to which all reference electrodes are measured (Hu, 2016; Jerkiewicz, 2020).

[00126] There are several limitations to using conventional ion selective electrodes in this type of scheme. The reference electrode is often exposed to the same medium as the sample which can promote error when measuring analytes in a complex sample matrix (interstitial fluid, whole blood, saliva etc.). Generally, this error manifests itself through interference of the working and reference electrodes by exogenous/endogenous competing ions and/or chemicals (Hu, 2016; Sophocleous, 2017. Each ion has a maximum allowed voltage variability that meets the percent error target can be calculated using the Nemst equation (Burtis, 1994; Kennedy, 1995). The maximum variability puts a stringent requirement on the reference/working electrode, especially in the case of continuous monitoring. Table 1 shows the error tolerance for various ions based on an iSTAT EC8+ from Abbott (Abbott Rapid Dx North America, LLC, Orlando, Florida). Additionally, when measuring against a reference electrode, factors such as pH, temperature, or ion concentration can drastically impact the reliability' and stability of the electrode.

[00127] Table 1 : Blood serum concentrations of 4 key ions (Calcium, Potassium, Sodium, and Chloride). All I-STAT measurable ranges and error came from i-STAT Test Cartridges | Abbott Point of Care procedures and guidelines (Microsoft Word - EC8+ IFU_Consolidated_2020 (zoetisus.com)).

00128] Many groups have worked to overcome these obstacles via implementation of new redox polymers, or conversion of the electrode into solid state capacitance based reference electrode (Hiraka, 2020). Another potentially attractive solution would be to separate the measurement cell from the sample cell, where the measurement cell could indirectly detect an electrochemical event occurring in the target compartment, and the two compartments could be controlled individually (Fosdick, 2013). This Example demonstrates that such a separation is achievable using bipolar electrochemistry, a concept that has attracted recent attention for use in reference electrode fabrication, as well as electrochemical to optical signal transduction (Walker, 2021; Jansod, 2018; Jansod, 2021). In this Example, the porous frit in commercially available reference electrodes was replaced with with an inert wire, setting up a closed bipolar system. Rather than using ion transfer to maintain electroneutrality, reactions are driven at each end of the wire to maintain communication between the working and reference electrodes.

[00129] The general schematic for the closed bipolar electrode is shown in Figure 6, where WEi, CEi, and REi represent the potentiostat holding a constant potential. A second potentiostat, represented by WE2, CE2, and RE2, is added to measure changes across a bipolar electrode coated in Prussian blue, a color-changing redox mediator (labeled PB in Figure 6). This setup allows us to detect OCP changes occurring strictly at the Prussian blue deposited electrode, independent of any ionic or molecular interferents in the sample compartment.

[00130] Several systems have been reported that monitor various electrolytes with the closed bipolar scheme using an optical array, in contrast to the OCP-based system described herein (Jansod, 2018; Jansod, 2021; Jansod, 2020). This study reports an electrochemical method to measure direct change of Prussian blue/white ratio using OCP and correlate the output to indirect changes in Cl' ion concentration. Such an electrochemical detection system has several benefits as compared to optical transduction. Electrochemical systems can be easily miniaturized compared to optical cells, which are often associated with higher capital costs for achieving quantitative detection (Chen, 2020). Also, OCP requires little or no power to drive the reaction, giving an advantage over other electrochemical methods such as amperometry.

[00131] EXPERIMENTAL

[00132] Materials:

[00133] All chemicals were purchased from Sigma Aldrich (St. Louis, MO, USA) unless specified otherwise. Gold disk electrode (2mm and micro electrodes) were purchased from CH Instruments (Texas, USA).

[00134] Electrode preparation:

[00135] Gold disk electrode and micro electrodes were cleaned by polishing with alumina oxide pads of 0.3, and 0.05pm in subsequent order. Then each electrode was soaked in piranha acid for 20 minutes and thoroughly rinsed with MQ water. To ensure cleanliness, cyclic voltammetry was performed on each electrode in 50mM KOH until a consistent response curve was seen. Electrodes were rinsed in stored under MQ until use.

[00136] Prussian blue deposition:

[00137] Deposition of Prussian blue was carried out using ImL volume of 2.5mM ferric chloride (FeCh), 2.5mM ferricyanide, in a supporting electrolyte of 0.1M hydrochloric acid and 0.1M KC1 solution. Deposition was performed using a Biologic VMP3 Potentiostat (Biologic, France), with Ag/AgCl as the reference electrode, platinum wire as the counter electrode, and the gold disk electrode as the working electrode. To deposit the Prussian blue an applied bias of 0.4V was applied to the working electrode for 3 cycles of 4 minutes while monitoring the current. After deposition, the electrode was vigorously rinsed in MQ water and activated using cyclic voltammetry. [00138] Prussian blue stabilization:

[00139] Cyclic voltammetry was run using 5 mL of 0. IM KC1 and 0. IM HC1, at a scan rate of 50mVsec , from -0.1V up to 0.5V. This was done for 25 cycles displaying a stable oxidation/reduction peak. Each electrode was rinsed with MQ and stored dry at 4°C until further use.

[00140] Prussian blue removal:

[00141] The Prussian blue film was removed off the gold electrodes by soaking in 50mM KOH, then rinsed abundantly by MQ water. Electrodes were stored in MQ water until further use.

[00142] Reference compartment:

[00143] The reference compartment was used to maintain a constant equilibrium between Ag/AgCl electrode of the bi-polar electrode. The was done using lOmL of lOmM potassium phosphate buffer. Potassium chloride was added serially to confirm the dependency of the Prussian blue oxidation/reduction state and the concentration of [Cl’]. In addition, a single junction Ag/AgCl reference electrode and platinum wire electrode were included to maintain the applied potential held between REi and WEi as shown in Figure 1.

[00144] Measurement compartment:

[00145] The measurement compartment contained the Prussian blue/white deposited bipolar electrode (PB/Ref2), the working electrode for the first potentiostat (WEi) which is used to hold the applied potential of Prussian blue across the bipolar electrode, a platinum counter electrode for the second potentiostat (CE2), and the working electrode for the second potentiostat (WE2). All experiments performed in the measurement compartment was under room temperature, in lOmM potassium phosphate buffer.

[00146] Chloride specific electrode fabrication:

[00147] For proof of concept, the chloride sensitive electrode was fabricated by taking a single junction Ag/AgCl electrode and removing the silver wire. This wire was then anodized using a positive potential of 0.4V held against a platinum counter electrode in IM KC1 solution. The electrode was stored in IM KC1 until further use.

[00148] RESULTS AND DISCUSSION

[00149] Prussian blue potential dependence on CP concentration

[00150] To confirm the Prussian blue/white’s potential dependency on chloride concentration, a concentration gradient of KC1 was added serially to lOmM potassium phosphate buffer solution. During this period OCP was monitoring across working electrode 2 (WE2). [Cl’] was added from a range of OmM up to 12mM this was tested using 3 separate 3mm diameter electrodes. Figure 7 shows the OCP response measured across Prussian blue electrode, and WE2 based on the concentration of [C1‘] added into the reference compartment. Error was measured using standard deviation of the 3 electrode trials. The calculated slope for this response was 25mV per decade, which is close to an ideal 29.5mV slope for a two-valence electron Nemstian response (Bakker; Chen, 2018).

[00151] Impact of electrode surface area

[00152] The [O’] concentration dependence was tested over several electrode diameters of 100pm, 25pm, and 10pm. Figure 8A shows each electrode fit to a logarithmic trendline to evaluate the response. Each electrode shown in Figure 8A was tested in the same manner as Figure 6, with a concentration range from OmM up to 55mM. The slope for each electrode was 22.5mV/decade [Cl’], 24.8mV/Decade Cl’, and 23.5mV/Decade [C1‘], respectively, as shown in Table 2, with an ideal Nemstian slope added for reference:

[00153] Table 2: Slope, and linear range of each electrode size against [C1-] tested using a closed bipolar scheme.

00154] All slopes are close to an ideal Nemst slope for a 2-valence electron system of 29.5mV/decade. Each electrode was tested for statistical significance against one another (based on surface area), and against changing chlonde concentration. Comparing electrode size to open circuit potential response gave no statistical significance using a one-way Analysis of Vanance (ANOVA (p-value = 0.875), verifying that surface area does not impact sensor response. A second one-way ANOVA was performed for all electrode sizes against concentration, resulting in statistical significance across the entire chloride concentration range (p-value = 0.00001). Each electrode has different initial OCP due to variations on the surface of the gold and different j unction potentials across the cell. Looking at the AOCP normalizes the impact of differences in electrode structure (Figure 8B). Both the upper and lower limit have variability as a result of the junction potentials, but across a linear range from ImM - 40mM the relative change in potential is equal across all electrode sizes. This, together with the data from Figure 8A, demonstrate that the system is capable of transducing reactions occurring on microelectrodes to larger surface areas, suggesting that closed bipolar electrodes may have promising application in an array format.

[00155] When KC1 is added to the measurement compartment there is a sudden spike and dissipation of current over WEi. This is caused by [Cl’] association with Ag⁺, which changes the potential across the Ag/AgCl electrode. To ensure conservation of charge, Prussian blue will shift the potential oppositely by dissociating with a cation into solution (Haghighi, 2010; Ozeki, 1987). This shifts the net potential away from the applied voltage across WEi and CEi, driving an anodic current which reverts the system to the initial state (see Figures 9 A and 9B).

[00156] Altering the Applied Bias

[00157] When using a bipolar electrode scheme, a bias is applied between WEi and REi that is maintained throughout the course of measurement. The selected bias should have a large impact on the measured OCP. To confirm this, this study used 10 pm gold disk electrodes, chose 5 different potentials, and measured chloride dependency (Figure 10). The applied potential varied from the calculated potential by ± 50 and ± lOOmV, causing a linear shift along the y-axis for the calibration curve. This ability to “tune” the potential detection range has been exploited optically using Prussian blue for the detection of various ions such as calcium, sodium, and potassium (Jansod, 2020). Additionally, this study tested the impact of alternating between reduction or oxidation potentials of Prussian blue. When each potential was applied, a sudden unexpected current flow across WEi and CEi was observed.

[00158] Prussian blue stability

[00159] The stability of the closed bipolar chloride electrode was evaluated by applying the FP of Prussian blue over the electrode (Figures 11A and 11B). The reference compartment was tested with saturated KC1 (approximately 4.56M) in lOmM potassium phosphate buffer, while the measurement compartment was held at lOmM potassium phosphate buffer. The electrode layout was identical to the scheme shown in figure 1, with two exceptions; first, the two compartments were covered with alumina foil to ensure that light would not interact with the Prussian blue layer (Gervais, 2014; Koncki, 2000). Then, a double junction Ag/AgCl electrode was used as WE2 instead of gold. The Ag/AgCl electrode has an inner filling solution of 4M KC1 and an outer filling solution of 10% potassium nitrate. Figure 11 A shows the average OCP measured at 3 different points during the day for 10 minutes each. The average OCP is lower in the first 24 hours than the following 7 days, which may result from the time required for Ag wire to reach equilibrium in saturated KC1 solution. In addition, after day 5, there is a steady decrease in OCP from about 186 mV down to 182 mV which could be due to the denaturation of Prussian blue (Haghighi, 2004; Navarro; Roig; Talagaeva, 2020). The total drift over the 8 days was found to be 13.8 pV/hr. which meets the potential drift requirement shown in table 1 for all 4 substrates. Using the Nemst equation, the expected OCP for saturated KC1 is 150mV.

[00160] Conclusion

[00161] This work demonstrated the ability to use a closed bipolar scheme to monitor [C1‘] changes across a wide range (l-55mM) while avoiding traditional interference that may im-pact the reference electrode. Microelectrodes were used to confirm signal transduction across differently sized electrodes in the same bipolar configuration, a property that can enable signal amplification and ease of interpretation. As a result, this technique facilitates size-independent microarray applications. Furthermore, OCP based detection eliminates the need for more complex optical tools and provides a platform for lower cost quantitative analysis.

[00162] Example 4 - Disease detection

[00163] BACKGROUND AND SIGNIFICANCE

[00164] Current limitations for the detection of infectious disease

[00165] Development of sensors for infectious disease is paramount for their early diagnosis, containment, and treatment across the world. There has been recent effort by the World Health Organization (WHO) (Kosack; Land, 2019) to standardize acceptable criteria for such types of sensors, which corresponds to real-time, affordable, sensitive, specific, user friendly, robust, rapid, equipment free, and deliverable, or “REASSURED.” However, current standard of care diagnostic platforms for infectious disease, represented by the detection of tuberculosis and multidrug-resistant tuberculosis (TB, MDRTB), rely on expensive capital equipment, chemical post-processing of samples, and/or skilled technicians to run and interpret data. For TB, effective diagnosis is particularly limited by a difficult balance between sensitivity/specificity and cost/rapidity.

[00166] Detection of TB and MDRTB currently falls short of the ideal REASSURED. In low-income developing countries, many clinics use a tuberculin skin test for TB that is inexpensive but has low sensitivity, specificity, and fails to give any direction to physicians regarding the risk of MDRTB (Diel, 2009; Gualano, 2019). Extraction of patient sputum, blood, or urine allows enables more sensitive and rapid detection of TB and even differentiation of rifampicin and isoniazid resistance via molecular analysis of protein and nucleic acid biomarkers (Thomson, 2005; Delacourt, 1993; Gamboa, 1997; Goletti, 2016). However, such type of analysis comes with several disadvantages from a point-of-care (POC) perspective. Clinically relevant concentrations of TB and MDRTB bacteria in the sputum might be as low as a few pg/mL, with single point mutations enough to differentiate between the wild-type and drug resistant strands (Song, 2018; Chakravorty, 2017; Huang, 2009). Therefore, current analyses of DNA extracted from samples typically require an amplification step such as polymerase chain reaction (PCR). PCR adds complexity, requires capital equipment, and if large amplification is needed, can negatively impact the specificity of detection. More recently, the use of Loop-mediated isothermal amplification (LAMP) has facilitated faster, less expensive, lower volume amplification by eliminating the need for temperature cycling. Once amplified, DNA recognition can be transduced into a measurable signal using methods such as fluorescence resonance energy transfer (FRET), a highly sensitive optical technique. While LAMP has substantially reduced capital costs as compared with PCR, it still requires skilled technicians and central laboratories for successful implementation (Gupta, 2018). Thus, developing platforms which do not require DNA amplification remains an attractive option for POC TB detection.

[00167] TB sensors that do not require the amplification of DNA have their own set of associated challenges

[00168] Amplification-free, electrochemical DNA sensors have previously been reported using techniques such as electrochemical impedance spectroscopy (EIS), which measures a change in impedance over the surface of an electrode upon binding of DNA and its recognition element (Teengam, 2018). Another approach involves differential pulse voltammetry (DPV), a technique whereby target DNA is labeled with a redox probe, such as methylene blue (MB), and immobilized on an electrode. DPV can be used to sensitively detect changes in peak current induced by binding of the labelled DNA and correlate that value to the concentration of the target (Eloi, 2020). However, these techniques suffer from several drawbacks; both EIS and DPV are very susceptible to noise and requires a sophisticated measurement apparatus. As an alternative to direct DNA detection, there are several reported methods to measure TB and MDRTB specific biomarkers in a POC format. Most of these methods involve molecular recognition elements that bind specifically to molecules such as lipoarabinomannan (LAM), antigen 85 complex (Ag85), ESAT-6, CFP- 10, MPT64, and 38 kDa antigen. Electrochemical and optical approaches have been developed for transducing these binding events, depending on the biomarker of interest. Two POC examples were reported including lab-on-chip ELISA-based assays and lab-on- chip amperometric assays (Evans, 2017). However, like the DNA-based assays, these techniques necessitate skilled technicians, and/or the use of tools only available in a central laboratory, which limits their widespread use.

[00169] CRISPR/Cas systems provide an innovative principle for DNA sensing

[00170] The discovery of Clustered Regularly Interspaced Palindromic Repeats (CRISPR)/CRISPR-associated system (Cas) and their development as molecular tools has led to a revolution in gene editing research. A fast-evolving area of research focused on the creation of CRISPR/Cas-based diagnostic tools builds on the specificity, programmability, and ease of use of the technology. Such systems may have promising application in the creation of POC diagnostic tests for detection of nucleic acids in routine clinical care. One notable example is the “DNA Endonuclease Targeted CRISPR Trans Reporter (DETECTR)” technology, first reported by Chen et al., which leverages the Casl2a protein in complex with a guide RNA (gRNA) for specific DNA detection. gRNA can be programmed to bind with virtually any complimentary strand of target DNA containing a proto-spacer adjacent motif (PAM) upstream of the target sequence. Upon binding of the DNA, a conformational change is induced in the protein’s structure that activates its RuvC domain to indiscriminately cleave the phosphodiester backbone of any single-stranded DNA (ssDNA) present in the system (Fig. 12). The non-specific trans-nuclease activity of this domain enables a single DNA binding event to be amplified thousands of times over. In this example, the detection system utilizes FRET, where the ssDNA is labelled with a fluorophore and quencher. Upon cleavage of the ssDNA, the proximity of the fluor ophore/quencher is altered and fluorescence increases. Systems have been designed capable of detecting downwards of attomolar concentrations post-amplification of target DNA, exhibiting specificity of a single point mutation (Xu, 2020; Chen, 2018; Dai 2019; Ibrahim, 2020). However, when comparing CRISPR/Cas based systems against the REASSURED criteria, several apparent drawbacks arise. First, most of the described methods still rely on traditional DNA amplification techniques to increase the detection range (Li, 2019; Aman, 2020). Specific high-sensitivity enzymatic reporter unlocking (SHERLOCK), One-hour Low-cost Multipurpose highly Efficient System (HOLMES), and DETECTR are examples of successful Casl2/13 fluorescent biosensing systems that rely on DNA amplification methods. In addition, performing a FRET assay requires skilled technicians and expensive capital equipment to measure the fluorescent signal output. Several types of CRISPR/Cas based electrochemical biosensors have been realized that do not require any kind of DNA amplification, using techniques such as EIS and DPV (Xu, 2020; Suea-Ngam, 2021; Sheng, 2021). These sensors have achieved limits of detection down to picomolar concentrations of disease-relevant nucleic acid biomarkers. However, like the DNA-based assays, these techniques necessitate skilled technicians, and/or the use of complicated electrochemical tools only available in a central laboratory, which limits their widespread use. A particularly innovative approach was developed by Dai et al.; nonspecific, single-stranded DNA modified with MB was covalently immobilized onto the surface of a gold electrode. Upon interaction of the target DNA with CRISPR-Casl2a, the immobilized ssDNA was cleaved, releasing the redox probe into solution and reducing the current derived from MB-mediated electron transfer (Dai, 2019). Although this system functions with high accuracy and specificity, the required incubation time for detection was over 30 minutes, and the technique involved a complicated voltametric measurement procedure, limiting the system’s suitability for rapid POC detection.

[00171] Current bipolar electrochemical sensors for various target analytes

[00172] Bipolar electrochemistry is a sub-field of electrochemistry where a single electrode is held under a constant bias potential and undergoes both oxidation and reduction reactions on either end. This electrochemical approach has been used recently for the measurement of various analytes, such as ions, heavy metals, metabolites, and even DNA (Fosdick, 2013; Jansod, 2020; Liu, 2016). In most of these detection systems, an interaction between the junction electrode and some charged analytes, causes a shift in the junctional potential. For instance, Bakker et al. demonstrated the electrooptical, colorimetric detection of various ions by using calcium, sodium, and potassium sensitive electrodes at one end of the bipolar electrode and electrodeposited Prussian blue on the other end (Jansod, 2018). Zhang et al., used a bipolar scheme to create an electrochemiluminescence sensor which was capable of electrooptically detecting glucose and DNAWang, 2019; Zhang, 2021). These systems have demonstrated the flexibility and efficacy of bipolar electrochemistry as a measurement tool for different analytes. Yet, to measure DNA with sufficient sensitivity, PCR was still required, introducing complexity, cost, and additional time to the tests.

[00173] There remains an unmet need for a simple, low-power, amplification-free, TB/MDRTB detection platform for use in environments where access to complicated equipment and experienced personnel may not be readily available. CRISPR/Cas and bipolar electrochemical sensing are two areas of developing research that show potential to push the current frontier of nucleic acid detection. A combination of these evolving tools presents an attractive strategy to address the limitations of current POC DNA sensors for infectious disease, using TB/MDRTB as a case study. [00174] To overcome the aforementioned limitations, to allow for direct detection of nucleic acids from biologically derived sample matrices, and to amplify the signal without the use of PCR or LAMP, or sophisticated measurement apparatus, the CRISPR/Casl2a system in tandem with a CBPE as described hereinabove is used. This combination makes possible the rapid and sensitive detection of DNA sourced from Mycobacterium tuberculosis as a case study, using low-power electronics and adhering to the criteria set forth in the REASSURED framework.

[00175] Signal amplification through the nonspecific trans-nuclease catalytic activity of CRISPR/Casl2a and the alteration of CBPE junction potential

[00176] The potential gradient applied across a CBPE drives anodic reactions on one end, and cathodic reactions on the other (Fosdick, 2013; Gamero-Quijano, 2019). To alter one end of the bipolar electrode’s junction potential, nonspecific, MB-modified ssDNA is immobilized onto the gold CBPE surface using thiol-gold interaction. As described in Dai et al., upon treating the sample compartment with CRISPR/Casl2a and its specific target double-stranded DNA (dsDNA) sequence, the RuvC domain of the protein is activated to non-specifically cleave the ssDNA and release the MB redox probe, which alters the charge of the electrodes surface from steady state (junction “2” in Figure 13). The difference is that, in the system described hereinbelow, the opposite end of the CBPE electrode in the measurement compartment has to oppositely shift in potential to balance (by conservation of charge). This phenomenon can be transduced electrooptically to create a colorimetric change using electrodeposited Prussian blue, a redox active molecule (junction “4” in Figure 13). When CRISPR-Casl2a cuts the MB-modified ssDNA, the junction will have an increase in voltage caused by removal of the negatively charged MB mediator. To keep the junction difference conserved, the opposite end of the bipolar electrode must uptake electrons, converting Prussian blue to Prussian white and changing the color from blue to transparent. This color shift can be visualized with the naked eye without the need of any skilled training or sophisticated measurement apparatus.

[00177] Electrooptical signal transduction is independent to surface area, providing an opportunity for further signal amplification

[00178] One considerable benefit of using a CBPE is the capability to transduce a signal from anode to cathode independent of the electrode surface area, as the technique is based upon potentiometric detection. The magnitude of signal measured over the ssDNA immobilized or Prussian blue deposited electrodes is the same for a 3mm² electrode as for a 10pm² electrode. This allows for substantial amplification of small electrochemical changes (surface removal of MB mediator) into a much larger colorimetric change. This property of the system improves its sensing capability in two key ways. First, by combining a microelectrode (~10 pm²) immobilized with MB-modified ssDNA with a large macroelectrode (~2 mm²) electrodeposited with Prussian blue, very few ssDNA need to be cleaved off the gold electrode to transduce a large relative change (Figure 14). The mathematics behind this phenomenon are detailed in the Nemst equation, where signal magnitude is based directly on the ratio of oxidized to reduced substrates (Equation 1). This implies that if the electrode surface area shrinks yet the ratio of oxidized to reduced substrates stays the same, signal transduction occurs at the same magnitude for both surface areas. Another advantage further in development is that the use of microelectrodes ultimately enable more facile fabrication of low-cost arrays and large-scale detection systems. Moreover, as the electrode surface area shrinks, there are many fewer total molecules of MB-modified ssDNA immobilized in the system. It is believed that that this will increase the sensitivity of the system, as a smaller number of ssDNA molecules need to be cleaved by Casl2a in order to appreciably alter the ratio of oxidized and reduced molecules. This property may also drastically reduce the time for signal transduction over the bipolar electrode to the Prussian blue electrodeposited electrode, facilitating more rapid detection.

[00179]

RT OX

Equation 1: E_signal = —Ln(—')

^M zF Red

[00180] In summary, this project is innovative for the following four reasons. The proposal realizes (1) thermodynamically driven electrooptical transduction across two separate compartments using a low-power source and minimal hardware, (2) detection of nucleic acids derived from TB/MDRTB via a CBPE with microelectrodes facilitating very rapid feedback of data, (3) Novel application of CRISPR-Casl2a which self amplifies the signal allowing for very high sensitivity and specificity of target DNA detection. (4) The ability to translate electrooptical signal independent of electrode surface area, which will facilitate a low-cost, visual readout.

[00181] DETAILS

[00182] In this Example, an innovative, low-power, and low-cost nucleic acid biosensing platform utilizing bipolar electrochemistry and CRISPR-Cas 12a for the detection of nucleic acid biomarkers relevant in infectious disease is described, specifically targeting dsDNA derived from TB/MDRTB as a test case. The biochemical rationale behind the platform is to leverage the non-specific nuclease activity of CRISPR-Casl2a to cut ssDNA in the presence of target dsDNA specific to TB/MDRTB. Then, the nuclease activity is transduced using a CBPE, with one end immobilized with MB-modified ssDNA, and the other electrodeposited with Prussian blue. CRISPR-Casl2a cleaves the ssDNA, dissociating the MB and causing Prussian blue to convert to Prussian white by conservation of charge, which is interpreted visually by the user. Finally, a disposable, miniaturized electrooptical sensor array is developed and evaluated using dsDNA contained within artificial sputum samples.

[00183] Establishment of bipolar electrooptical platform for amplification-free DNA detection

[00184] To establish rapid and size-independent measurement across the CBPE, sample and measurement compartments are constructed and evaluated separately, then integrated using individually optimized conditions to realize an electrooptical platform for amplification-free DNA detection.

[00185] Development and testing of the sample compartment for the CBPE sensor

[00186] The two most common drug resistances that constitute MDRTB are Rifampin and Isoniazid resistance, respectively. Rifampin resistance is derived primarily from select mutations in a conserved sequence of 81 base pairs (residues 436-463) that exist in the rpo0 gene of Mycobacterium tuberculosis. Isoniazid resistance is not as simply conferred; mutations in the katG, inhA, and ahpC genes can contribute to its development. In this study, Casl2a protein derived from Lachnospiraceae bacterium (LbCasl2a) is used, with gRNA sequences complementary to the regions of interest on the rpo/3 and katG genes, which both contain proto-spacer adjacent motifs (TTCV) upstream of the target sequence, as a proof of concept (Chen, 2020; Miller, 1994; Zhang, 2019). All of these biologies are purchased commercially. Conditions for target recognition are optimized using a microplate FRET assay similar to the DETECTR platform, utilizing the nuclease activity of LbCasl2a to separate fluorophore and quencher. The specificity of gRNA and target DNA binding are evaluated using this system to determine if LbCasl2a can differentiate point mutants. Key variables for evaluation include 1) Buffer components. 2) gRNA to LbCasl2a ratio. 3) Specificity of LbCasl2a. (differentiating a 1-2 base-pair mismatch). 4) Determining the maximal sensitivity of LbCasl2a with the FRET-based assay.

[00187] Further, the evaluation of the second junction electrode is performed starting with the immobilization of ssDNA on the surface of gold disk electrodes with varying surface areas, ranging from 2mm² and 10pm². The DNA sample is added to the compartment together with the Casl2a protein + gRNA under the optimum reaction conditions as determined via FRET. Casl2a activity is monitored using the differently sized electrodes and the system are evaluated for response time, magnitude, and stability. Open circuit potential (OCP) measurements are taken across a bipolar electrode against a standard Ag/AgCl reference electrode as a quantitative metric. From these experiments, the relationship between electrode surface area and sensor capability (response time, magnitude, and stability) is ascertained.

[00188] Development and testing of the measurement compartment for the CBPE sensor

[00189] Development of the bipolar measurement compartment involves electrodeposition of Prussian blue, which exhibits colorimetric change upon junction potential shift in the sample compartment. Evaluation of the benchtop system is broken down into 3 key performance metrics: 1) Sensor sensitivity 2) Response time for electrooptical readout, and 3) limit of miniaturization for transducing electrooptical signals across the CBPE.

[00190] Development and testing of integrated sample and measurement compartment for CBPE sensor

[00191] Once each compartment has been developed independently, integration and system testing is done to combine them (Figure 13). The specificity of this sensor design is derived from two components: 1) the ability of Cast 2a + gRNA to bind its target sequence and not non-specific dsDNA, and 2) the ability of the electrode itself to spontaneously interact with electroactive materials such as acetaminophen or ascorbic acid. This is evaluated by measuring the change in absorbance as well as OCP of the Prussian blue electrode. Limit of detection is evaluated by serially decreasing the amount of target dsDNA added to the sample compartment. The operational protocol for CBPE detection of TB/MDRTB is based on the visual change of Prussian blue to Prussian white. The system is configured with the parameters determined above and is used to develop a visual calibration curve correlating colorimetric change to concentration of target TB/MDRTB dsDNA.

[00192] Development of miniaturized and disposable electrooptical sensor array for rapid detection of TB and MDRTB

[00193] To prepare the bipolar electrooptical platform for amplification-free DNA detection in a POC format, a disposable TB/MDRTB sensor array is constructed by miniaturizing and integrating the components discussed above. The system is evaluated using artificial sputum samples containing nucleic acid sequences derived from TB/MDRTB without DNA amplification.

[00194] Design and fabrication of screen-printed CBPE array capable of multianalyte nucleic acid detection

[00195] The test system outlined above is miniaturized to a disposable sensor array using custom screen-printed CBPEs. Such sensor chips are available commercially (Metrohm Dropsens, 2021) for the purposes of translating benchtop developments to a more POC- accessible format. The electrodes are built as described. A polyimide substrate is used as the base, followed by deposition of thin film gold or indium tin oxide (ITO) thin film. Conductive traces are deposited between the two electrodes. Following this, the conductor is insulated with polyimide, and Ag/AgCl paste is deposited for maintaining the applied potential. Before use, Prussian blue is electrodeposited onto the thin film ITO, and the MB- modified ssDNA is immobilized onto the gold electrode (Figure 15). The sample is applied to the gold screen-printed electrode (sample compartment) together with CRISPR-Casl2a in the optimized solution composition. The second buffer is added to the measurement compartment for Prussian blue-mediated optical readout. To extrapolate this into an array format, screen-printed electrodes are tested in parallel using a single potential across all the electrodes.

[00196] Characterization and validation of disposable electrooptical sensor array using artificial sputum sample

[00197] To validate the overall system, artificial sputum loaded with various concentrations of dsDNA specific to wild-type TB and the mutations associated with MDRTB is used. The expected sensitivity and specificity is above 90% to meet the average response of the Xpert MTB/RIF assay, which is commonly used as a gold standard laboratory test for TB/MDRTB (Deng, 2019; Solanki, 2020; Sin, 2014). The target response time is less than 5 minutes for full operation to allow for immediate feedback.

[00198] Many modifications and other embodiments of the inventions set forth herein will come to mind to one skilled in the art to which the inventions pertain having the benefit of the teachings presented in the foregoing descriptions and the associated drawings. Therefore, it is to be understood that the inventions are not to be limited to the specific embodiments disclosed and that modifications and other embodiments are intended to be included within the scope of the appended claims.

[00199] Although specific terms are employed herein, they are used in a generic and descriptive sense only and not for purposes of limitation. Each embodiment disclosed herein is contemplated as being applicable to each of the other disclosed embodiments. All combinations and sub-combinations of the various elements described herein are within the scope of the embodiments.

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Claims

WHAT IS CLAIMED IS:

1. A closed bipolar electrode system comprising: a) a sample compartment; b) a measurement compartment; and c) a bipolar electrode having a first end and a second end, wherein the first end of the bipolar electrode is connected to the sample compartment and the second end of the bipolar electrode is connected to the measurement compartment; and wherein at least one of the first end and the second end, independently, have a biosensing molecule or a redox mediator immobilized thereto.

2. The closed bipolar electrode system of claim 1, wherein the first end has a biosensing molecule immobilized thereto, and the second end has a redox mediator immobilized thereto.

3. The closed bipolar electrode system of claim 1 or 2, wherein the biosensing molecule is selected from the group consisting of redox probe-modified redox enzymes, direct electron transfer-type enzymes, a reversible oxidoreductase, a redox probe, a nucleotide, an aptamer, a binding protein, or a conjugation of two or more thereof.

4. The closed bipolar electrode system of claim 3, wherein the redox probe- modified redox enzyme is a lactate oxidase; lactate dehydrogenase; glucose oxidase; glucose dehydrogenase; amino acid oxidase; glycerol-3-phosphate oxidase; glycerol phosphate-3 -dehydrogenase; D-amino acid oxidase; fructosyl amino acid oxidase; fructosyl peptide oxidase; peroxidase; cholesterol oxidase; fructose dehydrogenase; beta- hydoxybutyrate dehydrogenase, alcohol oxidase, alcohol dehydrogenase, fructose dehydrogenase, diaphorase, glycerol dehydrogenase, an enzyme comprising flavin adenin dinucleotide (FAD) or flavin mononucleotide (FMN) as a cofactor; a fusion enzyme comprising FAD or FMN redox enzyme without a heme domain or heme subunit, and a heme protein; or an electron mediator-modified redox enzyme, with or without heme domain or heme subunit.

5. The closed bipolar electrode system of claim 3, wherein the direct electron transfer-type enzyme is a glucose dehydrogenase, lactate dehydrogenase, fructose dehydrogenase, or cellobiose dehydrogenase.

6. The closed bipolar electrode system of claim 3, wherein the aptamer is an antithrombin aptamer, anti-insulin aptamer, anti-glucagon aptamer, anti-incretin aptamer, anti- bevacizumab aptamer, anti-cortizol aptamer, anti-C reactive protein (CRP) aptamer, anti- neuro peptide Y (NPY) aptamer, anti- Cry j 2 aptamer, anti-synuclein aptamer, anti thyroglobulin aptamer or anti- vascular endothelial growth factor (VEGF) aptamer.

7. The closed bipolar electrode system of claim 1 or 2, wherein the biosensing molecule is an oxidoreductase.

8. The closed bipolar electrode system of claim 7, wherein the oxidoreductase can directly transfer electrons to the bipolar electrode surface without the need of an external cofactor.

9. The closed bipolar electrode system of claim 7, wherein the oxidoreductase is conjugated with a reversible electron mediator capable of transferring electrons to the electrode surface.

10. The closed bipolar electrode system of claim 7, wherein the oxidoreductase is immobilized using a thiol-based self-assembly monolay er onto a gold bipolar electrode surface.

11. The closed bipolar electrode system of claim 7, wherein the oxidoreductase is immobilized using vapor deposition of glutaraldehyde onto a gold bipolar electrode surface.

12. The closed bipolar electrode system of claim 1 or 2, wherein the biosensing molecule is an oligonucleotide conjugated with a reversible electron mediator.

13. The closed bipolar electrode system of claim 12, wherein the oligonucleotide is an aptamer.

14. The closed bipolar electrode system of claim 12, wherein the nucleotide is conjugated with amine-reactive phenazine ethosulfate or amine-reactive methylene blue, conjugated to a gold bipolar electrode surface.

15. The closed bipolar electrode system of claim 12, wherein the nucleotide is immobilized using a thiol-modified oligonucleotide and/or thiol-based self-assembly monolayer onto a gold bipolar electrode surface.

16. The closed bipolar electrode system of claim 1 or 2, wherein the biosensing molecule is a binding protein conjugated with a reversible electron mediator.

17. The closed bipolar electrode system of claim 16, wherein the binding protein is conjugated with amine-reactive phenazine ethosulfate or amine-reactive methylene blue, conjugated to a gold bipolar electrode surface.

18. The closed bipolar electrode system of claim 12, wherein the binding protein is immobilized using a thiol-based self-assembly monolay er onto a gold bipolar electrode surface.

19. The closed bipolar electrode system of any one of claims 1-18, wherein a reversible mediator is immobilized to the second end of the bipolar electrode.

20. The closed bipolar electrode system of claim 19, wherein the reversible mediator has been immobilized via electrodeposition.

21. The closed bipolar electrode system of claim 19, wherein the reversible mediator has been immobilized via electrodeposition on gold or platinum.

22. The closed bipolar electrode system of any one of claims 19-21, wherein the reversible mediator is selected from the group consisting of Prussian blue, thionine, or methylene blue.

23. The closed bipolar electrode system of any one of claims 1-22, further comprising a first potentiostat and a second potentiostat.

24. The closed bipolar electrode system of claim 23, wherein the first potentiostat comprises a first working electrode and a first reference electrode in the measurement compartment.

25. The closed bipolar electrode system of claim 23 or 24, wherein the second potentiostat comprises a second working electrode in the measurement compartment, a counter electrode in the sample compartment, and a second reference electrode in either the measurement compartment or the sample compartment.

26. The closed bipolar electrode system of any one of claims 1-25, wherein the bipolar electrode comprises gold, platinum, carbon, palladium, or any combination thereof, coated by an insulator material.

27. The closed bipolar electrode system of any one of claims 1-26, wherein the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area and the second surface area are the same.

28. The closed bipolar electrode system of any one of claims 1-26, wherein the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area is greater than the second surface area.

29. The closed bipolar electrode system of any one of claims 1-26, wherein the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the first surface area is less than the second surface area.

30. The closed bipolar electrode system of any one of claims 1-26, wherein the bipolar electrode comprises a first surface area in the sample compartment and a second surface area in the measurement compartment, wherein the ratio of the first surface area to the second surface area ranges from about 1 : 1000 to about 1000: 1.

31. The closed bipolar electrode system of any one of claims 27-30, wherein the first surface area is from about 10 pm² to about 7.14 mm².

32. The closed bipolar electrode system of any one of claims 27-31, wherein the second surface area is from about 10 pm² to about 7. 14 mm².

33. The closed bipolar electrode system of any one of claims 1-32, wherein the bipolar electrode is a single electrode.

34. The closed bipolar electrode system of any one of claims 1-32, wherein the bipolar electrode is an array comprising two or more electrodes.

35. The closed bipolar electrode system of claim 34, wherein each of said two or more electrodes comprises an identical biosensing molecule immobilized thereto.

36. The closed bipolar bipolar electrode system of claim 34, wherein said two or more electrodes comprise more than one different biosensing molecule immobilized thereto.

37. A wearable device comprising the closed bipolar electrode system of any one of claims 1-36.

38. A test strip comprising the closed bipolar electrode system of any one of claims 1-36.

39. A method of detecting or quantifying an analyte, comprising: a) providing the closed bipolar electrode system of any one of claims 1-36; b) introducing an analyte into the sample compartment; c) applying a sweeping potential to the system; and d) measuring the change in open-circuit potential (OCP) at the bipolar electrode in the measurement compartment.

40. The method of claim 39, wherein the potential is a square wave.

41. The method of claim 39, wherein the potential is a cyclic, sweeping potential.

42. The method of claim 41, wherein the scan rate of the potential ranges from about 10 mV/sec to about 200 mV/sec.

43. The method of any one of claims 39-42, wherein the potential ranges from about -1.0 V to about 1.0 V.

44. The method of any one of claims 39-43, wherein the analyte is selected from the group consisting of an enzyme, a sugar, a nucleic acid, an amino acid, an ion, a heavy metal, a bacteria, a virus, and a metabolite.

45. The method of any one of claims 39-43, wherein the analyte is selected from the group consisting of D-serine, lactate, glucose, glycated proteins, glycated amino acid, hydrogen peroxide, cholesterol, glycerol, glycerol-3-phosphate, fructose, urate, ethanol, galactose, 1,5-anhydro-D-glucitol, NAD(P)H, dopamine, 3 -hydroxy butyrate, Levodopa (L-DOPA), L-glutamate, L-glutamine, sarcosine, creatine, and creatinine.

46. The method of any one of claims 39-43, wherein the analyte is a bacteria or a virus.