Aptamer biosensor with charge-mixed monolayer antifouling ability

Abstract

A device for continuously sensing at least one analyte in a sample liquid via measurement of the analyte is presented. The device has a sensor with a surface having a plurality of aptamers that bind to the analyte. The plurality of aptamers has at least one tag. The surface also has a protective layer between the aptamers that protects the surface from fouling. The protective layer has a monolayer of molecules that forms an interface with the sample liquid. Furthermore, the monolayer of molecules is a mixed-charge monolayer.

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims priority to U.S. Provisional Patent Application Serial No. 63 / 451,011, filed March 9, 2023, which is hereby incorporated by reference in its entirety.

[0002] The present invention relates generally to the placement of biosensors into the skin at multiple penetration depths accessible by a single device. [Background technology]

[0003] This section is intended to introduce the reader to various aspects of the art that may be related to various aspects of the present invention, which are described and / or claimed below. This discussion is believed to be helpful in providing the reader with background information to facilitate a better understanding of the various aspects of the present invention. As such, it should be understood that these statements are to be read in this light, and not as admissions of prior art.

[0004] Electrochemical aptamer sensors can determine the presence and / or concentration of an analyte of interest through the use of an aptamer sequence that specifically binds to the analyte. These sensors contain an aptamer linked to an electrode, each of which has a redox-active molecule (redox tag) linked to the electrode. The redox pair can transfer charge to or from the electrode. When the analyte binds to the aptamer, the shape of the aptamer changes, and the redox pair moves closer to or further away from the electrode on average. This causes a measurable change in current that can be converted into an indication of the presence or concentration of the analyte. When used in this way, aptamers are an example of affinity-based biosensors.

[0005] A significant unresolved issue for aptamer sensors and other affinity-based biosensors (especially sensors in which the aptamer is bound to the working electrode) is sensor lifetime, especially for applications requiring continuous operation ("continuous" refers to multiple measurements made over time by the same device). Such aptamer sensors are prone to degradation due to, among other things, the detachment of the aptamer itself from the electrode and / or the detachment of protective layer molecules (e.g., mercaptohexanol, etc.) from the electrode. The aptamer and protective molecules together form a monolayer that may be referred to as the sensing monolayer. The protective layer portion of the sensing monolayer (1) ensures that the conformational change of the aptamer upon binding to the analyte is not physically obstructed by fouling substances and (2) reduces background currents (including oxygen reduction currents) that would otherwise cancel out the measurement signal generated by the aptamer and redox tag.

[0006] Current fabrication methods for these devices use a very simple and convenient approach: forming a partial monolayer of aptamer by thiol-coupling to a gold electrode via incubation of the electrode in a solution containing the aptamer, followed by a more complete monolayer containing a protective molecule such as mercaptohexanol (via incubation of the electrode in a mercaptohexanol solution). This process is highly convenient for researchers because not only does the mercaptohexanol monolayer reduce background current, but the mercaptohexanol monolayer, as typically formed, also has at least one feature, e.g., defects, that allows electron transfer between the redox tag and the electrode; these defects are few and / or small enough to minimize oxygen reduction current and other major sources of background current. Furthermore, the mercaptohexanol monolayer has the right defects for electron transfer to support a zero-gain frequency, enabling two-frequency or equivalent self-calibration techniques. Finally, mercaptohexanol has sufficient surface fouling resistance to allow short-term in-lab experiments in biological fluids such as blood and serum.

[0007] Thus, researchers have had at their disposal a very "convenient" method for creating aptamer sensors for research applications. However, most aptamer researchers have not previously been motivated to address the lifetime of aptamer sensors, and similarly convenient monolayer approaches are inherently fragile because the monolayer may detach over time. This detachment is due in part to the fact that each portion of the monolayer is a single molecule with a single bond to the electrode, and statistically or energetically, it is not very difficult to break one of these bonds to the electrode with traditional monolayer chemistry, especially at elevated temperatures such as body temperature. Multiple bonds to gold may alleviate this challenge, but may still lack the close packing density required to suppress background current during measurements. According to leading experts in a 2022 review article (see Shaver et al., "The challenge of long-term stability for nucleic acid-based electrochemical sensors," Current Opinion in Electrochemistry (2022), 32: 100902 (https: / / doi.org / 10.1016 / j.coelec.2021.100902)), "Unfortunately, these chemistries desorb over time upon exposure to environmental or experimental factors, such as dry air, high temperatures, voltage pulses, and biological fluids. This desorption process simultaneously removes the sensing moiety and passivating thiol from the electrode surface, preventing their deployment for more than a few hours." Clearly, even for experts in the field, aptamer sensor lifetime remained an open problem with no clear solution in sight, even when attempting to achieve sensor lifetimes of days or weeks. Even if alternative methods are developed to extend the lifetime of the protective layer, these methods must also support proper electron transfer to transmit the sensor signal, ideally allow the use of one or more calibration-free operating methods, and prevent excessive fouling that would otherwise inhibit aptamer movement and thus proper sensor signaling.Inventor Young and co-inventor Heikenfeld recently reported a breakthrough in aptamer sensor longevity by stabilizing an alkylthiolate-based protective layer system such that aptamer sensors have operational lifetimes exceeding one week. See Watkins et al., "Week-Long Operation of Electrochemical Aptamer Sensors: New Insights into Self-Assembled Monolayer Degradation Mechanisms and Solutions for Stability in Serum at Body Temperature," ACS Sens. 2023, 8, 3, pp. 1119–1131, published March 8, 2023, https: / / doi.org / 10.1021 / acssensors.2c02403 (hereinafter referred to as the "Young and Heikenfeld lifetime paper"). This paper details the degradation mechanisms of aptamer sensors and demonstrates the utility of membrane protection to preserve aptamer sensor function by preventing monolayer fouling over multiple days.

[0008] For sensing larger analytes (more than a few kDa or tens of kDa) or for sensing over multiple weeks, membrane protection may not be adequate, as if the target analyte reaches the sensing surface, some fouling material may also reach the sensor surface. Motivated to address the operation of aptamer sensors without membrane protection, even for short-term (<24 h) sensor operation, researchers have used dual-frequency "calibration-free" operation as taught in, for example, Li H, Dauphin-Ducharme P, Ortega G, Plaxco KW. Calibration-Free Electrochemical Biosensors Supporting Accurate Molecular Measurements Directly in Undiluted Whole Blood. J Am Chem Soc. 2017 Aug 16;139(32):11207-11213. doi: 10.1021 / jacs.7b05412. Epub 2017 Aug 2. PMID: 28712286; PMCID: PMC6519131, and other methods to mitigate the effects of sensor degradation in whole biological fluids, such as those taught in Arroyo-Curras N, Somerson J, Vieira PA, Ploense KL, Kippin TE, Plaxco KW. Real-time measurement of small molecules directly in awake, ambulatory animals. Proc Natl Acad Sci U S A. 2017 Jan 24;114(4):645-650. doi: 10.1073 / pnas.1613458114. Epub 2017 Jan 9. PMID: 28069939; PMCID: PMC5278471.

[0009] The above-referenced lifetime paper by Young and Heikenfeld was the first to demonstrate, in a subsection of their paper entitled "Fouling Affects Electron Transfer Rates and Has a Significant Impact on Freedom of Mobility of The Aptamer," that fouling is a remaining challenge for aptamer biosensors that achieve multi-day or even multi-week operation. Young and Heikenfeld's lifetime paper and previous papers: Li H, Dauphin-Ducharme P, Arroyo-Curras N, Tran CH, Vieira PA, Li S, Shin C, Somerson J, Kippin TE, Plaxco KW. A Biomimetic Phosphatidylcholine-Terminated Monolayer Greatly Improves In Vivo Performance of Electrochemical Aptamer-Based Sensors. Angew Chem Int Ed Engl. 2017 Jun 19;56(26):7492-7495. doi: 10.1002 / anie.201700748. Epub 2017 Mar 28. PMID: 28371090; PMCID: As demonstrated in PMC5660315, zwitterion-terminated monolayers can be used to reduce fouling of the sensor surface and reduce or eliminate the need for a protective membrane. However, zwitterion-terminated monolayers can also be detrimental to the operation of the sensor, and alternative anti-fouling approaches may still be needed.

[0010] Therefore, there remains a need for alternative anti-fouling chemistries that allow optimal sensor operation in terms of maintaining accuracy and achieving desired detection ranges over multiple days or even weeks. [Prior art documents] [Non-patent literature]

[0011] [Non-Patent Document 1] Shaver et al., “The challenge of long-term stability for nucleic acid-based electrochemical sensors,” Current Opinion in Electrochemistry (2022), 32: 100902(https: / / doi.org / 10.1016 / j.coelec.2021.100902) [Non-patent document 2] Watkins et al., "Week-Long Operation of Electrochemical Aptamer Sensors: New Insights into Self-Assembled Monolayer Degradation Mechanisms and Solutions for Stability in Serum at Body Temperature," ACS Sens. 2023, 8, 3, pp. 1119-1131, Published: March 8, 2023, https: / / doi.org / 10.1021 / acssensors.2c02403 [Non-patent document 3] Li H, Dauphin-Ducharme P, Ortega G, Plaxco KW. Calibration-Free Electrochemical Biosensors Supporting Accurate Molecular Measurements Directly in Undiluted Whole Blood. J Am Chem Soc. 2017 Aug 16;139(32):11207~11213. doi: 10.1021 / jacs.7b05412. Epub August 2, 2017. PMID: 28712286; PMCID: PMC6519131 [Non-patent document 4] Arroyo-Curras N, Somerson J, Vieira PA, Ploense KL, Kippin TE, Plaxco KW. Real-time measurement of small molecules directly in awake, ambulatory animals. Proc Natl Acad Sci US A. 2017 Jan 24; 114(4):645-650. doi: 10.1073 / pnas.1613458114. Epub January 9, 2017. PMID: 28069939; PMCID: PMC5278471 [Non-patent document 5] Li H, Dauphin-Ducharme P, Arroyo-Curras N, Tran CH, Vieira PA, Li S, Shin C, Somerson J, Kippin TE, Plaxco KW. A Biomimetic Phosphatidylcholine-Terminated Monolayer Greatly Improves In Vivo Performance of Electrochemical Aptamer-Based Sensors. Angew Chem Int Ed Engl. June 19, 2017; 56(26):7492~7495. doi: 10.1002 / anie.201700748. Epub March 28, 2017. PMID: 28371090; PMCID: PMC5660315 Summary of the Invention [Means for solving the problem]

[0012] Certain illustrative aspects of the present invention are described below. It should be understood that these aspects are presented merely to provide the reader with a brief summary of certain forms that the invention may take, and that these aspects are not intended to limit the scope of the invention. Indeed, the present invention may encompass a variety of aspects that may not be explicitly set forth below.

[0013] Many of the above-mentioned drawbacks and limitations can be resolved by creating new and advanced interactions of chemicals, materials, sensors, electronics, microfluidics, algorithms, computing, software, systems, and other features or designs in a manner that brings sensing technology into proximity with biological fluids and analytes affordably, effectively, conveniently, intelligently, or reliably.

[0014] One embodiment of the present invention relates to a device for continuously sensing at least one analyte in a sample liquid via analyte measurement. The device includes at least one sensor. The sensor surface has a plurality of aptamers that bind to the analyte. The aptamers carry at least one tag that causes a change in at least one parameter when the analyte binds to the aptamer. The sensor surface also has a protective layer between the aptamers that protects the surface from fouling. The protective layer further includes a monolayer of molecules that forms an interface with the sample liquid. The monolayer of molecules is a mixed-charge monolayer of at least a first molecule and a second molecule. The first molecule has a first charge that confers a net negative charge at the interface with the sample liquid. The second molecule has a second charge that confers a net positive charge near the interface with the sample liquid. Compared to other antifouling surface chemistries, this novel approach to aptamer sensors can improve the electron transfer rate, improve the overall magnitude of the electron transfer current, and create optimal binding affinity of the analyte to the aptamer and the resulting detection range of the analyte.

[0015] The present invention relates to a device for continuously sensing at least one analyte in a sample liquid via analyte measurement. The device includes at least one sensor having a surface. The surface has a plurality of aptamers that bind to the analyte. The plurality of aptamers carry at least one tag that causes a change in at least one parameter when the analyte binds to the aptamer. The surface also has a protective layer between the aptamers that protects the surface from fouling. The protective layer has a monolayer of molecules that forms an interface with the sample liquid. Furthermore, the monolayer of molecules is a mixed-charge monolayer consisting of at least a first molecule and a second molecule. The first molecule has a first charge that imparts a net negative charge at the interface with the sample liquid. The second molecule has a second charge that imparts a net positive charge near the interface with the sample liquid.

[0016] In one embodiment, the first charge and the second charge have a magnitude of at least one electron charge. In another embodiment, the monolayer of molecules comprises a repeating hydrophobic molecular chain selected from the group consisting of two or more methyl groups, two or more phenyl groups, and combinations thereof. In one embodiment, the mixed-charge monolayer comprises one or more subsets of molecules, each subset of molecules having termini, and further the mixed-charge monolayer has mixed charges at the termini of at least one subset of molecules. In another embodiment, the mixed-charge monolayer has mixed charges located along the interior chains of at least one subset of molecules. In one embodiment, the mixed-charge monolayer has mixed charges located both at the termini of at least one subset of molecules and along the interior chains of at least one subset of molecules.

[0017] In another embodiment, the protective layer has a resistance of 200 μF / cm 2 Less than 20μF / cm 2 Less than 10 μF / cm 2 Less than 5 μF / cm 2 Less than 3 μF / cm 2 and less than 1 μF / cm 2The device has a capacitance that is a value selected from the group consisting of less than 1000 kJ / cm. In one embodiment, the capacitance is limited by a non-monolayer blocking layer to which the monolayer is attached. In another embodiment, the non-monolayer blocking layer comprises a material selected from the group consisting of semiconductors, metal oxides, nitrides, carbides, and mixtures thereof. In one embodiment, the sample fluid is a biological fluid selected from the group consisting of interstitial fluid, serum, and blood. In another embodiment, the device also includes an electrode, the electrode carrying a protective layer.

[0018] In one embodiment, the sensor has an initial redox tag current when placed in a sample solution, and the redox tag current decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during the first 6 hours of operation at a temperature of at least 30° C. In another embodiment, the sensor has a sensor response that decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during at least 3 days of operation at a temperature of at least 30° C. In one embodiment, the sensor has a sensor response that decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during at least 7 days of operation at a temperature of at least 30° C.

[0019] In another embodiment, the sensor has a measurable oxygen reduction current at -0.5V, which increases by an amount selected from the group consisting of less than 10%, less than 30%, and less than 100% relative to the magnitude of the background current over at least three days of operation. In one embodiment, the sensor has a measurable oxygen reduction current at -0.5V, which increases by an amount selected from the group consisting of less than 10%, less than 30%, and less than 100% relative to the magnitude of the background current over at least seven days of operation. In another embodiment, the mixed charge monolayer has a net charge density of less than 1 nm 2The mixed charge monolayer has a net charge equal to a four electron charge per ion, and the net charge is less than a percentage selected from the group consisting of 50%, 25%, 5%, and 1%. In one embodiment, the mixed charge monolayer has a net charge, and the net charge is negative.

[0020] In another embodiment, the mixed-charge monolayer has a net charge, the net charge is positive, and a voltage is used to operate the sensor, the voltage being sufficiently negative so that the net potential at the interface between the sample liquid and the monolayer is negative. In one embodiment, most of the first charges and most of the second charges are separated by a distance selected from the group consisting of 0.5 nm, 1.0 nm, and 1.5 nm. In another embodiment, most of the first charges and most of the second charges are separated by a number of methyl groups selected from the group consisting of 5, 10, and 15. In one embodiment, most of the first charges and most of the second charges are separated by a number of phenyl groups selected from the group consisting of 1, 2, and 3.

[0021] In another embodiment, the first charge has a first distance from the electrode while the positive charge has a second distance from the electrode, and the first distance exceeds the second distance such that the interface between the monolayer and the sample liquid has a net negative charge. In one embodiment, the mixed-charge monolayer has a density of at least 0.5 molecules / nm 2 , 1 molecule / nm 2 , 2 molecules / nm 2 , 3 molecules / nm 2 , 4 molecules / nm 2 , and 5 molecules / nm 2In another embodiment, the device also includes a background current, and the background current increases by less than a percentage selected from the group consisting of 5%, 10%, 20%, and 40% over a period of at least 3 days. In one embodiment, the device also includes a background current, and the background current increases by less than a percentage selected from the group consisting of 5%, 10%, 20%, or 40% over a period of at least 7 days. In another embodiment, the mixed-charge monolayer molecules comprise a long hydrophobic alkyl chain length selected from the group consisting of 6 methyl groups in length and 8 methyl groups in length.

[0022] In one embodiment, the mixed-charge monolayer molecules are linked to the device using a chemical structure selected from the group consisting of thiol, silane, phosphonic acid, trichlorosilane, trimethoxysilane, triethoxysilane, phosphate, alkene, and alkyne. In another embodiment, the tag is a redox tag, the device has an electron transfer rate, and the electron transfer rate changes by less than a percentage selected from the group consisting of 50%, 20%, 10%, and 5% over a period of at least three days. In one embodiment, the tag is a redox tag, the device has an electron transfer rate, and the electron transfer rate changes by less than a percentage selected from the group consisting of 50%, 20%, 10%, and 5% over a period of at least seven days. In another embodiment, the boundary between the protective layer and the sample liquid is within a range of values ​​selected from the group consisting of ±5 mV, ±10 mV, ±20 mV, ±40 mV, and ±80 mV of a zero volt surface potential.

[0023] In one embodiment, the boundary between the protective layer and the sample liquid has a net positive charge when no voltage is applied to the electrode, and a negative voltage is used during operation of the electrode. In another embodiment, the boundary between the protective layer and the sample liquid has a net negative charge when no voltage is applied to the electrode, and a positive voltage is used during operation of the electrode. In one embodiment, the device also includes a plurality of molecular brush molecules interspersed between the aptamers. In another embodiment, the boundary layer and the mixed-charge monolayer have a Debye length, and further, the molecular brush has a net charge in the vicinity of the aptamer and extends at least twice the Debye length from the boundary layer including the mixed-charge monolayer.

[0024] In one embodiment, the sample fluid is interstitial fluid within a human body. In another embodiment, the device also includes a first surface area between the sample fluid and the sensor, wherein there is less than 30% area coverage of foreign body response-elicited cells on the first surface area for a period selected from the group consisting of 3 days, 7 days, and 14 days. In one embodiment, the device also includes a first surface area between the sample fluid and the sensor, wherein there is less than 10% area coverage of foreign body response-elicited cells on the first surface area for a period selected from the group consisting of 3 days, 7 days, and 14 days. In another embodiment, the at least one tag is a redox tag and the parameter is a redox current. In one embodiment, the parameter is electrical impedance. In another embodiment, the parameter is light intensity.

[0025] In one embodiment, the device also comprises a binding affinity between the analyte and multiple aptamers, wherein the binding affinity is within at least one of ±10, 50, 100, 200, 300, 500, and 1000% of the analyte binding affinity for the same aptamer when tested in the same manner except using mercaptohexanol or mercaptooctanol instead of the mixed-charge monolayer.

[0026] In another embodiment, the present invention relates to a method for sensing at least one analyte in a sample liquid via measurement of the analyte. The method comprises exposing the sample liquid to a sensor as described above and using data relating to changes in one or more aptamer parameters to measure the analyte. A voltage is applied to the electrodes such that the potential at the interface between the monolayer and the sample liquid has a net negative potential, causing most foulants and aptamers to be slightly repelled from the surface.

[0027] The objects and advantages of the disclosed invention will be further understood in light of the following detailed description and drawings. [Brief explanation of the drawings]

[0028] [Figure 1A] FIG. 1 shows a schematic diagram of one embodiment of a conventional prior art sensor device. [Figure 1B] FIG. 1 shows a schematic diagram of another embodiment of a conventional prior art sensor device. [Figure 2] 1 shows a schematic diagram of one embodiment of a device according to the principles of the present invention. [Figure 3] 1 shows a schematic diagram of another embodiment of a device according to the principles of the present invention; [Figure 4] 1 shows a schematic diagram of yet another embodiment of a device according to the principles of the present invention; [Figure 5A] 1 shows a schematic diagram of a prior art device. [Figure 5B] 1 shows a schematic diagram of yet another embodiment of a device according to the principles of the present invention; [Figure 5C] 1 shows a schematic diagram of yet another embodiment of a device according to the principles of the present invention; [Figure 6A] FIG. 1 shows a schematic of an electrical measurement technique using square wave voltammetry. [Figure 6B] FIG. 1 shows graphs representing "forward," "backward," and "net" voltammograms. [Figure 6C]FIG. 1 shows a graph representing frequency, where the sensor response is "ON" (redox current increases) or "OFF" (redox current decreases) when the analyte binds to the aptamer. [Figure 7A] 1 shows a schematic diagram of yet another embodiment of a device according to the principles of the present invention; [Figure 7B] 1 shows a schematic diagram of yet another embodiment of a device according to the principles of the present invention; [Figure 8] FIG. 10 shows a graph depicting titration results as a plot of sensor response against total cortisol concentration. DETAILED DESCRIPTION OF THE INVENTION

[0029] definition As used herein, "continuous sensing" by "continuous sensor" refers to a sensor that changes in response to changes in the concentration of at least one solute, e.g., an analyte, in a solution. Similarly, as used herein, "continuous monitoring" refers to the ability of a device to provide multiple measurements of an analyte over time.

[0030] As used herein, the term "about," when referring to a value or amount of mass, weight, time, volume, pH, size, concentration, or percentage, is meant to encompass variations of, in some embodiments, ±20%, in some embodiments, ±10%, in some embodiments, ±5%, in some embodiments, ±1%, in some embodiments, ±0.5%, and in some embodiments, ±0.1% from the specified amount, since such variations are appropriate to perform the disclosed methods.

[0031] As used herein, the term "electrode" refers to any electrically conductive material, such as gold, platinum, nickel, silicon, conductive liquid infusion materials such as ionic liquids, PEDOT:PSS, conductive oxides, carbon, boron-doped diamond, nanotube or nanowire mesh, or other suitable electrically conductive materials.

[0032] As used herein, the term "protective layer" means a single layer protective layer or a non-single layer protective layer, or a combination of both.

[0033] As used herein, the term "protective monolayer" refers to a uniform or heterogeneous layer of material or one or more types of molecules that can enable at least one of: a reduction in electrochemical background current and / or current due to electrochemical interference; a reduction in fouling in the sample liquid; and promote adequate freedom of movement of aptamers required to elicit a measurable response to analyte concentration.

[0034] As used herein, the term "non-monolayer protective layer" refers to a uniform or non-uniform layer of material or one or more types of molecules on an electrode that does not exhibit a monolayer configuration and that can reduce electrochemical background current and / or current due to electrochemical interference; reduce fouling in the sample solution; and facilitate adequate freedom of movement of aptamers required to elicit a measurable response to analyte concentration. For example, a metal or semiconductor oxide can be a non-monolayer protective layer, or a thin polymer film can be a non-monolayer protective layer because it is composed of multiple layers of atoms or molecules. For example, a single atomic monolayer of SiO2 is a monolayer, while 3 nm of SiO2 is a non-monolayer. A non-monolayer protective layer can be, for example, a nitride, such as silicon nitride, or other suitable materials that provide similar functions, including carbides and other suitable materials.

[0035] As used herein, the term "mixed-charge monolayer" refers to a monolayer of at least partially vertically oriented molecules on a surface, which may be a protective monolayer, comprising at least a first plurality of molecules having a first charge polarity at or near their end facing the sample liquid and at least a second plurality of molecules having a second charge polarity at or near their end facing the sampling liquid, wherein the first and second polarities have opposite charges. Mixed charges can also include monolayers containing uncharged molecules such as mercaptohexanol or mercaptooctanol or other uncharged molecules, even if not 100% of the molecules in the monolayer, since mixed charges can be used to increase the hydration of the monolayer surface and its ability to prevent fouling and / or modify the net surface potential, which may affect the binding affinity of aptamers for analytes. Mixed-charge monolayers can also include monolayers containing uncharged molecules such as mercaptohexanol or mercaptooctanol or other uncharged molecules, so long as at least one percent (1%) of the molecules in the monolayer are mixed-charge.

[0036] As used herein, the term "aptamer" refers to a molecule that undergoes a conformational or binding change upon binding of an analyte to the molecule, and satisfies the general operating principle of the sensing method described herein. Such molecules include, for example, natural or modified DNA, RNA, or XNA oligonucleotide sequences, spiegelmers, peptide aptamers, and affimers, as well as other affimer-based probes. Modifications may include substituting non-natural nucleobases for natural bases in the aptamer sequence, substituting non-natural sequences for natural sequences, or other suitable modifications that improve sensor function but behave similarly to conventional aptamers. Two or more aptamers bound together may also be referred to as aptamers (i.e., do not separate in solution). Aptamers may have a molecular weight of at least 1 kDa, 10 kDa, or 100 kDa.

[0037] As used herein, the term "tag" refers to a molecule carried on an aptamer that has a measurable response when an analyte binds to the aptamer, such as a redox tag, or a fluorescent tag or quencher tag such as those used in molecular beacons, or some other suitable tag that is measurable. In the case of a redox tag, the aptamer is typically carried on or near an electrode; when a fluorescent tag is employed, the aptamer is often carried on or near an optical waveguide.

[0038] As used herein, the term "redox tag" or "redox molecule" refers to any species, e.g., a small or large molecule, having a redox-active moiety that can reversibly donate or accept at least one electron to or from an electrode when introduced adjacent to the electrode. Examples of redox tags or molecules include methylene blue, ferrocene, quinones, or other suitable species that meet the definition of a redox tag or molecule. In some cases, redox tags or molecules are referred to as redox mediators. Redox tags or molecules may also exchange electrons or change behavior when introduced adjacent to other redox tags or molecules. Exogenous redox molecules are those that are added to the device, e.g., they are not endogenous but are provided by the sample solution being tested.

[0039] As used herein, the term "change in electron transfer" refers to a redox molecule whose electron transfer to or from an electrode has been altered in a measurable manner. This change in electron transfer may be due, for example, to electron transfer availability, distance from the electrode, diffusion rate to or from the electrode, a shift or increase or decrease in the electrochemical activity of the redox molecule, or any other embodiment taught herein that results in a measurable change in electron transfer between the redox molecule and the electrode.

[0040] As used herein, the term "sensing monolayer" refers to at least a plurality of aptamers on a working electrode, and may also include a plurality of molecules or a mixture of molecules that form a non-monolayer protective layer or a monolayer protective layer.

[0041] As used herein, the term "analyte" means any solute in a solution or liquid or sample fluid that can be measured using a sensor. An analyte can be a small molecule, protein, peptide, electrolyte, acid, base, antibody, small molecule-bound molecule, DNA, RNA, drug, chemical, pollutant, or other solute in the solution or liquid.

[0042] As used herein, the term "sample liquid" is a liquid that contains an analyte.

[0043] As used herein, the term "device" includes at least one sensor based on at least one aptamer and at least one sample solution. The device can sense multiple samples and can take multiple configurations, such as a device for measuring a finger stick of blood, or a microneedle or indwelling sensor needle for measuring interstitial fluid, or a device for measuring saliva, tears, sweat, or urine, or a device for measuring water contaminants or food processing solutes, or other device that measures at least one analyte found in the sample solution.

[0044] As used herein, "redox tag current" refers to the amplitude of the Faradaic redox tag peak current (e.g., as collected in a voltammogram) minus the amplitude of the background current outside the redox peak in a given voltammetric scan. "Normalized redox tag current" refers to the redox tag current normalized to the initial measurement. "Background current" refers to the current measured when the aptamer molecule is not tagged with a redox reporter, including, for example, capacitive current and competing redox processes such as oxygen reduction. "Adjusted current" refers to the combination of the redox tag current and background current in a square wave voltammogram, where the minimum current is set to 0 A in the voltammogram presentation, thus allowing voltammograms to be plotted side-by-side and more easily compared. "Sensor response" is the change in redox tag current due to binding of the target analyte to the aptamer, also known as signal gain, which may increase or decrease depending on the aptamer and the voltammetric time scale (e.g., operating frequency in square wave voltammetry, etc.). The above term can be used to describe or interpret other types of sensor scanning techniques, e.g., chronoamperometry, and should not be narrowly construed to specifically refer to only scanning methods, e.g., square wave voltammetry.

[0045] Detailed Description of the Invention One or more specific embodiments of the present invention are described below. While an effort has been made to describe these embodiments concisely, not all features of an actual implementation are described herein. It should be understood that in developing any such actual implementation, as with any engineering or design project, numerous implementation-specific decisions must be made to achieve the developer's specific goals, such as adhering to system-related and business-related constraints that may vary from implementation to implementation. It should further be understood that such a development effort may be complex and time-consuming, but would nevertheless be a routine undertaking of design, fabrication, and manufacture for those of ordinary skill in the art having the benefit of this disclosure.

[0046] Certain embodiments of the disclosed invention present sensors as individual, simple elements. It is understood that many sensors require two or more electrodes, reference electrodes, or additional supporting technologies or features not captured in the description herein. Sensors may be dual, triple, or more to provide improved data and readings. Sensors may provide continuous or discrete data and / or readings. Certain embodiments of the disclosed invention highlight the subcomponents of sensing devices with many more subcomponents (e.g., reference or counter electrodes, batteries, antennas, adhesives) that are required and known for using the device in various applications; for the sake of brevity and to focus on the inventive aspects, such components may not be explicitly shown in the diagrams or described in the embodiments of the disclosed invention. All ranges of parameters disclosed herein include the endpoints of the ranges.

[0047] 1A and 1B, a conventional prior art sensor device 100 is presented that is initially placed in a sample fluid 130, such as interstitial fluid. The device 100 includes at least one working electrode 120, e.g., gold, carbon, or other suitable electrode material; a single protective layer 126, e.g., at least one protective layer of a plurality of molecules, e.g., mercaptohexanol, thiol-linked to the electrode; at least one aptamer 124 responsive to binding to an analyte 180; and a redox tag 170, e.g., methylene blue, associated with, e.g., bound to, the at least one aptamer. In the general example taught in FIGS. 1A and 1B, the aptamer 124 is a simple stem-loop (hairpin) aptamer in which binding of the analyte 180 forms a stem-loop, resulting in an increase in the measured redox-tag current from the redox tag 170, as measured using square-wave voltammetry ( FIGS. 6A and 6B ), chronoamperometry, or other suitable techniques. In the absence of binding of the analyte 180 to the aptamer 124, the stem-loop conformation is not formed, and therefore the redox current does not increase. Alternatively, binding of the analyte 180 to a different instance of the aptamer can cause a decrease in the redox-tag current. Thus, when using an aptamer sensor, changes in the electrical redox-tag current measurement can be used as a signal to convert into changes in the concentration of the analyte 180.

[0048] With further reference to FIG. 1B, a challenge with aptamer sensors is that when placed in sample solution 130 during initial operation, over a period of minutes to hours to days, depending on the protective layer 126, the signal (e.g., redox tag current) may decrease by 30%, 50%, or more due to effects such as fouling by small molecules 186, proteins 188, or other solutes in the sample solution 130, as well as due to detachment of the sensing monolayer containing the aptamer 124 and / or protective layer 126.

[0049] As can be seen, currently available sensor devices contain several drawbacks and limitations, particularly due to the protective layers found in such current devices. However, various aspects of the present invention address such drawbacks and limitations, including initial or long-term release, instability, and / or fouling of the sensing monolayer. In this regard, one aspect of the present invention relates to a device for detecting the presence or measuring the concentration or amount of at least one analyte in a sample liquid. In particular, the present invention reduces or overcomes fouling while achieving at least one of maintaining a fast electron transfer rate; enabling a strong electron transfer current; and enabling optimal binding affinity of the analyte to the aptamer and resulting detection range for the analyte.

[0050] Referring to Figures 2, 3, and 4, several embodiments illustrating device structures of the present invention are first presented, where like numerals refer to like features. The device includes electrodes 220, 320, and 420, which can be metals such as gold or platinum, semiconductors such as silicon, conductive oxides such as In2O3:SnO2, carbon such as glassy carbon or diamond-like carbon, or other suitable electrodes. For example, electrodes such as glassy carbon or diamond-like carbon are considered preferable because they suppress the oxygen reduction current, a form of background current, by at least a factor of three at an applied voltage of approximately -0.3 V compared to electrodes such as gold. The device also includes at least one protective layer, which can be included in at least a portion of a monolayer, such as 226, 326, or 426. The protective layer can be included in at least a portion of a non-monolayer layer 322, 422, such as 1-2 nm of SiO2, Al2O3, or other suitable material that is sufficiently porous for redox electron transfer. The device may have aptamers linked to the electrodes 220, 420, or to the non-monolayer 322. Linking chemistries for the aptamers and monolayers may include thiols, silanes, phosphonic acids, diazonium salts, or other suitable chemistries optimal for each electrode type, as will be understood by those skilled in the art of monolayer chemistry. Monolayers may be formed from molecules consisting of chains of methyl groups such as in alkylthiolates or phenyl groups in organosilanes, or other suitable chains.

[0051] Referring to Figure 5, the prior art is illustrated in Figure 5A, and an embodiment of the present invention is taught in Figure 5B. One embodiment of the present invention, shown in Figure 5B, uses a protective layer that significantly reduces fouling of electrochemical aptamer sensors by introducing mixed charges into the monolayer spanning distances 590 and 592, preferably using a monolayer that exhibits less shedding and is more stable than the prior art, e.g., mercaptohexanol, illustrated in Figure 1B. The prior art in Figure 5A is an example of a zwitterionic monolayer, e.g., a phosphatidylcholine-terminated monolayer (containing both + and - charges per molecule in the monolayer at the pH of the sample solution), which can be coupled to a substrate using thiols or other suitable chemistries. The prior art of FIG. 5A provides two useful anti-fouling properties: the zwitterionic charge promotes water binding in a monolayer (extending partially to depth 590) that acts as a partial shield against fouling by repelling fouling substances 588, and the surface charge is negative in region 592, which is useful since most fouling substances, such as albumin, have a negative net charge. While the prior art of FIG. 5A can reduce fouling, it may suffer from several potential drawbacks, including, but not limited to, poor monolayer stability and resulting monolayer detachment; poor electron transfer through the monolayer; field effects or steric effects on negatively charged aptamers (not shown), such as a negative surface charge as shown in FIG. 5A, which repel and weaken the binding affinity of the aptamer for its analyte; field effects or steric effects on charged aptamers (not shown), such as a positive surface charge, which may attract the negatively charged aptamer to the surface and reduce its ability to bind the analyte or have a significant change in electron transfer due to binding; field effects or steric effects on charged analytes, such as a negative surface charge as shown in FIG. 5A, which may repel the negatively charged analyte of interest and inadvertently shift the binding affinity between the aptamer and the analyte; poor packing density between adjacent molecules within the protective monolayer; or other potential drawbacks not specifically described herein.

[0052] Referring to FIG. 5B , the mixed charge in region 592 creates a stronger dipole locally because oppositely charged dipole elements are closer near the boundary between sample liquid 592 and the monolayer, thus binding more water and reducing fouling. While the charges are presented as highly periodic (alternating) in FIG. 5B , this need not be the case. The present invention may further include zwitterionic charges in region 590 (not shown in FIG. 5B ). If mixed charges are found at the boundary between the monolayer and sample liquid in region 592 (the end portions facing sample liquid 520) and within monolayer region 590 (referred to herein as internal charges, or internal mixed charges when mixed), even more bound water is possible, as illustrated in the embodiment of FIG. 5C . As a result, if a fouling substance penetrates region 592, it must also penetrate region 590 of FIG. 5C and its bound water, which may be much more difficult to foul than the monolayer of FIG. 5B . The chemistry of the mixed-charge monolayer spanning some regions, e.g., regions 592 and / or 590, can also connect to non-monolayer surfaces, e.g., 322 and 422. Further referring to embodiments of the present invention, in one embodiment, the monolayer is composed of zwitterions, where one molecule has a phosphorylcholine group at its terminus and the other molecule has a sulfobetaine group at its terminus. This molecular configuration provides a zwitterionic monolayer with both a positively charged group on one molecule closer to the liquid interface and a negatively charged group on the other molecule closer to the liquid interface, with the opposite charge on each zwitterionic molecule located deeper within the monolayer. In another embodiment of the monolayer, each molecule is individually and oppositely charged, where one molecule has a positively charged group, e.g., trimethylammonium, at its terminus and the other has a negatively charged group, e.g., carboxylate, at its terminus.

[0053] To present embodiments of the present invention, we further describe the operation of the aptamer sensor with respect to one of several possible measurement techniques, called square-wave voltammetry. Figure 6A shows a square waveform with frequency f and voltage amplitude Esw. This is superimposed on a staircase waveform that gradually increases the average voltage (average potential per pulse pair, E) by a voltage step size, Estep, every cycle. In Figure 6B, in response to this fluctuating voltage, the system generates an oscillatory faradaic current. This current is then deconvolved into "forward" and "backward" voltammograms (ifwd and ibwd) by extracting the current at specific times after the capacitive background current has at least partially dissipated. The difference between these two voltammograms then returns a more sensitive "net" voltammogram because the effects of the capacitive charging current are largely removed (they are roughly equal in magnitude regardless of polarity). As shown in Figure 6C, sensors can be designed to have optimal frequencies at which the sensor response is either "ON" (increase in redox current) or "OFF" (decrease in redox current) as the analyte binds to the aptamer. These frequencies are important not only because they represent the maximum sensor response, but also because their ratio can be used to correct for sensor drift, allowing for "calibration-free" operation as long as the electron transfer rate remains stable.

[0054] With further reference to embodiments of the present invention, a protective layer composed at least in part of elements 226, 322, 326, 422, 426 may have a mixed charge in the vicinity of the sample solution 230, 330, 430, but the total capacitance between the sample solution 230, 330, 430 and the electrodes 220, 320, 420 must not cumulatively be too high. Otherwise, techniques such as square wave voltammetry may not be usable or may result in increased electrical background current relative to the measured redox tag current. For example, a protective layer of a monolayer of alkylthiolate with a dielectric constant of about 2.6 and a thickness of about 0.7 nm or 1.4 nm may have a C=2.6 * 8.854E-14F / cm / 7E-8cm=3.3μF / cm2 The capacitance C=e * eo * A / t or C=2.6 * 8.854E-14F / cm / 1.4E-7cm=1.6μF / cm 2 C=e * eo * A / t. The capacitance of the electric double layer in most liquids (e.g., liquids 230, 330, 430, etc.) is so large that it can be ignored in this calculation (due to the thin double layer in high salt conditions such as biological fluids and the very high dielectric constant of water). This capacitance generally allows square wave voltammetry to be performed reliably up to about 1 kHz for methylene blue and about 10 kHz for redox tags, e.g., Os(II / III). Thus, the total capacitance of the protective layer at the measurement frequencies used in the device is 200 μF / cm 2 Less than 20μF / cm 2 Less than 10 μF / cm 2 Less than 5 μF / cm 2 Less than or 3μF / cm 2 Less than or equal to 1 μF / cm 2 For example, at least one of the following: 2 In this case, robust square wave voltammetry measurement frequencies may be limited to a maximum of about 10 Hz, and / or signals measured amperometrically or chronoamperometrically may be limited to slow electron transfer rates and / or low intensity signals. SiO2 is therefore one preferred material for layers 322, 422 because it can advantageously reduce the capacitance between the sample liquid and the electrode; SiO2 can be doped with carbon to reduce its dielectric constant (capacitance) from about 4 to about 2.6-2.8, referred to herein as SiOC, and has a dielectric constant of 1 μF / cm. 2 This is considered ideal to enable a capacitance of less than 2 nm. When capacitance is equal, a 2 nm thick SiO2 film is equivalent to 3 nm of SiOC. The thickness and packing density of the charge-mixed monolayer can also be designed to reduce capacitance.

[0055] With further reference to embodiments of the protective layer of the present invention, anti-fouling performance when directly exposed to human interstitial fluid or human serum at 33°C as a testable surrogate can be characterized in several ways, including, but not limited to, the following. The first measurement is raw redox current. When placed in interstitial fluid or serum unprotected (without membrane protection), aptamer sensors typically suffer from an initial 50% or even 80% decrease in redox tag current when using roughened gold electrodes, primarily due to surface fouling that impedes the transfer of redox electrons. The present invention, using a mixed-charge monolayer, can limit the loss of redox tag current to at least one of less than 5%, less than 10%, and less than 30% during the first 6 hours of sensor operation in interstitial fluid or serum. The second measurement is sensor response, which is the change in redox tag current in response to changes in analyte concentration. Generally, in the absence of protection, sensor response decreases due to fouling substances that interfere with the freedom of aptamer movement or due to other effects. The present invention, using a charge-mixed monolayer, can limit sensor response loss to less than 5%, 10%, or 30% over at least three days of operation. For example, if a sensor's redox tag current changes (increases or decreases) by 1 μA when exposed to a first concentration of analyte, the change in redox tag current will be greater than 0.8 μA when the sensor is again exposed to the same first concentration of analyte three days later. As can be seen in this example, a robust sensor response can be maintained for seven days or more if the data on day four is nearly identical to the initial data and the data on day seven declines by less than 5%, 10%, or 30%. The third measurement, measurable at -0.4 V or -0.5 V, is the prevention of a measurable increase in oxygen reduction current in the sensor voltammogram during square-wave voltammetry measurements of the sensor. The present invention, using a mixed-charge monolayer, can provide an indication of oxygen reduction current that increases by at least 10%, 30%, or 100% relative to the magnitude of the background current, where no significant oxygen reduction current occurs, over at least 7 days of operation.For a properly constructed sensor, the initial test of the sensor will result in an oxygen reduction current that is near zero or less than 10% of the baseline current, which is a preferred, non-limiting example of an initial oxygen reduction current to be used in subsequent measurements of oxygen reduction current increase.

[0056] With further reference to embodiments of the protective layer of the present invention, anti-fouling performance when directly exposed through a membrane of at least 100 kDa and using human interstitial fluid or human serum at 33°C as a testable surrogate can result in a decrease in redox tag current of at least 10% in the first two hours of operation and a loss in sensor response of at least 30% over at least three days of operation. A 100 kDa membrane is defined as maintaining the concentration of large solutes and foulants (e.g., antibodies or fibrinogen) greater than 100 kDa in the fluid between the membrane and the sensor at less than 10% of the concentration of solutes greater than 100 kDa in the sample fluid during the test period. For example, if a sensor has a 1 μA change in redox tag current upon exposure to an analyte, after three days the change in redox tag current will be greater than 0.8 μA. The present invention can further result in a loss in redox tag current of less than 30% over at least seven days of operation. Membranes are generally suitable for use with the present invention (not shown) and can be based on hydrogels, dialysis membranes, nanofiltration membranes, or other suitable materials. Mixed-charge monolayers aid in such fouling reduction, with the 100 kDa membrane reducing large foulants, whose size can significantly impact the freedom of movement of aptamers in the vicinity of the foulant, even with only slight adsorption. Some examples of membranes useful for protecting such mixed-charge monolayer sensors include the 100 kDa Repligen Spectrum® Hollow Fiber Membrane or the 100 kDa Millipore NMWL Biomax polyethersulfone ultrafiltration membrane. Other membranes capable of reducing the access of large proteins to the electrode surface can also be used to protect the electrode.

[0057] With further reference to the protective layer embodiment of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as testable surrogates, the mixed-charge monolayer exhibited a 1 nm 2 The aptamer may have a net charge that is at least one of 50%, 25%, 5%, or less than 1% of a charge density equivalent to a net charge density of 2 or 4 electrons per nanoparticle. If the net charge is too strong, it may increase fouling by attracting foulants with opposite charges, or may exert electrical forces on the aptamer, which may adversely affect the sensor response in some cases. Aptamers and foulants are generally, but not always, negatively charged. For example, a mixed-charge monolayer of 1 nm thick may be formed. 2 A monolayer is negative if it has a density of 4 monolayer molecules per nm, with 40% having positive charges and 60% having negative charges at or near its end facing the sample liquid. 2 has a net charge of 0.2 * 4 molecules / nm 2 =0.8 charge / nm 2 This is 1 nm 2 A net charge (1 nm) that is at least 50% less than the net charge equal to 1 charge per nm. 2A mixed-charge monolayer is filled with a mixed-charge monolayer having less than one charge per electrode. In another embodiment, the net charge is preferably negative, so that most foulants and aptamers are slightly repelled from the surface. In one embodiment, a monolayer filled with a negative net charge contains molecules terminated with a sulfonate group, e.g., 11-sulfanyl-1-undecanesulfonate, as well as molecules terminated with a trimethylammonium group, e.g., N,N,N-trimethyl-11-sulfanyl-1-undecanaminium. In this monolayer, the sulfonate-terminated molecules have a higher surface density than the trimethylammonium-terminated molecules, providing a negative net charge. In another embodiment, the net charge is preferably positive, but applying a voltage to the electrodes causes the potential at the interface between the monolayer and the sample to have a negative net potential, which causes most foulants and aptamers to be slightly repelled from the surface. For example, in the absence of an applied voltage, the net charge can result in a surface potential of +20 mV in the sample solution, and a continuous square-wave voltage sweep between -200 mV and -350 mV can create a net negative potential of at least -20 mV at the interface between the charge-mixed monolayer and the sample. The amount of applied voltage that appears at the interface between the monolayer and the sample depends on the porosity of the protective layer and the charge accumulation within the protective layer. In one embodiment, a charge-mixed monolayer is constructed for an aptamer sensor, containing molecules terminated with trimethylammonium groups and other molecules terminated with carboxylate groups, with the trimethylammonium-terminated molecules having a higher density than the carboxylate groups, resulting in a positive surface potential in the sample solution in the absence of an applied voltage. This higher density of positively charged molecules is achieved by varying the ratio of positively charged molecules to negatively charged molecules used during sensor preparation. When a negative voltage is applied, the surface potential becomes closer to neutral than sensors developed with mercaptohexanol, but still has an overall negative net surface potential to repel negatively charged foulants.

[0058] With further reference to embodiments of the protective layer of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as a testable surrogate, the mixed-charge monolayer can have charges appropriately separated along the molecular chain of which it is a part, or with respect to adjacent molecular chains. As a result of the appropriate separation, acid-base reactions are avoided (charges are not neutralized), and therefore sufficient distance exists to facilitate the incorporation of multiple bound water molecules. Charge separation can be defined in several ways; for example, in the case of charges at the interface between the monolayer and the sample liquid, oppositely charged charges are separated by 0.5 nm, 1.0 nm, or 1.5 nm in alternative embodiments by at least one of 5, 10, or 15 methyl groups (carbons in an alkane chain), or at least one of 1, 2, or 3 phenyl groups. In one embodiment, the charges separated along the same chain are choline groups attached to the molecule as head groups and phosphate groups incorporated into the alkane chain, where the oppositely charged molecules are separated by five carbons to facilitate proper charge separation and allow for multiple water molecules to bind. Another embodiment of this concept is a sensor having a monolayer of molecules, one terminated with a negative charge, e.g., carboxylate, and the other with a positive charge, e.g., trimethylammonium. To facilitate proper spacing of these molecules, each molecule has a spacer side chain attached to a phenyl group, which forces charge spacing during monolayer formation. In another embodiment, the negative charges are at a first distance from the electrode, while the positive charges are at a second distance from the electrode, the first distance being greater than the second distance so that a negative potential and Debye shielding occurs near the surface for the negative charges, and so that negatively charged species (e.g., aptamers or foulants, e.g., albumin) are repelled from the surface of the monolayer. In one embodiment, the monolayer is created entirely from the same molecule, where the negatively charged group, e.g., sulfonate, is located 1.5 nm from the electrode surface and separated from the positively charged group, e.g., trimethylammonium, by 0.75 nm to provide adequate Debye shielding of the negative charge and to allow the negatively charged species to repel common foulants.

[0059] With further reference to embodiments of the present invention relating to protective layers, when directly exposed to human interstitial fluid or human serum at 33°C as testable alternatives, the mixed-charge monolayer exhibits, in alternative embodiments, a resistance of at least 0.5 molecules / nm 2 , 1 molecule / nm 2 , 2 molecules / nm 2 , 3 molecules / nm 2 , 4 molecules / nm 2 , or 5 molecules / nm 2 The monolayer may have a density of 1000 Å. Generally, the denser the monolayer, the more difficult it is for fouling agents to physically penetrate and foul because the closer the mixed charges are brought together, the tighter the binding of water between them (up to the point where acid-base reactions can occur for some charged molecules). In one embodiment, the mixed-charge monolayer is composed of molecules terminated with a negatively charged group, e.g., phosphate, on one side and a positively charged group, e.g., trimethylammonium, on the other side, with the remainder of the molecules having a packing density of 1 molecule / nm after construction of the monolayer. 2 The carbon or other functional group is long enough to promote the formation of the hydroxyl group.

[0060] With further reference to embodiments of the protective layer of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as a testable surrogate, the mixed-charge monolayer may, in alternative embodiments, maintain a background current that increases by less than 5%, 10%, 20%, or 40% for at least 3 or 7 days. The background current occurs outside the range of the redox potential peak; for example, when using a methylene blue redox tag against an Ag / AgCl reference electrode, the peak redox current is approximately -300 mV, and the background is measured at, for example, -100 mV, -400 mV, or -500 mV. The background current may increase due to several factors, including increased porosity of the protective layer, which can increase both the capacitance and the charging current and / or oxygen reduction current, both of which may reduce the accuracy of the sensor. Thus, while the protective layers of the present invention may contain mixed charges within the monolayer, the incorporation of such charges may destabilize the monolayer compared to a monolayer, e.g., a pure alkylthiolate monolayer, e.g., mercaptooctanol. In the case of a mercaptooctanol monolayer on a gold electrode, when properly constructed on an appropriate roughened gold surface and subjected to appropriate electrical scanning, the hydrophobic interactions between the alkylthiolate molecules in the monolayer may result in the monolayer being stable for more than three days, or even more than seven days. Thus, the present invention may utilize monolayers that also contain long hydrophobic alkyl chain lengths of at least six, and more preferably at least eight, methyl groups (carbon units) to provide such hydrophobic interactions, where such molecules may terminate with mixed charges as taught herein. To enable even greater monolayer stability, the electrode or non-monolayer blocking layer may be constructed of a non-gold material, e.g., carbon or oxide, or other suitable material, such that monolayer molecules with mixed charges can bind to the surface with greater stability than monothiol bonds on gold, regardless of the number of carbon units or degree of hydrophobicity. Chemical structures useful in the present invention include trimethoxysilane, triethoxysilane, phosphonic acid, trichlorosilane, phosphate, alkene, or alkyne.In one embodiment, a non-monolayer protective layer, such as SiO2, is deposited on the gold electrode to provide a surface for alternative chemical structures. A combination of two phosphonic acids is used to create a mixed-charge monolayer, where one has a primary amine at its end and the other has a carboxylate group at its end. In one embodiment, they self-assemble on the SiO2 substrate. Under physiological conditions, the primary amine has a positive charge and the carboxylate group has a negative charge, providing a mixed charge within the monolayer.

[0061] With further reference to embodiments of the protective layer of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as a testable surrogate, the mixed-charge monolayer can provide a redox tag current that is strong enough to achieve accurate sensor measurements, e.g., ±10% or ±20% sensor accuracy. A protective layer that is too thick or too dense can suppress fouling and background current, but does so at the expense of redox tag current flow. Therefore, the protective layer must remain permeable to electron transfer between the redox tag and the electrode. Furthermore, to reduce measurement inaccuracies and enable dual-frequency or continuous square-wave calibration-free operation, the electron transfer rate should remain stable. In alternative embodiments, the electron transfer rate should change by less than 50%, 20%, 10%, or 5% over a period of at least three days or at least seven days. In one embodiment, a mixed-charge monolayer for an aptamer sensor is constructed from a combination of molecules, one of which has a phosphorylcholine group at its end and the other of which has a sulfobetaine group at its end, where the chain length of each molecule is long enough to promote favorable stabilizing interactions between the molecules but short enough not to interfere with the initial electron transfer rate. In alternative embodiments, after continuous scanning for at least 3 days or at least 7 days, the sensor remains stable due to the mixed-charge monolayer, and thus the electron transfer rate changes by less than 20%.

[0062] Further referring to embodiments relating to the protective layer of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as a testable alternative, the mixed-charge monolayer can be optimally fabricated to be compatible with the voltage scan used for the aptamer sensor, so that a strong positive or negative potential does not occur at the interface between the protective layer and the sample solution. If a strong positive or negative potential occurs at the interface between the protective layer and the sample solution, some of the charged or strongly dipole-charged fouling substances may be able to permeate through the bound water present at the interface between the protective layer and the sample solution, thereby promoting fouling. For example, a protective layer having a net negative charge at the interface between the protective layer and the sample can be used, and the voltage applied during measurement is continuously applied so that the net surface potential at the interface between the protective layer and the sample is within at least one of ±5 mV, ±10 mV, ±20 mV, ±40 mV, or ±80 mV of the zero-volt surface potential. In the case of methylene blue, which requires a negative voltage for scanning, the mixed-charge monolayer will have a positive net charge. In the case of Os(II / III) or ferrocene redox tags, which require a positive or near-neutral voltage for scanning, the mixed-charge monolayer will have a negative or near-neutral net charge, respectively. Other combinations are possible, such as redox tags operating near zero volts with respect to the redox tag current, and thus a mixed monolayer with a zero or near-zero net charge. While voltage scans may be incompatible with fouling reduction, keeping the duration short can minimize the resulting fouling. For example, consider a monolayer with a near-neutral or neutral net charge and an Os(II / III) or ferrocene redox tag scanned with a positive voltage, which can attract many foulants to or into the mixed-charge monolayer. Instead of using a full square-wave voltammetric scan spanning hundreds of millivolts, which can easily take a minute to complete, one can scan only at potentials near the redox peak potential. For example, with a 300 Hz scan (3.3 ms total for the negative and positive trending pulses in FIG. 6A), a total of five current samples can be taken in a total scan time of only 16.5 ms.If measurements are taken every 5 minutes using this method, the voltage duty cycle is 16.5E-3 / (5. * 60) = 0.0055%, which is a very short duty cycle, and the scan is performed for that very short period during which fouling can increase. Furthermore, once the surface is fouled, the applied voltage causes the fouling substances (e.g., albumin, etc.) to reorganize into a different geometric shape, allowing for the attachment of more fouling substances, which then gradually foul the surface. Thus, the present invention may further include a measurement circuit, e.g., a wearable and programmable potentiostat, and may utilize a partial voltammogram scan method that captures only a portion of the peak redox tag current. Thus, the present invention may utilize a scan method using a scan duty cycle of at least one of less than 10%, 1%, or 0.1% during device operation.

[0063] With further reference to embodiments of the present invention relating to protective layers, a mixed-charge layer can strongly affect the freedom of movement of an aptamer and / or its binding affinity to an analyte. Some non-limiting examples are provided here. For example, in some cases, using aptamers with a mercaptohexanol protective monolayer often enables a strong sensor response because these aptamers have a net negative potential, and the mercaptohexanol interface with the sample solution has a net negative potential that repels the aptamer and redox tag. This repulsion and distance between the aptamer and redox tag are maintained until the aptamer binds to the analyte and forms a stem-loop structure (e.g., as shown in Figure 1 ) that brings the redox tag closer to the electrode. Such a structure results in a strong sensor response (change in redox tag current). If the protective monolayer were positively charged, electrostatic attraction could cause the aptamer and redox tag to be more frequently or always in close proximity to the electrode, resulting in a weaker sensor response. As another example, the electric field and Debye shielding layer may extend far enough outward from the mercaptohexanol blocking layer to inhibit, for example, binding of the analyte to the aptamer. Therefore, the charge of the aptamer and the mixed protective monolayer may be crucial to the sensor response and binding affinity of the aptamer in some cases. Therefore, the present invention can impart a net positive charge to the aptamer or to the vicinity of the redox tag on the aptamer by tagging the aptamer with a positively charged species, chemically modifying the aptamer, or incorporating a positively charged primary amine, for example, through a flexible linker. As a result, the binding affinity or sensor response of the aptamer can be optimized compared to natural DNA aptamers, which have only negative charges due to the phosphate backbone of the aptamer. Numerous combinations of the mixed-charge protective layer and aptamer charge, along with the applied voltage to the electrode, are possible, thereby optimizing sensor performance, such as binding affinity or sensor response.For example, in the case of protein sensing, the aptamer can have at least +1, +2, or +3 electron charges within 10 nucleotide units of the redox tag, such that a mixed-charge protective layer with a net negative potential at the sample interface will have a strong redox tag current (so that the redox tag is electrostatically attracted closer to the electrode). Upon protein binding, the aptamer will most likely need to change its shape to accommodate the larger protein, but in most cases, the aptamer will need to displace the redox tag away from the electrode, resulting in a strong sensor response (even though the redox tag current decreases as the protein analyte concentration increases).

[0064] With further reference to embodiments of the protective layer of the present invention, in human interstitial fluid or human serum at 33°C, the Debye length is approximately 0.7 nm, which is the distance over which the applied voltage and resulting potential in the sample solution drops to 1 / e of the potential at the surface (assuming the voltage reaches the sample solution instantaneously without voltage loss). As a result, the redox tag cannot be too far away from the potential drop near the electrode. As a further consequence, there is an upper limit to the thickness of the mixed-charge monolayer, for several reasons: (1) if the mixed-charge monolayer is highly electrically insulating and too thick, excessive applied voltage will drop too far across the monolayer rather than into the sample solution (where the redox tag is located); and (2) if the mixed-charge monolayer is partially electrically conductive (e.g., charge can penetrate the monolayer within the measurement time scale (e.g., microseconds to milliseconds)), the applied voltage will also drop too far across the monolayer rather than into the sample solution (where the redox tag is located). This is disadvantageous because it prevents the use of long-chain molecular brushes and other antifouling layers of several nanometers in thickness, which have excellent antifouling properties. However, in the present invention, it is possible to incorporate these antifouling chemistries with the charge-mixed monolayer at a density such that the freedom of movement of the aptamer is not hindered and the redox tag is not further removed from the electrode. Such a density is typically in the range of 1013 pieces / cm 2 Less than 10 12 pieces / cm 2 Less than 10 11 pieces / cm 2 Less than or 10 10 pieces / cm 2 Examples of suitable molecular brushes include, but are not limited to, at least one of: (a) a molecular brush having a length of less than 1 nm; (b) a molecular brush having a length of less than 1 nm; and (c) a molecular brush having a length of less than 1 nm. As illustrated in FIG. 7A, these long-chain molecular brushes, e.g., polyethylene glycol (PEG) or polyacrylamide, can be longer than 1 nm in length and thus can contain mixed charges, allowing them to repel very large fouling substances without the need for a protective membrane, such as a 100 kDa membrane. Furthermore, as illustrated in FIG. 7B, the aptamer-proximal portion of such long-chain molecular brushes can be negatively charged to extend the Debye length, as in the case of negative-voltage redox tags, e.g., methylene blue, and / or a negative surface potential at the protective layer interface with the sample liquid. Thus, the present invention can further include multiple molecular brush molecules interspersed between the aptamers. Thus, the present invention may further include a plurality of molecular brush molecules interspersed between the aptamers, the plurality of molecular brush molecules having a net charge in the vicinity of the aptamer and extending in length from the boundary layer with the mixed-charge monolayer by at least two Debye lengths.

[0065] With further reference to embodiments of the present invention, sensors can be constructed with mixed-charge antifouling capabilities that limit cell deposition and fibrinogen growth on the sensor. Such embodiments may further include molecular brushes as taught herein. Such embodiments may include a zwitterionic crosslinked membrane film, such as polybetaine, suspended above the sensor to separate it from the biological fluid, or layered on the sensor at a density that still allows freedom of aptamer movement and therefore proper sensor response. For wearable glucose monitors, the lag time between blood and glucose (greater than 10 minutes) is limited by the foreign body response layer and tissue deposits on the sensor membrane, which can be several cell layers thick. While very small analytes, such as glucose, on the order of 180 Da, can find a tortuous path and penetrate this deposited tissue layer on the sensor, analytes larger than 1 kDa, particularly proteins, would have much more difficulty passing through the deposited tissue layer to reach the sensor surface. Embodiments of the present invention can be tested for two use cases: insertion into the skin or into a blood vessel or vein (both of which have been demonstrated in the art for short-term use of aptamer sensors, by way of example). Embodiments of the present invention can be tested in these use case scenarios, including a surface located between the sample fluid and the sensor, and molecules and materials carried on that surface, where there is at least less than 10% or less than 30% area coverage with foreign body response-eliciting cells over at least one of 3, 7, or 14 days. This cell coverage can be easily tested by removing the sensor after in vivo use and examining the sensor surface via staining for cells or fibrinogen or other substances that make up the deposited tissue layer.

[0066] With further reference to embodiments of the present invention relating to protective layers, the present invention can also be deployed as an impedance aptamer sensor measured using electrical impedance measurements when directly exposed to human interstitial fluid or human serum at 33°C as a testable alternative. While these sensors have struggled to perform in vivo due to nonselective binding and fouling, the present invention remedies this limitation, enabling their in vivo operation. Thus, the present invention can include electrical impedance measurements of the sensor. The present invention is similar to commercially available bio-layer interferometry devices, but the device architecture, using the advanced anti-fouling capabilities taught herein, can similarly enable bio-layer interferometry in which the aptamer forms a layer that undergoes a refractive index change upon binding with the analyte, and the aptamer is suspended at the end of an optical fiber.

[0067] With further reference to embodiments of the protective layer of the present invention, a mixed-charge monolayer can be engineered for optimal compatibility with molecular pendulum motion when directly exposed to human interstitial fluid or human serum at 33°C as a testable surrogate. Here, the analyte is a large protein, and the aptamer, upon binding to the protein, has a slowed-down rate at which the voltage on the electrode can repel or attract the negatively charged aptamer and protein. The slowed-down timescale is measurable as a slowed-down change in the increase or decrease in redox tag current as the aptamer or aptamer and linked protein are electrostatically attracted to or repelled from the electrode, respectively.

[0068] With further reference to embodiments of the protective layer of the present invention, when directly exposed to human interstitial fluid or human serum at 33°C as testable surrogates, the mixed-charge monolayer may enable an aptamer sensor to have an analyte binding affinity of at least one of ±10%, 50%, 100%, 200%, 300%, 500%, or 1000% of the analyte binding affinity for the same aptamer when tested in the same manner except for using mercaptohexanol or mercaptooctanol instead of the mixed-charge monolayer. For example, if a vancomycin aptamer sensor has a binding affinity for vancomycin of 50 μM when introduced with a mercaptohexanol monolayer, a mixed-charge monolayer of the present invention can have its surface charge optimized (as taught herein) to maintain a binding affinity of at least 45 μM but no more than 55 μM. [Example]

[0069] Example 1 The working electrodes were first prepared as follows. Briefly, gold electrodes were polished for 1 minute in 0.3 μm and 0.05 μm alumina slurries using a rotary disk polisher. They were then sonicated in ethanol for 5 minutes, rinsed in deionized water, and sonicated again for another 5 minutes in deionized water. After sonication, they were briefly rinsed again in deionized water. The electrodes were then electrochemically cleaned in 0.5 M NaOH by cyclic voltammetry between -1 and -1.6 V vs. a standard Ag / AgCl reference electrode for 1400 cycles, and then in 0.5 M H2SO4 by cyclic voltammetry between 0 and +1.6 V vs. a standard Ag / AgCl reference electrode for 300 cycles. After electrochemical cleaning, the electrodes were electrochemically roughened using the NaOH roughening technique. Briefly, the electrodes were placed in 5 M NaOH and investigated by chronoamperometry between +1.32 and -4.48 V vs. a mercury oxide reference electrode. After electrochemical roughening, the electrode was again electrochemically cleaned in 0.5 M H2SO4 and rinsed before deposition of the aptamer.

[0070] To layer the aptamer on the electrodes, 400 nM cortisol aptamer was prepared in HEPES buffer containing 20 mM HEPES, 1 M NaCl, 10 mM MgCl2, and 5 mM KCl, and the pH was adjusted to 7.5 with HCl. 20 μL of aptamer was drop-cast onto each electrode for 1 h and placed in a humidity-controlled, light-proof box. After 1 h, a charge-mixed monolayer solution was prepared, and each electrode was placed in 500 μL of the charge-mixed monolayer solution overnight. Four electrodes were placed in an equimolar solution consisting of 5 mM 8-mercaptooctanoic acid and 5 mM 8-amino-1-octanethiol hydrochloride in 1x PBS. All chemicals were purchased from Sigma-Aldrich.

[0071] After incubation in the mixed-charge monolayer, the electrodes were thoroughly rinsed with deionized water and scanned using square-wave voltammetry (in 1x PBS) 100 times per electrode to remove any physisorbed species and prepare the electrode surface for better stabilization before testing. The electrodes were then placed in serum at 33°C and stabilized for 3 hours using intermittent square-wave voltammetry (SWV) scans every 15 minutes. After stabilization, the electrodes were titrated with various concentrations of cortisol to examine the current response of cortisol over various frequencies. After titrating to the desired concentrations, the electrodes were placed in fresh serum and scanned using SWV every 15 minutes for 4, 7, 12, and 14 days. On each of these days, the electrodes were titrated and returned to fresh serum, and scanning continued for the following days. The resulting titration results are plotted in Figure 8 as a plot of sensor response versus total cortisol concentration. In serum, approximately 95% of cortisol is protein-bound, so the concentration measured by the sensor is approximately 5% of the total added concentration displayed on the x-axis.

[0072] Other steps not described in detail herein but readily apparent from or incorporated with the disclosed embodiments are intended to be included as part of the present invention. The embodiments described herein provide specific examples to illustrate elements of the invention, but do not necessarily encompass all possible embodiments known to those skilled in the art. [Explanation of symbols]

[0073] 100 Sensor Devices, Devices 120 working electrode 124 Aptamers 126 Protective layer 130 Sample liquid 170 Redox Tag 180 Analyte 186 Small molecules 188 Protein 220 electrode 226 Single Layer, Elements 230 Sample liquid, liquid 320 electrode 322 Non-single layer, non-single layer surface, element, layer 326 Single Layer, Elements 330 Sample liquid, liquid 420 electrode 422 Non-single layer, non-single layer surface, element, layer 426 Single Layer, Elements 430 Sample liquid, liquid 520 Sample liquid 588 Polluting substances 590 Distance, Depth, Area 592 distance, area

Claims

1. 1. A device for continuously sensing at least one analyte in a sample liquid via measurement of the analyte, comprising: a plurality of aptamers that bind to the analyte, the plurality of aptamers carrying at least one tag that causes a change in at least one parameter when the analyte binds to the aptamer; a protective layer between the aptamers to protect the surface from fouling, the protective layer comprising a monolayer of molecules that form an interface with the sample liquid; at least one sensor having a surface comprising the monolayer of molecules is a mixed-charge monolayer of at least first and second molecules, the first molecules having a first charge that imparts a net negative charge at the interface with the sample liquid, and the second molecules having a second charge that imparts a net positive charge near the interface with the sample liquid; device.

2. The device of claim 1 , wherein the first charge and the second charge have a magnitude of at least one electron charge.

3. The device of claim 1 , wherein the monolayer of molecules comprises repeating hydrophobic molecular chains selected from the group consisting of two or more methyl groups, two or more phenyl groups, and combinations thereof.

4. 10. The device of claim 1, wherein the mixed-charge monolayer comprises one or more subsets of molecules, each subset of molecules having an end, and wherein the mixed-charge monolayer has mixed charges at the end of at least one subset of molecules.

5. The device of claim 1 , wherein the mixed-charge monolayer has mixed charges located along the interior chains of at least a subset of the molecules.

6. The device of claim 1 , wherein the mixed-charge monolayer has mixed charges located both at the ends of at least a subset of the molecules and along the interior chains of at least a subset of the molecules.

7. The protective layer has a resistance of 200 μF / cm ２ Less than 20 μF / cm ２ Less than 10 μF / cm ２ Less than 5 μF / cm ２ Less than 3 μF / cm ２ and less than 1 μF / cm ２ 10. The device of claim 1, having a capacitance that is a value selected from the group consisting of less than 1000 kJ / cm.

8. 8. The device of claim 7, wherein the capacitance is limited by a non-monolayer blocking layer to which the monolayer is attached.

9. 10. The device of claim 8, wherein the non-monolayer blocking layer comprises a material selected from the group consisting of semiconductors, metal oxides, nitrides, carbides, and mixtures thereof.

10. The device of claim 1 , wherein the sample fluid is a biological fluid selected from the group consisting of interstitial fluid, serum, and blood.

11. The device of claim 10 further comprising an electrode carrying the protective layer.

12. 12. The device of claim 11, wherein the sensor has an initial redox tag current when placed in the sample liquid, and wherein the redox tag current decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during an initial 6 hours of operation at a temperature of at least 30°C.

13. 11. The device of claim 10, wherein the sensor has a sensor response that, when placed in the sample liquid, decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during operation for at least 3 days at a temperature of at least 30°C.

14. 11. The device of claim 10, wherein the sensor has a sensor response that, when placed in the sample liquid, decreases by an amount selected from the group consisting of less than 5%, less than 10%, and less than 30% during operation for at least 7 days at a temperature of at least 30°C.

15. 12. The device of claim 11, wherein the sensor has a measurable oxygen reduction current at −0.5 V, and the oxygen reduction current increases by an amount selected from the group consisting of less than 10%, less than 30%, and less than 100% compared to the magnitude of the background current over at least three days of operation.

16. 12. The device of claim 11, wherein the sensor has a measurable oxygen reduction current at −0.5 V, and the oxygen reduction current increases by an amount selected from the group consisting of less than 10%, less than 30%, and less than 100% compared to the magnitude of the background current over at least 7 days of operation.

17. The mixed charge monolayer has a net charge density, and the net charge density is ２ 11. The device of claim 10, wherein the net charge is less than a percentage selected from the group consisting of 50%, 25%, 5%, and 1% of a net charge equivalent to a four electron charge per ion.

18. The device of claim 10 , wherein the mixed charge monolayer has a net charge, and the net charge is negative.

19. 12. The device of claim 11, wherein the mixed-charge monolayer has a net charge, the net charge is positive, and further wherein a voltage is used to operate the sensor, the voltage being sufficiently negative so that the net potential at the interface between the sample liquid and the monolayer is negative.

20. 10. The device of claim 1, wherein most of the first charges and most of the second charges are separated by a distance selected from the group consisting of 0.5 nm, 1.0 nm, and 1.5 nm.

21. 2. The device of claim 1, wherein most of the first charges and most of the second charges are separated by a number selected from the group consisting of 5, 10, and 15 methyl groups.

22. 2. The device of claim 1, wherein most of the first charges and most of the second charges are separated by a number selected from the group consisting of 1, 2, and 3 phenyl groups.

23. 11. The device of claim 10, wherein the first charge has a first distance from the electrode while the positive charge has a second distance from the electrode, and the first distance exceeds the second distance such that the boundary between the monolayer and the sample liquid has a net negative charge.

24. The mixed-charge monolayer has a density of at least 0.5 molecules / nm ２ , 1 molecule / nm ２ , 2 molecules / nm ２ , 3 molecules / nm ２ , 4 molecules / nm ２ , and 5 molecules / nm ２ 10. The device of claim 1, having a density selected from the group consisting of:

25. 12. The device of claim 11, further comprising a background current, wherein the background current increases by less than a percentage selected from the group consisting of 5%, 10%, 20%, and 40% over a period of at least three days.

26. 12. The device of claim 11, further comprising a background current, wherein the background current increases by less than a percentage selected from the group consisting of 5%, 10%, 20%, and 40% over a period of at least seven days.

27. 10. The device of claim 1, wherein the mixed-charge monolayer molecules comprise a long hydrophobic alkyl chain length selected from the group consisting of 6 methyl groups and 8 methyl groups in length.

28. 10. The device of claim 1, wherein the mixed-charge monolayer molecules are linked to the device using a chemical structure selected from the group consisting of thiols, silanes, phosphonic acids, trichlorosilanes, trimethoxysilanes, triethoxysilanes, phosphates, alkenes, and alkynes.

29. 12. The device of claim 11, wherein the tag is a redox tag, the device has an electron transfer rate, and the electron transfer rate changes by less than a percentage selected from the group consisting of 50%, 20%, 10%, and 5% over a period of at least three days.

30. 12. The device of claim 11, wherein the tag is a redox tag, the device has an electron transfer rate, and the electron transfer rate changes by less than a percentage selected from the group consisting of 50%, 20%, 10%, and 5% over a period of at least 7 days.

31. 12. The device of claim 11, wherein the boundary between the protective layer and the sample liquid is within a range of values ​​selected from the group consisting of ±5 mV, ±10 mV, ±20 mV, ±40 mV, and ±80 mV of a zero volt surface potential.

32. 32. The device of claim 31, wherein the interface between the protective layer and the sample liquid has a net positive charge when no voltage is applied to the electrodes, and a negative voltage is used during operation of the electrodes.

33. 32. The device of claim 31, wherein the interface between the protective layer and the sample liquid has a net negative charge when no voltage is applied to the electrodes, and a positive voltage is used during operation of the electrodes.

34. The device of claim 1 , further comprising a plurality of molecular brush molecules interspersed between the aptamers.

35. The device of claim 34, wherein the boundary layer and the mixed-charge monolayer have a Debye length, and further wherein the molecular brush has a net charge in the vicinity of the aptamer and extends in length at least twice the Debye length from the boundary layer with the mixed-charge monolayer.

36. The device of claim 10, wherein the sample fluid is interstitial fluid within the human body.

37. 37. The device of claim 36, further comprising a first surface area between the sample liquid and the sensor, wherein there is less than 30% area coverage of foreign body response-eliciting cells on the first surface area for a period selected from the group consisting of 3 days, 7 days, and 14 days.

38. 37. The device of claim 36, further comprising a first surface area between the sample liquid and the sensor, wherein there is less than 10% area coverage of foreign body response-eliciting cells on the first surface area for a period selected from the group consisting of 3 days, 7 days, and 14 days.

39. The device of claim 1 , wherein the at least one tag is a redox tag and the parameter is a redox current.

40. The device of claim 1 , wherein the parameter is electrical impedance.

41. The device of claim 1 , wherein the parameter is light intensity.

42. 2. The device of claim 1, further comprising a binding affinity between the analyte and the plurality of aptamers, wherein the binding affinity is within at least one of ±10%, 50%, 100%, 200%, 300%, 500%, and 1000% of the analyte binding affinity for the same aptamer when tested in the same manner except for using mercaptohexanol or mercaptooctanol instead of the charge-mixed monolayer.

43. A method for sensing at least one analyte in a sample liquid through measurement of the analyte, comprising exposing the sample liquid to the sensor described in claim 1 and using data regarding changes in one or more aptamer parameters to measure the analyte, wherein a voltage is applied to the electrodes so that the potential at the interface between the monolayer and the sample liquid has a net negative potential and most fouling substances and aptamers are slightly repelled from the surface.

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