Electrochemical sensors and methods cross-reference to related applications

Abstract

Sensors described in this disclosure have unexpectedly improved stability in challenging environments, including whole blood. The increased stability is at least partly result of a polymer coating that covers a nanoporous layer. The polymer coating allows analytes to travel through to recognition elements within a nanopore in the nanoporous layer. The polymer coating and the nanopore reduce or eliminate the travel of interferents to degrade or foul the recognition elements. The nanoporous layer may define a nanopore or plurality of nanopores. The recognition element may be in contact with the nanoporous layer and disposed in the nanopore. The recognition element may be configured to bind to the analyte to form a complex that sends a signal. Systems, methods of making the systems, and methods of using the systems are described.

Description

PATENT Attorney Docket No.079445-1473006-013310PC Client Ref. No. S23-420 ELECTROCHEMICAL SENSORS AND METHODS CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to U.S. Provisional Application No.63 / 610,584, filed on December 15, 2023, the disclosure of which is incorporated by reference in its entirety for all purposes. BACKGROUND

[0002] Blood tests are an important component of clinical care and health monitoring, but most blood tests are designed to provide instantaneous measurements for a certain point in time. Compact or wearable sensors that enable continuous measurement of circulating analytes could greatly extend the utility of blood-based monitoring by tracking disease states, drug dosing, or metabolic and physiological activity in a wide range of clinical and non- clinical setting. Aptamer-switch-based electrochemical biosensors and other suitable sensors can provide a generalizable molecular detection framework that achieves good sensitivity and specificity for relatively short-term continuous in vitro and in vivo measurements of a variety of small-molecule analytes. In the aptamer-switch design, analyte binding induces a conformational change in a redox-tagged, electrolyte-coupled aptamer strand; this in turn alters the relative proximity of the redox tag to the electrode surface, producing a measurable change in current. The subsequent dissociation of the target then allows the aptamer switch to “reset”, restoring baseline current and offering the potential for continuous, real-time detection. Electrochemical aptamer-based sensor platforms can deliver a sensitive, rapid, and highly specific response to changes in analyte concentration. However, none of the systems are suitable for extended use in the complex and challenging environment of the bloodstream, and the longest duration demonstrated for such an in vivo sensor to date is less than half day, which cannot match the clinical and daily-use requirement of health monitoring. Extended use of biosensors and other challenges are addressed with systems and methods described herein.BRIEF SUMMARY

[0003] Sensors described in this disclosure have unexpectedly improved stability in challenging environments, including whole blood. The increased stability is at least partly result of a polymer coating that covers a nanoporous layer. The polymer coating allows analytes to travel through to recognition elements within a nanopore in the nanoporous layer. The polymer coating and the nanopore reduce or eliminate the travel of interferents to degrade or foul the recognition elements.

[0004] In embodiments, a sensor may include a sensing surface. The sensing surface may include a nanoporous layer defining a plurality of nanopores. The nanoporous layer may be functionalized with a plurality of recognition elements in contact with the nanoporous layer and disposed in the plurality of nanopores. Each recognition element of the plurality of recognition elements may be configured such that upon binding of the recognition element to the analyte, a detectable signal from the recognition element is generated. The sensor may in addition include a polymer coating overlaying the nanoporous layer and the recognition elements disposed thereon. The nanopore and the polymer coating may be configured to reduce an amount of interferents from contacting the recognition elements.

[0005] In embodiments, methods may include functionalizing a nanoporous layer with a plurality of recognition elements. Methods may in addition include depositing a polymer coating over the nanoporous layer. Other embodiments of this aspect include corresponding computer systems, apparatus, and computer programs recorded on one or more computer storage devices, each configured to perform the actions of the methods.

[0006] A better understanding of the nature and advantages of embodiments of the present invention may be gained with reference to the following detailed description and the accompanying drawings. BRIEF DESCRIPTION OF THE DRAWINGS

[0007] FIG.1A shows an example of a sensor device according to embodiments of the present invention.

[0008] FIG.1B shows how different coatings are screened according to embodiments of the present invention.

[0009] FIG.1C is a graph that characterizes the extent of degradation in real time by electrochemically measuring signal loss over time according to embodiments of the present invention.

[0010] FIG.1D is a graph of the signal gain after a 30-minute exposure to S1 nuclease for different aptamer-based biosensors with different polyethylene glycol (PEG) and for planar gold and nanoporous gold according to embodiments of the present invention.

[0011] FIG.2A is a graph of the square wave voltammetry (SWV) assessment of electrochemical drift after exposing a PEG-free pAu (MCH-only) electrode to undiluted human serum for a month.

[0012] FIG.2B is a graph of the square wave voltammetry (SWV) assessment of electrochemical drift after exposing a sensor electrode to undiluted human serum for a month according to embodiments of the present invention.

[0013] FIG.2C is a graph of drift from baseline electrochemical signal over the course of a month-long incubation in human serum for a nanoporous sensor versus the pAu sensor according to embodiments of the present invention.

[0014] FIG.2D shows calibration curves obtained at multiple timepoints during a month- long incubation in undiluted human serum for pAu.

[0015] FIG.2E shows calibration curves obtained at multiple timepoints during a month- long incubation in undiluted human serum for the long-term Stable Electrochemical Nanostructured Sensor for Blood In situ Tracking (SENSBIT) according to embodiments of the present invention.

[0016] FIG.2F shows a graph of the drift from fitted maximum signal gain of the sensor, B_max (R2 > 0.80) over the course of a month-long incubation in human serum for SENSBIT according to embodiments of the present invention versus the pAu sensor. The shaded area indicates the 95% confidence intervals of Langmuir isotherm model fitting (n = 4).

[0017] FIG.3A shows top-view SEM images of PEG-free pAu control and SENSBIT sensor electrodes after 14 h incubation with fresh human blood according to embodiments of the present invention.

[0018] FIG.3B is a graph that shows the extent of blood cell fouling for PEG-free and PEG-coated pAu and npAu sensors according to embodiments of the present invention.

[0019] FIG.3C shows in situ focused ion beam scanning electron microscopy (FIB-SEM) cross-section image of a blood cell attached to a sensor according to embodiments of the present invention.

[0020] FIG.3D shows SEM image of SENSBIT electrodes after incubation for 19 days in 0.2-μm filtered chicken plasma according to embodiments of the present invention.

[0021] FIG.3E shows fluorescence micrographs of sensor electrodes after 2 h incubation with fresh human blood containing fluorescently-labeled fibrinogen according to embodiments of the present invention.

[0022] FIG.3F is a graph of fibrin coverage for PEG-free and PEG-coated pAu and npAu electrodes according to embodiments of the present invention.

[0023] FIG.4A illustrates the workflow of intravenous real-time continuous monitoring according to embodiments of the present invention.

[0024] FIG.4B is an optical image of an intravenous probe compared to 22G clinical catheter according to embodiments of the present invention.

[0025] FIG.4C illustrates the intravenous probe implanted within femoral vein to achieve a direct interface with the blood in anesthetized rats according to embodiments of the present invention.

[0026] FIG.4D is a graph of real-time SWV scanning with intravenously implanted pAu over several hours after the injection of a single bolus of drug (time marked with dashed line) according to embodiments of the present invention.

[0027] FIG.4E shows real-time tracking of kanamycin concentration with SENSBIT over several hours after the injection of a single bolus of drug (time marked with dashed line) without the need for signal correction according to embodiments of the present invention.

[0028] FIG.5A shows the workflow for evaluation of long-term implantation in free- moving rats according to embodiments of the present invention.

[0029] FIG.5B shows signal degradation after two days of continuous implantation of a PEG-free pAu sensor.

[0030] FIG.5C shows signal degradation after two days of continuous implantation of a SENSBIT electrode according to embodiments of the present invention.

[0031] FIG.5D shows baseline electrochemical signals of the pAu sensors and SENSBIT sensors after multi-day implantation according to embodiments of the present invention.

[0032] FIG.5E is a graph summarizing papers on device longevity of existing aptamer- based electrochemical sensors interfacing blood derivatives and whole blood according to embodiments of the present invention.

[0033] FIG.5F shows calibration curves of pAu before intravenous implantation and after 1 day implant according to embodiments of the present invention.

[0034] FIG.5G shows continuous SWV scanning to extract real-time kanamycin concentrations in buffer with pAu before and after 1 day implantation according to embodiments of the present invention.

[0035] FIG.5H shows calibration curves of SENSBIT after one day or four days of chronic implantation according to embodiments of the present invention.

[0036] FIG.5I shows continuous kanamycin monitoring with SENSBIT before and after 1- and 4-day implantation in femoral veins of free-moving rats according to embodiments of the present invention.

[0037] FIG.6 shows an example of a microfabricated sensor according to embodiments of the present invention.

[0038] FIG.7 shows connections between the sensor and a computer configured to record the electrical signal according to embodiments of the present invention.

[0039] FIG.8 is a schematic illustration of the fabrication process of a nanoporous gold sensor according to embodiments of the present invention.

[0040] FIG.9 is a flowchart of an example process of making a sensor according to embodiments of the present invention.

[0041] FIG.10 is a flowchart of an example process for detecting an analyte according to embodiments of the present invention.

[0042] FIG.11 shows a fabrication workflow according to embodiments of the present invention.

[0043] FIG.12 shows a graph of intravenous real-time monitoring of kanamycin pharmacokinetics during anesthesia according to embodiments of the present invention. DETAILED DESCRIPTION

[0044] There are no compact or wearable sensors that continuously measure circulating analytes in blood over an extended amount of time. The failure to achieve longer-term continuous sensing is attributable to the inability to maintain sensing functionality when exposed to complex biological matrices. First, any biofluid-interfacing system must be capable of resisting surface fouling by proteins, cells, and other biomolecules in the sample matrix. Such fouling will limit the longevity time that systems for routine medical procedures, such as dialysis and cardiopulmonary bypass, can be used to just a few hours. Additionally, other biochemical sensors rely on antibodies, enzymes, or aptamers that are vulnerable to chemical and enzymatic degradation as well as non-specific binding to interferents. Other strategies have been developed to overcome these challenges, including the use of passive filtration systems that protect the sensor and its associated bioreceptors from fouling and degradation. For example, a microfluidic system with two stacked laminar flows can be leveraged to dynamically filter out blood cells and other high-molecular-weight interferents in flowing blood prior to aptamer-based sensing. This approach suffers from measurement time lag due to blood passing through the device as well as mechanical vulnerability in terms of maintaining laminar flow conditions. One can also employ surface modification with anti-fouling self-assembled monolayers (SAMs), but gradual monolayer desorption due to thermal, electrical, or fouling effects erodes the effectiveness of such measures, ultimately leading to device failure. Nanostructured biosensors have also exhibited the potential to suppress non-specific protein binding by limiting plasma protein access to the sensor surface, and nanoporous electrode designs can protect immobilized aptamers against mechanical shearing during implantation. Nevertheless, complex interactions with the surrounding biological milieu have generally limited the operational duration of such electrodes to less than three hours in vivo. As such, there is still no reliable and generalizable framework for achieving stable and accurate biomolecular detection of a broad range of analytes in vivo over protracted periods of time.

[0045] Described herein is a biomimetic multi-component sensor design that enables continuous analyte detection with an electrochemical aptamer-switch-based sensor. The Stable Electrochemical Nanostructured Sensor for Blood In situ Tracking (SENSBIT) systemretains excellent sensitivity and stability for detecting the antibiotic kanamycin even after protracted exposure to tissue environments and complex biological fluids including blood and serum. As an example, SENSBIT features a 3D nanoporous gold surface that sequesters electrochemically-modified aptamer switch “receptors” from interferents in the sample matrix. An additional coating of highly-branched polymer molecules on the porous gold surface confers additional and unexpectedly effective protection, insulating the sensor against degradation and fouling. When challenged with undiluted serum in vitro, SENSBIT can, in some implementations, maintain a stable baseline and consistent sensitivity for at least a month.

[0046] In one demonstration of the novel technologies disclosed herein, as set forth in the Examples section below, the SENSBIT is implanted in rat femoral veins and achieved real- time pharmacokinetic monitoring of kanamycin in the bloodstream with no need for signal correction over hours of operation. Additionally, longer-term intravenous implantation was performed using the sensor into free-moving rats and showed that SENSBIT maintains >75% baseline stability and robust sensitivity for kanamycin detection after four days—far longer than any other published in vivo sensor described to date. SENSBIT’s resilience and stability even after prolonged exposure to complex biological fluids and tissue environments suggests that it could offer a robust and generalizable platform for the development of electrochemical aptamer-based sensors that are compatible with long-term in vivo implantation and sensing of blood biomarkers.

[0047] In this disclosure, sensors that are called SENSBIT sensors may be used for applications that are not necessarily in blood. SENSBIT sensors refer to the hierarchical configuration of the sensor and do not specify the use of the sensor unless context clearly dictates otherwise. I. DESIGN CONCEPT FOR SENSOR SYSTEM

[0048] SENSBIT integrates multiple technological advances to overcome the challenges that have previously prevented development of blood-interfacing biosensors for long-term in vivo molecular monitoring. Examples of biosensors include materials acting as a mechanical and physicochemical barrier against shear-induced damage and harmful molecules, while also serving as a frontline sensor for the real-time acquisition of bacteria-derived immunomodulatory signals. The sensor surface itself consists of a 3D bicontinuous nanoporous surface, which is modified with aptamer switch receptors as well as ahyperbranched polymer coating. In the context of electrochemical sensing, SENSBIT’s redox-tagged aptamer switches undergo reversible binding-induced molecular switching to generate a measurable electrochemical readout in the presence of their target analyte. Target concentration can be calculated from the signal gain of the electrochemical signals, which can be quantified with the redox-tag current peaks from square-wave voltammetry (SWV), enabling real-time molecular quantitation.

[0049] In some examples, SENSBIT can permit long-term in vivo analyte monitoring in complex biological environments, such as undiluted plasma and whole blood, without sacrificing performance to signal degradation or excess background. As a proof-of-concept demonstration, an exemplary aptamer-based electrochemical SENSBIT system was developed that tracks changing concentrations of the small-molecule antibiotic drug kanamycin. This is a clinically relevant example, as kanamycin-based therapy to treat serious bacterial infections usually lasts weeks, and the dose could theoretically be tuned and optimized in real-time based on data from a continuous sensor. Additionally, there is an existing electrochemical aptamer switch that has been used to sensitively and specifically report concentrations of this analyte in other sensor systems, binding to kanamycin with an equilibrium dissociation constant KD of ~80 nM. II. DETECTION SYSTEMS A. Sensor Systems

[0050] The scope of the disclosure encompasses various analyte detection systems, also referred to herein as “sensor systems,” in which the sensor systems are constituted to detect a selected analyte and in which the sensor configurations provide protection from fouling and other factors that confound effective detection of the target species. The sensor configurations include a novel combination of pores and polymer coatings, which elements, in combination, provide unexpectedly and massively superior performance over previous sensor designs. The sensors, by this superior performance, may be deployed for long-term operation in challenging sampling environments, for example, in live animals, enabling novel diagnostic, therapeutic, and research applications that were not achievable by other systems.

[0051] The protective elements of the invention may be applied to all manner of extant sensor platforms in which the sensing functions of the device are improved by anti-fouling protection. The improved sensor designs of the present disclosure are not limited to anyparticular sensor type and may be advantageously applied to improve the function of any sensor platform. For example, the protective elements of the invention may be applied to electrochemical sensing platforms such as aptamer-based electrochemical sensors. The protective elements may further be applied to optical sensor platforms and other sensor types. The protective elements may also be deployed on enzymatic-based sensors.

[0052] A first element of the sensing system is a sensor element. The sensor element includes a sensing surface, which, in operation will be exposed to a sample fluid and, by the action of recognition elements thereon, will provide a detectable signal that is responsive to the presence of a target analyte. In the example sensor systems, the sensing systems may further include the protective elements disclosed herein, which enable superior performance.

[0053] The sensor element can achieve analyte detection by a plurality of recognition elements. Recognition elements may include molecular switches, such as moieties, that selectively bind to the target analyte, and where such binding induces a phsysiochemical change in the recognition element that may be transduced to a detectable signal. For example, target binding may induce a conformational change in the recognition element molecule. The recognition elements may include any molecular switch composition of matter. Exemplary recognition elements include aptamers and other nucleic acids, antibodies and antigen- binding fragments thereof, target binding peptides, binding domains of proteins, small molecules, and other compositions having affinity for an analyte or class of analytes. Molecular switches are described, for example, in K. Fu et al., Advance Science (2021) and J- W. Seo et al., Science Advances (2022), the entire contents of which are incorporated herein by reference for all purposes.

[0054] The sensor elements, in certain embodiments, can include one or more reporter moieties, which are compositions of matter that facilitate transduction of binding-induced changes to detectable signals. For example, in the context of electrochemical sensors, the one or more reporter moiety may include a redox probe. In the context of optical sensors, the reporter moiety may include a fluorescent tag.

[0055] Electrochemical Sensors. In some implementations, the sensor system may include an electrochemical sensor. Embodiments may include any type of electrochemical sensor in which the sensing electrode is functionalized with the protective elements described herein. A typical electrochemical sensing system includes a reference electrode and a sensing electrode including a sensing surface in which the sensing electrode surface is functionalized withrecognition elements. Potentials applied to the electrodes may result in faradaic current flows between the electrodes in which analyte binding to recognition elements alters their properties in such a way as to modulate the current between the electrodes, for example, in a concentration-dependent manner. For example, the oxidation-reduction potential of the sensing electrode can be modulated by the binding of analytes to the recognition elements.

[0056] The sensing electrode surface may include any number of suitable conductive materials, such as metals including gold, titanium, copper, aluminum, tungsten, platinum, and palladium, conductive ceramics including indium tin oxide (ITO), conducting polymers including poly (3,4-ethylenedioxythiophene):poly (styrene sulfonate) (PEDOT:PSS) and polyaniline (PANi) and other carbon-based or silicon-based conductors. The sensing electrode surface may include a solid object composed of the foregoing or may include a support material coated with electrode material, or onto which the electrode material is deposited. Support materials may be or include any material, such as polymeric materials, glass, carbon, etc.

[0057] In some implementations, the electrochemical sensor may include an aptamer-based electrochemical sensor. In this context, the sensing element includes a sensing electrode in which the sensing electrode is functionalized with recognition elements including molecular switch aptamers. The molecular switch aptamers may include nucleic acid sequences that selectively and reversibly bind a target species in which such binding changes the conformation of the aptamer. In the context of an aptamer-based electrochemical sensor, the aptamer may be immobilized on the sensing electrode surface, for example by thiol conjugation chemistries known in the art for joining nucleic acids to metals such as gold. The aptamer may further be functionalized with one or more reporter moieties where the conformational change induced by target binding will alter the position of the reporter moiety with respect to the electrode surface, which modulates currents when probative potentials are applied to the pair of electrodes in the sensing system. In the context of electrochemical sensors, the reporter moiety may include a redox probe, for example, methylene blue, cytochrome c, ferrocene, anthraquinone, plastocyanin, viologen, ethidium bromide, daunomycin, ruthenium, bis-pyridine, tris-pyridine, bis-imidizole, and ethylenetetracetic acid. The reporter moiety or moieties may be conjugated to the aptamer by suitable chemistries, for example by N-hydroxysuccinimide ester modified probe conjugation to an amine-terminated aptamer.

[0058] Over 1,000 aptamers to various targets are known in the art, for example as cataloged at the University of Texas Aptamer Database, accessible at sites.utexas.edu / aptamerdatabase. Novel aptamers are being developed by systematic evolution of ligands by exponential enrichment (SELEX) and other aptamer evolution technologies known in the art. The aptamer may include any nucleic acid composition, for example, DNA, RNA, or an aptamer including non-natural nucleic acids, as well as hybrids of the foregoing. Typical aptamers are about 15-60 bases in length.

[0059] The sensing systems may include any aptamer-based sensing architecture known in the art, for example, as described in Xiao, Y, Rowe, A. A., and Plaxco, K. W. (2007) Electrochemical detection of parts per billion lead via an electrode-bound DNAzyme assembly. J Am. Chem. Soc.129, 262-263; U.S. Pat. No.8,003,374 by Heeger, Fan, and Plaxco; Ferguson et al., “Real-time, aptamer-based tracking of circulating therapeutic agents in living animals,” Sci Transl Med.2013 Nov.27; 5(213): 213ra165; and Swensen et al., “Continuous, Real-Time Monitoring of Cocaine in Undiluted Blood Serum via a Microfluidic, Electrochemical Aptamer-Based Sensor,” J Am Chem Soc.2009 Apr.1; 131(12): 4262-4266, the entire contents of which are disclosed herein by reference for all purposes. The aptamer-based electrochemical sensor may include a signal-on system, a signal-off system, or other suitable configuration.

[0060] In some implementations, the recognition element includes an antibody-based molecular switch, such as target-responsive “antibody-switches” described in I. Thompson et al., Science Advances (2023), monolithic dual-antibody clamp described in N. Maganzini et al., biorxiv (2023), and molecular pendulum described in J. Das et al., Nature Chemistry (2021), and A. Mahmud et al., Journal Angewandte Chemie - International Edition (2023), the entire contents of which are incorporated herein by reference for all purposes. The antibody-based molecular switches, while immobilized on the electrode surface, can dynamically bind and release target molecules according to the real-time environmental target concentrations, coupling with conformational changes. The reporter moiety or moieties to transduce the conformational changes may include electrochemical tags, such as, methylene blue or ferrocene.

[0061] Electrochemical sensors may be operated by suitable devices, including potentiostats for the application of controlled potentials in selected waveforms and elements for measuring and recording the resulting faradaic currents.

[0062] Electrochemical sensors may be interrogated by suitable methods and as appropriate for the selected sensor configuration. In some embodiments, the sensors are configured for pulsed voltametric measurements where a series of potentials is applied in a pulsed waveform to the functionalized sensing electrode of the sensor. One or more measurements of faradic current may be assessed in each cycle. In some embodiments, the sensor is interrogated by square-wave voltammetry (SWV).

[0063] In some embodiments, the nanoporous layer may be in electrical connection with elements for electrochemical interrogation of the redox state of the nanoporous layer. Elements may include a probe to deliver signal from the nanoporous layer to a detector.

[0064] Optical Sensors. In some implementations, the sensor system includes an optical sensor system. In optical sensor system implementations, the sensor may include a set of recognition elements conjugated to or otherwise present on a sensing surface in which the recognition elements are functionalized with one or more optical reporter moieties such as fluorophore-quencher pairs. Analyte binding to the recognition element induces a change in conformation that alters an optical property or output of the sensor element. In some implementations, the recognition element is a molecular beacon. Additionally or alternatively, the recognition element is an aptamer. Exemplary reporter moieties include: Förster Resonance Energy Transfer (FRET) pairs, for example, including donors-acceptor pairs such as fluorescein and DABCYL; chemi / bio-luminescent reporters, such as the luminol-H2O2-horseradish peroxidase-fluorescein system; and fluorescent dyes utilized in Plasmon field-enhanced fluorescence systems.

[0065] In the context of optical sensors, the sensing surface may include various materials, such as gold, silver, platinum, copper, silica, graphene, graphene-oxide, molybdenum disulfide nanosheets, indium tin oxide (ITO), and / or cobalt oxyhydroxide nanosheets. The sensing surfaces may be functionalized with the protective elements described herein, for example, to enhance performance in complex sample environments.

[0066] Optical sensors may be interrogated by suitable elements such as light sources for excitation of fluorophores, such as laser diodes, and fluorescence microscopy systems for measuring optical emissions from sensing surfaces.

[0067] Nanoporous Sensing Surface. Embodiments may include any nanoporous sensing surface such as a sensing surface including a set of nanopores. The sensor may include a set of nanopores. The nanopores are characterized by a bicontinuous network of nanometer-sizedstruts and interconnected pores. The sensing surface may include any material, for example, gold, silver, platinum, titanium, copper, titanium oxide, silica, ceramic material, graphene, graphene-oxide, molybdenum disulfide nanosheets, indium tin oxide (ITO), and / or cobalt oxyhydroxide nanosheets and other materials, for example materials commonly deployed or suitable for deployment in electrochemical and optical biosensors.

[0068] The nanoporous sensing surface may be prepared by any number of suitable methods for the creation of nanoporous structures. The sensing surface may include a nanoporous layer. In some embodiments, the nanoporous surface may be formed by a co- deposition and dealloying process. For example, in the case of a sensor element including a gold electrode, the nanoporous surface may be fabricated by co-sputtering of gold and silver, for example, in an Au:Ag ratio ranging from 5:1 to 1:1, for example about 2:1. Other co- deposition methods for generating nanoporous surface include co-electroplating metals with mixed electrolytes, and sol-gel deposition of mixed precursors. The deposited layer can be dealloyed by immersion in strong acid, for example nitric acid at about 69-70%, to selectively remove most of the non-resistant material (e.g., silver) and create a nanoporous surface. Commercially available gold alloys can also be used to dealloy for creating the nanoporous surface, for example, 10k gold. The size features of the nanoporous surface may be tuned by varying the proportion of metals in the co-sputtered alloy or adjusting the dealloying process time. Additional methods to tune the size features of the nanoporous surface include electrochemical coarsening, for example in H2SO4 at a scan rate of 50 mV / s over a potential range of 0–1.8 V, as well as thermal annealing, for example, baked at 200-500 °C for 1-10 min. Additional techniques of facilitating pore formation include surface patterning by laser ablation or electron beam lithography.

[0069] Methods of forming nanoporous electrodes include co-sputtering described in Matharu et al., 2017, Nanoporous-Gold-Based Electrode Morphology Libraries for Investigating Structure–Property Relationships in Nucleic Acid Based Electrochemical Biosensors, ACS Appl. Mater. Interfaces 9: 12959–12966; Fu et al, 2021, Accelerated Electron Transfer in Nanostructured Electrodes Improves the Sensitivity of Electrochemical Biosensors, Adv. Sci.8: 2102495; Daggumati, et al., 2015, Biofouling-resilient nanoporous gold electrodes for DNA sensing. Anal. Chem.87, 8618–8622; and Patel et al., 2013, Electrochemical properties of nanostructured porous gold electrodes in biofouling solutions. Anal. Chem.85, 11610–11618. The methods of forming nanoporous electrodes by co- electroplating metals include those described in Lang et al.2013, Nanoporous gold supportedcobalt oxide microelectrodes as high-performance electrochemical biosensors, Nat. Commun. 2169; Sattayasamitsathit et al.2012, Highly ordered tailored three-dimensional hierarchical nano / microporous gold–carbon architectures, J. Mater. Chem.,22, 11950-11956. The methods of forming nanoporous electrodes by directly etching commercial alloys were described in Bhattarai et al.2021, Nanoporous Gold Monolith for High Loading of Unmodified Doxorubicin and Sustained Co-Release of Doxorubicin-Rapamycin. Nanomaterials, 11(1), 208. The methods of 3D printing nanoporous gold were described in Zhu et al.2018, Toward digitally controlled catalyst architectures: Hierarchical nanoporous gold via 3D printing. Sci. Adv.4,eaas9459. The methods of tuning the pore sizes of the nanoporous structures described in Fu et al, 2021, Accelerated Electron Transfer in Nanostructured Electrodes Improves the Sensitivity of Electrochemical Biosensors, Adv. Sci.8: 2102495; Jeon et al, 2017, Self- similarity in the structure of coarsened nanoporous gold, Scr. Mater.46-49; Dixon et al, 2007, Preparation, Structure, and Optical Properties of Nanoporous Gold Thin Films, Langmuir, 23, 5, 2414–2422. Methods of forming metallic, ceramic, or polymeric nanoporous surface described in Koya et al., 2021, ACS Nano, Nanoporous Metals: From Plasmonic Properties to Applications in Enhanced Spectroscopy and Photocatalysis, 15, 4, 6038–6060; Jones, B. H., & Lodge, T. P.2009. High-temperature nanoporous ceramic monolith prepared from a polymeric bicontinuous microemulsion template. Journal of the American Chemical Society, 131(5), 1676-1677; Dawson, R., Cooper, A. I., & Adams, D. J. 2012. Nanoporous organic polymer networks. Progress in Polymer Science, 37(4), 530-563. The contents of the foregoing references are incorporated herein by reference for all purposes.

[0070] In some embodiments, the nanoporous layer defines a set of nanopores. Each nanopore of the set of nanopores, or any subset thereof, may have an effective diameter in a range from 1 to 5 nm, from 5 to 10 nm, from 10 to 20 nm, from 20 to 30 nm, from 30 to 50 nm, or over 50 nm. The effective diameter is the diameter of a circle having the same area as the cross section of the nanopore. The effective diameter may be a mean or median or may be a diameter taken at a certain depth of the nanopore (including at the surface, or 25% or 50% or 75% of the depth of the nanopore). Nanopores may include substantially cylindrical cavities, or may include irregular or tortuous cavities, for example, including branches. In some embodiments, the nanopores are substantially interconnected to form a continuous mesh-like network of pores. The term “substantially” may allow for some variation ordeviation that does not depart from the fundamental nature (e.g., cylindrical or interconnect). Substantially may mean at least 80%, 85%, 90%, 95%, or 99% of the feature is present.

[0071] The nanoporous layer may be a bicontinuous network of nanometer-sized solid struts and interconnected pores. The nanoporous layer may be defined by the strut width, pore diameters, and the distance between the interior lines of neighboring struts or pores, individually or collectively. The effective sizes of the strut width, pore diameters, or the distance may each be in a range from 1 to 5 nm, from 5 to 10 nm, from 10 to 20 nm, from 20 to 30 nm, from 30 to 50 nm, or over 50 nm. The tortuosity may have a tau of 1-1.5, 1.5-2, 2- 3, or greater than 3.

[0072] The nanoporous layer may overlay a body or support structure, for example, configured to define the shape of the sensor. In some embodiments, the nanoporous layer overlays a substrate including a metal, ceramic, or polymer substrate. The metal substrates may include gold, tungsten, platinum, silver, and / or copper. The polymer substrates may include polyimide (PI), polyethylene terephthalate (PET), photoresist (e.g., SU-8), cellulose, rubber and elastomer (e.g., polydimethylsiloxane, polyurethane, styrene ethylene butylene styrene block copolymers). The body or support structure may be characterized by a surface having an area of 1,000 to 5,000 μm2, 5,000 to 10,000 μm2, 10,000 to 15,000 μm2, 15,000 to 20,000 μm2, or over 20,000 μm2. The nanoporous layer may overlay and may be in contact with a gold layer. The gold layer may overlay and may be in contact with a metal layer that may include titanium and gold. The metal layer may overlay and may be in contact with the polymer substrate.

[0073] Recognition Elements. Following formation of the nanoporous sensing surface, the nanoporous sensing surface is functionalized with a set of recognition elements. In some embodiments, the recognition element includes a molecular switch or enzymes. Molecular switches are described in K. Fu et al., Advance Science News (2021) and J-W. Seo et al., Science Advances (2022), the entire contents of both of which are incorporated herein by reference for all purposes. In some implementations, the recognition element includes an aptamer, for example, an aptamer functionalized with one or more reporter moieties, for example, one or more redox probes or fluorophores. Aptamers may be conjugated to the sensing surface by appropriate chemistries. In the context of gold sensing surfaces, aptamers may be conjugated to the sensing surfaces by nucleic acid thiol conjugation chemistry as known in the art. In some embodiments, the recognition element includes an enzyme, forexample, an enzyme embedded in a matrix. In some embodiments, the enzyme is hexokinase, glucose oxidase, or glucose-1-dehydrogenase, for use in a glucose sensor. The recognition elements may be deposited at any density suitable for generating measurable signal, for example, 1010to 1011, 1011to 1012, 1012to 1013, or over 1013molecules per cm2. The set of recognition elements in the nanopore and / or sensor may be from 2 to 5, from 5 to 10, from 10 to 50, from 50 to 100, or from 100 to 1,000.

[0074] Sensor Configurations. The sensor systems may be configured for a number of uses. In some embodiments, the sensor element of the sensor system is configured for implantation or other in vivo placement. For example, the sensor element may include a shank, needle, needle array, microneedle array, disc, or other body shape, for example, including a sensing surface deposited or otherwise present on a support which defines the form of the sensor element. In some embodiments, the sensor element may include a wire or other elongated body. The sensor element may be integral to an implanted housing, for example a housing including a needle, catheter, or cannula, which may be inserted into a

[0075] Target Analytes. The sensor systems described herein may be configured for the detection of one or more analytes, the analyte being any composition of matter detectable by the selected recognition element. Exemplary analytes include, for example: ions, proteins, peptides, carbohydrates, lipids, drugs, metabolites, growth factors, nutrients, neuropeptides, gases, small molecules, cells and other compositions of matter, for example compositions including biomarkers, therapeutics, pathogens, wastes, and contaminants.B. Protective Elements

[0076] The sensor elements may include protective elements, which act to inhibit contact between the recognition elements and species present in the sample that interfere with sensor functions. The protective elements of the system may allow dynamic fluid contact between the recognition element and the sample, such that the sensor is responsive in real time to the concentration of analyte present in the sample, while at the same time, preventing, reducing, or inhibiting interferent contact with and / or activity on the recognition elements. Interferents may include cells (e.g., erythrocytes, platelets, leukocytes, fibroblasts, macrophages, glia cells, and astrocytes), proteins (e.g., fibrinogen, albumin, and other nonspecific binding proteins), nucleases, or any material that may foul or degrade the molecular switch. The protective elements of the invention include a nanoporous sensing surface functionalized with recognition elements, overlaid by a coating of polymeric material.

[0077] The protective elements may be deployed on enzymatic based sensors such as glucose-oxidase or glucose dehydrogenase-based sensing elements. In such contexts, enzymatic moieties, by reaction with target analytes, result in the generation of electrochemically active species that are detectable by an underlaying electrode. In such contexts, the recognition elements include enzymes that react with target analytes. In such implementations, the sensing surface includes enzymatic moieties (in some embodiments, embedded in a scaffold or support material) disposed on a nanoporous sensing surface overlaid with a polymer coating.

[0078] Polymeric Coating. The sensor surface also may include a polymer coating overlaying the nanoporous layer functionalized with recognition elements. The polymer coating may be polymeric coating 108 or any polymeric coating described herein. The polymer coating may be deposited, grafted, to and / or from the nanoporous surface. The polymer coating may be adhered to the sensing surface by any interactions, for example, non- covalent interactions, electrostatic interactions, covalent bonding, etc. The polymer coating may be applied in any manner, for example by spin coating, dip coating, sputtering, thermal evaporation, chemical vapor deposition, atomic layer deposition, electrochemical deposition, electron beam deposition, Langmuir-Blodgett deposition, or colloidal deposition.

[0079] The polymer coating may include a natural polymer, a synthetic polymer, mixtures of the foregoing, or a crosslinked hydrogel network. The polymer coating may include a copolymer or homopolymer. The polymer coating may include a hydrogel. The polymercoating may include a polyethylene glycol (PEG), for example, a polyacrylamide, a polysaccharide, a polyoxazoline, a polyglycerol fatty acid ester, a poly(thioglycidyl glycerol) (PTGG), a polycarbonate, a propylene glycol, a polyaminoacid, a polyvinylpyrrol (PVP), a zwitterionic polymer, a polybetaine, a polyester, or a polysarcosine. The polymer coating may include a glycoprotein or glycoprotein-glycan hybrid material.

[0080] In some embodiments, the polymer coating includes polymers of hyaluronic acid, for example, having a molecular weight between 1-2×106KDa, for example, as described in Bauer et al, 2013, Adhesion of Marine Fouling Organisms on Hydrophilic and Amphiphilic Polysaccharides, Langmuir 29: 4039–4047; polymers of alginic acid, for example, as described in Morra and Cassinelli, 2012, Non-fouling properties of polysaccharide-coated surfaces, Non-fouling properties of polysaccharide-coated surfaces, Journal of Biomaterials Science, Polymer Edition, 10:1107-1124.

[0081] In some embodiments, the polymer coating includes polyoxazoline: for example, polymers of poly(2-methyl-2-oxazoline) (PMOXA), for example, having a molecular weight of 100-100,000 KDa; for example, as described in Konradi et al., 2012, Polyoxazolines for Nonfouling Surface Coatings — A Direct Comparison to the Gold Standard PEG, 33: 1663- 1676; or polymers comprising poly(2-methyl-2-oxazoline) / poly(acrylic acid) mixed brushes, for example, as described in Mahand et al., 2022, Polyoxazoline: A review article from polymerization to smart behaviors and biomedical applications, European Polymer Journal 178: 111484.

[0082] In some embodiments, the polymer coating includes polyglycerol fatty acid esters, for example, polymers of Linear methylated and hydroxylated oligo(glycidylethers), for example, as described in Calderon et al, Dendritic Polyglycerols for Biomedical Applications, Advanced Materials 22: 190-218; and polymers of Linear polyglycerol.

[0083] In some embodiments, the polymer coating includes poly(thioglycidyl glycerol) (PTGG), for example, having a molecular weight of 1000-100,000 KDa, for example, as described in d’Arcy et al., A Reactive Oxygen Species-Scavenging ‘Stealth’ Polymer, Poly(thioglycidyl glycerol), Outperforms Poly(ethylene glycol) in Protein Conjugates and Nanocarriers and Enhances Protein Stability to Environmental and Biological Stressors, J. Am. Chem. Soc.2022, 144, 46, 21304–21317.

[0084] In some embodiments, the polymer coating includes polycarbonate, for example, having a molecular weight of 1000-100,000 KDa, for example, as described in Engler et al.,2015, Hydrophilic Polycarbonates: Promising Degradable Alternatives to Poly(ethylene glycol)-Based Stealth Materials, Macromolecules 2015, 48, 6, 1673–1678.

[0085] In some embodiments, the polymer coating includes propylene glycol, for example, polymers comprising Poly(propylene glycol) methyl ether, -carboxy-terminated, forexample, as described in Wanka et al., Dendritic polyglycerols (PGs) were synthesized and postmodified by grafting of poly(ethylene glycol) (PEG) and polypropylene glycol (PPG) diglycidyl ether groups, Langmuir 2019, 35, 50, 16568–16575.

[0086] In some embodiments, the polymer coating includes polyaminoacids, for example, polymers comprising Poly ( , -l-aspartic acid), for example, having a molecular weight of100-100,000 KDa, for example, as described in Wang et al., 2011, Synthesis of a novel zwitterionic biodegradable poly ( , -l-aspartic acid) derivative with some l-histidine side-residues and its resistance to non-specific protein adsorption, Colloids and Surfaces 86: 237- 241.

[0087] In some embodiments, the polymer coating includes polyvinylpyrrol (PVP), for example, polymers comprising Polyvinylpyrrolidone, for example, having a molecular weight of 100-100,000 KDa, for example, as described in Guo et al., A polyvinylpyrrolidone-based surface-active copolymer for an effective marine antifouling coating, Progress in Organic Coatings 150: 105975.

[0088] In some embodiments, the polymer coating includes zwitterionic polymers, for example, polymers comprising zwitterionic poly (sulfobetaine methacrylate) (PSB) and polydopamine (PDA) for example, having a molecular weight of 100-100,000 KDa, for example, as described in Golabchi et al., 2021, Zwitterionic polymer / polydopamine coating reduce acute inflammatory tissue responses to neural implants, Biomaterials 225: 119519; and polymers comprising Zwitterionic carboxybetaine acrylamide, for example, having a molecular weight between 200-100,000, for example, 3-((3- ACRYLAMIDOPROPYL)DIMETHYLAMMONIO)-PROPANOATE.

[0089] In some embodiments, the polymer coating includes polybetaine, for example, polymers comprising polysulfobetaines (Poly (ethylene glycol)-block-Poly (sulfobetaine methacrylate)), for example, having a molecular weight of 100-100,000 KDa, for example, as described in Schonemann et al., 2021, Sulfobetaine Methacrylate Polymers ofUnconventional Polyzwitterion Architecture and Their Antifouling Properties, Biomacromolecules 2021, 22, 4, 1494–1508.

[0090] In some embodiments, the polymer coating includes polyester, for example, comprising Poly(PEG-hexadeca-di-oate); polyester based on poly(ethylene glycol) and hexadecanedioic acid, for example, having a molecular weight of 100-100,000 KDa; and polymers comprising 8-arm PLGA-SH, for example, having a molecular weight of 100- 100,000 KDa, for example, comprising Thiol capped 8-arm poly(lactide-co-glycolide).

[0091] In some embodiments, the polymer coating includes polymers of polysarcosine.

[0092] The polymer may be branched, highly branched, or hyperbranched. Hyperbranched may refer to the branch points of the overall polymer. Hyperbranched may include multi- armed polymers (e.g.4-Arm PEG compounds). Hyperbranched may also include multiple endpoints per monomer. The polymer may have an average of 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 20 more branches. Each branch of the polymer may have a length, for example, of at least 1 nm, at least 2 nm, at least 3 nm, or more than 3 nm. The height of the branched polymer may be greater than that of the recognition elements anchored on the nanoporous surface, for example, being, on average, at least 1.2.1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.9, 2, 3, 4, 5, 6, 7, 8, 9 or 10 times the height of the recognition elements. For example, aptamers have an average height of about 0.5 to 1, 1 to 2, 2 to 3, 3 to 4, or 4 to 5 nm. In some embodiments, the polymer is 8-branch PEG having a molecular weight of about 20,0000 Da. The PEG-SH is used as received without any pretreatment from Creative PEGworks Inc. PSB-852. The average height of this polymeric material deposited as a monolayer is about 4 to 5, 5 to 7, 7 to 10, or 10 to 15 nm. The polymer coating may include a polymer with a molecular weight of average molecular weight of 1,000 to 2,000, 2,000 to 3,000, 3,000 to 4,000, 4,000 to 5,000, 5,000 to 6,000, 6,000 to 7,000, 7,000 to 8,000, 8,000 to 9,000, 9,000 to 10,000, 10,000 to 11,000, 11,000 to 12,000, 12,000 to 13,000, 13,000 to 14,000, 14,000 to 15,000, 15,000 to 16,000, 16,000 to 17,000, 17,000 to 18,000, 18,000 to 19,000, 19,000 to 20,000, 20,000 to 21,000, 21,000 to 22,000, 22,000 to 23,000, 23,000 to 24,000, 24,000 to 25,000, 25,000 to 26,000, 26,000 to 27,000, 27,000 to 28,000, 28,000 to 29,000, 29,000 to 30,000 Da, or greater than 30,000 Da.

[0093] Exemplary polymers coating may include polyacrylamide-based copolymer hydrogels, for example as described in D. Chan et al., 2022, Combinatorial Polyacrylamide Hydrogels for Preventing Biofouling on Implantable Biosensors, Adv. Mater.34, 2109764;cholesterylated peptides, for example, as described in Stuhr-Hansen et al., 2016, Synthesis of Cholesterol-Substituted Glycopeptides for Tailor-Made Glycocalyxification of Artificial Membrane Systems, ChemBioChem 17: 403-406; poly-N-substituted glycines, for example, as described in Ham et al., 2013, Antifouling Glycocalyx-Mimetic Peptoids, J. Am. Chem. Soc.135: 13015–13022, the entire contents of all of which are incorporated herein by reference for all purposes.

[0094] The polymer coating may include functional groups. The functional groups may modify the hydrophobicity, charge conditions, packing density, and specific binding conditions of the surface to adjust the sensitivity and selectivity of sensing. The modifications may affect aptamers (e.g., enriching local analyte concentrations), reduce non-specific binding (e.g., protein and other interferent adhesion on sensor surfaces), or affect sensor kinetics (e.g., mass transport of analytes). The exemplary functional groups may include hydroxyl group or carboxyl group, for example as described in Kesler et al., 2023, Tailoring electrode surface charge to achieve discrimination and quantification of chemically similar small molecules with electrochemical aptamers. Advanced Functional Materials, 33(1), 2208534; amine group, biotin group, azide group, or silane group, for example, as described in Mertgen et al., 2022, A low-fouling, self-assembled, graft co-polymer and covalent surface coating for controlled immobilization of biologically active moieties. Applied Surface Science, 584, 152525. The entire contents of all of which are incorporated herein by reference for all purposes.

[0095] The polymer coating may include a polymer having a hydrodynamic size less than the average (e.g., mean or median) pore size of the nanoporous surface. For example, the polymer may have a size that is 10% to 15%, 15% to 20%, 20% to 25%, 25% to 30%, 30% to 35%, 35% to 40%, 40% to 45%, or 45% to 50% of the pore size. The polymer coating may extend 1 to 5 nm, 5 to 10 nm, 10 to 20 nm, or over 20 nm into the channel defined by the nanopore. The polymer coating has a thickness of 1 to 5 nm, 5 to 10 nm, 10 to 20 nm, or over 20 nm above a surface defining an opening of the nanopore.

[0096] In some embodiments, the polymer coating may include a layer covering the openings of the plurality of nanopores. For example, the polymer coating may overlay a surface including the openings of the nanopores. In some embodiments, the polymer coating may conformally or substantially conformally coat the surface of the nanopores. For example, a percentage (e.g., 5-10%, 10-20%, 20-30%, 30-40%, or 40-50%) of the volume of thenanopore may not be filled with a layer formed by the polymer coating. In some embodiments, the nanopores may be filled with the polymer coating. For example, no additional polymer coating could be added to the nanopores, even though the polymer coating itself may include some space between branches of the polymer.

[0097] The polymer coating may, in some embodiments, define one or more channels, a discontinuous series of pockets, or other complex geometries. The polymer coating may be a layer that includes these spaces as a result of the branching structure. In some embodiments, the polymeric coating overlays the opening of the channel and does not substantially ingress into the pores. In some embodiments, the polymeric coating substantially coats the inner surface of the pore, with channels or pockets within the pore not being covered with polymeric coating. In some embodiments, the pores are substantially filled with the polymeric material. The pore defines a path for the analyte to pass through to reach the molecular switch recognition element. The path may include continuous polymeric material, or intermittent or discontinuous polymeric material, for example including gaseous or hydrodynamic voids. The nanopore and the channel may be configured to permit the analyte to move from outside the pore to the molecular switch recognition element. The channel may be configured to reduce an amount of interferents from entering the nanopore. The channel may have an effective diameter in a range from 1 to 5 nm, 5 to 10 nm, 10 to 20 nm, 20 to 30 nm, 30 to 50 nm, or over 50 nm.

[0098] The polymer coating is deposited onto the nanoporous, recognition element functionalized surface. The deposition may be achieved by grafting the polymer to and / or from the nanoporous surface. In some embodiments, the polymers may be grafted to and / or from the nanoporous surface via dip coating the surface within the polymer solution in isopropyl alcohol. The density and polymer brush height can be tuned with solvent compositions, as described in Milner, S. T., Witten, T. A., & Cates, M. E.1988. Theory of the grafted polymer brush. Macromolecules, 21(8), 2610-2619, the entire contents of which are incorporated herein by reference for all purposes. Alternative methods for depositing polymers to the nanoporous surface include for example, grafting-to polymer deposition include spray coating, spin coating, roll coating, inkjet coating, screen printing, or dispensing.

[0099] FIG.6 shows an example of a sensor 600. Sensor 600 includes three polymer substrates with nanoporous layers (polymer substrates 604, 608, 612). The polymer substrates are on probe 616, which has a length of 20 mm and a width of 400 μm. The nanoporouslayers of the three polymer substrates are in electrical communication with wire 620 and other wires.

[0100] FIG.7 shows connections between the sensor 704 and a computer 708 configured to record the electrical signal. Computer 708 may be in communication with a potentiostat and multiplexer 712, which may detect and / or measure the electrical signal. FIG.7 shows wired connections between sensor 704 and computer 708 but any or all of the wires may be replaced with wireless communication (e.g., wireless transmitters and receivers). III. METHODS A. Fabricating the sensor

[0101] FIG.9 is a flowchart of an example process 900 of making a sensor (e.g., SENSBIT sensor). The sensor may be any sensor described herein.

[0102] At block 910, process 900 may include depositing an alloy of a first metal and a second metal over a substrate. The deposition may be simultaneous so that the metals are well mixed. Block 910 may be similar to stage 824. The alloy may be deposited onto a first metal layer. The first metal layer may include titanium and gold. The first metal layer may be deposited onto a second metal layer. The second metal layer may include gold. The second metal layer may be deposited onto a polymer substrate (e.g., stage 812). Block 910 is optional. In some embodiments, process 900 may include depositing a first material and a second material simultaneously, where the materials are not metals.

[0103] At block 920, process 900 may include removing the second metal to form a nanoporous layer may include the first metal. Block 920 may be similar to stage 828. Removing the second metal may include etching the second metal with nitric acid. Block 920 is optional. In some embodiments, the nanoporous layer may be formed by surface patterning (e.g., laser ablation or electron beam lithography).

[0104] At block 930, process 900 may include functionalizing the nanoporous layer with a set of recognition elements.

[0105] At block 940, process 900 may include depositing a polymer coating over the nanoporous layer. Process 900 may include adjusting the sizes of pores in the nanoporous layer by electrochemical coarsening or thermal annealing. The polymer coating may be dip coating described herein, including polymeric coating 108.

[0106] The nanoporous layer and the polymer coating may be passivated. As an example, the passivation may be with 6-mercapto-1-hexanol (MCH).

[0107] Process 900 may include additional implementations, such as any single implementation or any combination of implementations described herein and / or in connection with one or more other processes described elsewhere herein.

[0108] Although FIG.9 shows example blocks of process 900, in some implementations, process 900 may include additional blocks, fewer blocks, different blocks, or differently arranged blocks than those depicted in FIG.9. Additionally, or alternatively, two or more of the blocks of process 900 may be performed in parallel. B. Detecting analytes

[0109] Embodiments include detecting an analyte by an exemplary sensor. For example, an exemplary method encompasses the steps of utilizing a sensor system configured for the detection of a selected analyte in which the sensor element is deployed such that the sensing surface thereof is in contact with a sample, interrogating the sensor for detection of the selected analyte at one or more time points. The sample may include or be derived from any material, for example whole blood, such as flowing whole blood in vivo, interstitial fluid, plasma, saliva, urine, cell culture exudates, foodstuffs, drinking water, groundwater, and other materials wherein target analytes are to be detected. The sample may include an unprocessed sample, or a processed sample, for example, processed by filtering, centrifugation, dilution, or any suitable technique.

[0110] In some embodiments, the sensor element is deployed in vivo. In some embodiments, the sensor element is deployed ex vivo, for example, in a point-of-care device. In some embodiments, the sensor element is deployed for a period of one or more days, one or more weeks, or one or more months. In some embodiments, the sensor element is interrogated continuously, near-continuously, or at fixed intervals, for example, at intervals of seconds, minutes, hours, etc., for example, every 1-10 minutes, hourly, etc.

[0111] Measurements attained using the sensors of the invention may be compared against a standard curve or like tool for calculating target analyte concentration in the sample.

[0112] FIG.10 is a flowchart of an example of a process 1000 for detecting an analyte. In some embodiments, the process may be for measuring the concentration of an analyte.Process 1000 may include implanting the sensor in a human subject or any organism, including a rodent or other mammal. The sensor may be implanted into a blood vessel or a tissue, including onto or into the skin, the heart, the brain, the kidney, the lung, the pancreas, the bladder, the intestine, the gut, the stomach, or a tumor tissue.

[0113] At block 1010, process 1000 may include contacting the sensor to a sample that includes an analyte. The sensor may include any sensor described herein, including sensor 600.

[0114] At block 1020, process 1000 may include binding the analyte to a recognition element. A signal may be generated. For example, process 1000 may include sending an electrical signal resulting from the binding through the nanoporous layer. As an example, the electrical signal may result from increased electron transfer from moiety 124 to nanoporous gold layer 104. For example, an electrical signal resulting from the binding through the nanoporous layer may be sent through wire 620 and other wires. In other embodiments, the signal may be an optical signal.

[0115] At block 1030, process 1000 may include detecting the signal, thereby detecting the presence of the analyte. Detecting the electrical signal may use the potentiostat and multiplexer 712 and / or computer 708. Detecting the optical signal may use an imaging system and / or a computer.

[0116] Process 1000 may include releasing the analyte from the molecular switch. The release may be after block 1030. The binding of the analyte to the molecular switch may therefore be reversible.

[0117] The sensor may be used to detect or measure a set of analytes. The analyte may be a first analyte of the set of analytes. The molecular switch may be a first molecular switch of a set of molecular switches. The nanopore may be a first nanopore of a set of nanopores. The set of molecular switches may be disposed in the set of nanopores. The method may further include binding the set of analytes to the set of molecular switches. The electrical signal may result from the binding of the set of analytes to the set of molecular switches. The electrical signal may be measured. An amount of analytes may be determined using the measured electrical signal. The magnitude of the electrical signal may be proportional or directly related to the amount of analytes. The amount may be a raw number, a concentration, or a mass.

[0118] Process 1000 further includes disposing the sensor in an environment, where the environment would degrade or foul the molecular switch within seven days if the polymer coating were absent. The environment may be a chemical reactor, an industrial reactor, a manufacturing facility, or a food production reactor. The environment may include a biological fluid. The biological fluid may include plasma, serum, blood, urine, or saliva. Process 1000 may further include detecting the analyte 2, 3, 4, 5, 6, 7, 10 to 20, 20 to 30, 30 to 60, or 60 or more days after disposing the sensor in the environment.

[0119] Process 1000 may include additional implementations, such as any single implementation or any combination of implementations described herein and / or in connection with one or more other processes described elsewhere herein.

[0120] Although FIG.10 shows example blocks of process 1000, in some implementations, process 1000 may include additional blocks, fewer blocks, different blocks, or differently arranged blocks than those depicted in FIG.10. Additionally, or alternatively, two or more of the blocks of process 1000 may be performed in parallel. IV. EXAMPLES A. Assembling the SENSBIT device

[0121] FIG.8 is a schematic illustration of the fabrication process of a nanoporous gold sensor. The sensor may be any sensor described herein. At stage 804, a polyimide layer is on top of an aluminum sacrificial layer, which is on a silicon wafer. At stage 808, photoresist is deposited on the polyimide layer and patterned. At stage 812, gold is deposited on top of the photoresist and exposed polyimide. At stage 816, the photoresist is removed.

[0122] At stage 820, additional photoresist is deposited and patterned. At stage 824, bottom protective layer (BPL) is deposited. The BPL may be gold and / or titanium. An alloy of gold and silver is deposited (e.g., sputtered) onto the BPL. At stage 828, silver is removed with nitric acid, leaving a nanoporous gold layer. The nanoporous gold layer may be nanoporous gold layer 104 or any nanoporous gold layer described herein. At stage 832, photoresist is removed, leaving a stack of gold, BPL, and nanoporous gold on the polyimide.

[0123] At stage 836, a photoresist (e.g., SU8) is deposited onto the nanoporous gold to encapsulate the sides of the stack and patterned. At stage 840, an aluminum sacrificial layer isdeposited and patterned. At stage 844, polyimide is removed. At stage 848, the sacrificial layer is removed.

[0124] After stage 848, a molecular switch may be added into the nanopores of the nanoporous gold. The molecular switch may include receptor 112 with moiety 120 or any molecular switch described herein. For example, an aptamer may be functionalized with a moiety and then used to coat surfaces of the nanopores.

[0125] As an example, a DOX aptamer is dissolved in nuclease-free water with a concentration of 100 μM. This solution is reacted with a 1000-fold molar excess of TCEP solution with a 1:1 volume ratio for 1 hour, leading to reduction of the MB moiety and thiol- end group on the aptamer. The freshly prepared sensor can be rinsed with deionized water and then functionalized with 1 μM TCEP-treated DOX aptamer in 1× SSC buffer for 2 hours at RT. The sensor can be washed with excess buffer and incubated with 7 mM 6-MCH solution for 24 hours at RT to passivate the remaining electrode surface. The sensor can be stored in 1× SSC at 4°C until used for electrochemical measurement.

[0126] After the molecular switches are added into the nanopores of the nanoporous gold, the samples may be washed twice with PBS buffer. The polymer coating (e.g., PEG) may be added onto and / or into the nanoporous gold layer and incubated at room temperature for 1-12 hours. PEG may be prepared by dissolving in 70% isopropyl alcohol (1 wt%) with 65 °C heating and sonication. The samples may be then washed with excess PBS twice and incubated with ~10 mM 6-mercapto-1-hexanol (MCH) for 3 hours at room temperature to fully passivate the remaining electrode surface. A cross section of the sensor may be similar to that shown in FIG.1A.

[0127] FIG.11 shows SENSBIT fabrication and functionalization workflow. At stage 1104, gold and silver can be co-sputtered onto pre-cleaned glass slides at a 1:2 ratio. Stage 1104 may be equivalent to stage 824. At stage 1108, silver can be selectively wet etched with nitric acid. Stage 1108 may be equivalent to stage 828.

[0128] At stage 1112, aptamers can be immobilized onto the gold surface via thiol-gold binding. Stage 1112 may occur after stage 848. At stage 1116, hyperbranched polyethylene glycol modified thiol groups (PEG-SH) can be immobilized onto the gold surface via thiol- gold binding. At stage 1120, 6-mercapto-1-hexanol (MCH) can be immobilized for full passivation.B. Testing the SENSBIT device

[0129] FIG.1A shows an example of a SENSBIT device. The SENSBIT design includes a nanoporous gold layer 104, a protective polymeric coating 108, and aptamer switch-based receptors (e.g., receptor 112) for detecting relevant small-molecule analytes (e.g., analyte 116). Moiety 124 may be attached to receptor 112. When receptor 112 binds the analyte, receptor 112 may increase electron transfer from moiety 124 to nanoporous gold layer 104. Top panel 124 shows an electron microscopy cross-section of the sensor interface after cutting with a focused Ga+ion beam. The bottom panel shows a simplified diagram of the SENSBIT system.

[0130] The bicontinuous npAu surface may be custom designed to achieve an apical pore size of ~20.84 ± 7.47 nm, accessible for subsequent modification with the 30-nucleotide kanamycin aptamer switches (< 10 nm in length) via thiol-gold interactions. The aptamers may be modified with a methylene blue (MB) functional group at their 3’ terminus, and this moiety acts as the redox reporter for detecting target-binding events. The aptamer-loaded nanoporous gold surface may be subsequentially modified with a PEG-based polymer coating. Finally, the electrode surface can be passivated with 6-mercapto-1-hexanol (MCH), which coats the spaces remaining between the polymer molecules and aptamer switches to suppress non-specific redox reactions with the electrode surface.

[0131] An experiment can be performed to identify suitable polymer candidates for the SENSBIT coating layer, assessing PEG coatings of different sizes (molecular weight (MW) = 2,000 or 20,000 Da) and structures (linear or 8-armed / hyperbranched). A nuclease-based assay can be deployed to rapidly quantify the extent to which different coatings minimize the extent of protein interactions with aptamer switches on the SENSBIT surface.

[0132] FIG.1B shows how different coatings are screened. The SENSBIT polymeric coating design can be rapidly screened with high resistance to protein interference by spiking S1 nuclease at >10-fold physiological concentrations to accelerate aptamer switches degradation on planar gold (pAu) and nanoporous gold (npAu) surfaces. Vulnerability to protein interactions can be quantified based on electrochemical SWV peak signal loss.

[0133] S1 nuclease can be selected for various reasons. First, the ubiquity of endogenous nucleases in biofluids means that any DNA- or RNA-based biosensor will require robust nuclease resistance for long-term use. Additionally or alternatively, S1 nuclease is anendonuclease that specifically degrades single-stranded nucleic acids; as such, S1 nuclease activity will result in loss of MB redox reporter moieties, leading to a measurable loss of electrochemical signal. Additionally or alternatively, the hydrodynamic size of S1 nuclease is comparable to or smaller than most serum proteins, and thus offers a robust test of the size- exclusion mechanism of our mucosa-inspired nanostructures. SENSBIT systems modified with either linear PEG MW 2,000 or 8-arm hyperbranched PEG MW 20,000 can be exposed to S1 nuclease concentrations of 60 U / mL, which is more than an order of magnitude higher than physiological levels, in order to accelerate nuclease-induced aptamer degradation.

[0134] FIG.1C is a graph that characterizes the extent of degradation in real time by electrochemically measuring signal loss over time. The graph shows signal gain on the y-axis and time on the x-axis. We monitored real-time electrochemical signal decay of various aptamer-based sensor designs—planar (pAu) and nanoporous gold (npAu) electrodes either with or without hyperbranched MW = 20 kDa PEG—before and after spiking in S1 nuclease. A 30-min baseline measurement can be performed before spiking in 60 U / mL S1 nuclease to assess real-time aptamer-protein interactions.

[0135] FIG.1D is a graph of the signal gain after a 30-minute exposure to S1 for different aptamer-based biosensors with different PEG (none, 2kDa PEG, or 20 kDa PEG) and for planar gold and nanoporous gold. All datapoints and error bars represent the mean and standard deviation of four replicates.

[0136] As illustrated in FIG.1D, and after a 30-min nuclease challenge, the PEG-free pAu electrode experienced exponential signal loss of 85.6%, essentially rendering the sensor unusable. In comparison, PEG-free npAu or hyperbranched PEG-coated pAu electrodes achieved considerably reduced degradation, with 45.4% and 39.6% signal loss, respectively. The combination of npAu and linear PEG yielded significantly better synergistic protection against electrochemical decay, with signal loss of just 9.8% after 30 min. When the simple linear PEG coating was replaced with hyperbranched PEG, even greater suppression of electrochemical decay, with negligible SWV signal change (1.92 ± 0.46%), was observed after the 30 min assay. This indicates that this hierarchical sensor configuration – molecular switches coupling with protruding nanostructures and hyperbranched coating – can effectively shield aptamer switches from proteins in the sample. Therefore, unless otherwise indicated, hyperbranched PEG can be adopted as the polymeric coating materials for SENSBIT.

[0137] The hydrodynamically-dense nanostructures formed by the hyperbranched PEG layer raised potential concerns about whether target analytes have sufficient access to the aptamer switches immobilized within the nanostructure, and whether those captured biomolecules can subsequently exit the nanostructure and re-enter the bulk biofluid. The short-term sensitivity and reversibility of SENSBIT can be evaluated in a series of in vitro experiments with undiluted human serum. To assess sensitivity, the SWV signal gain produced by SENSBIT can be evaluated and pAu-based kanamycin sensors with and without a hyperbranched PEG protective layer in undiluted human serum ~1 min after spiking in kanamycin to a final concentration of 10 mM. Introducing a PEG coating onto the pAu surface undermined sensor sensitivity (signal gain = 31.06 ± 1.98%), resulting in a 17.59% lower signal change relative to a PEG-free pAu electrode (signal gain = 48.65 ± 3.99%). This confirmed the potential for additional steric hindrance that could affect target access to aptamer switches. However, the use of the npAu electrode with a hyperbranched PEG coating mitigated these deleterious effects, yielding a superior signal gain relative to the pAu systems of 81.18 ± 1.45%. This enhancement can be explained by the accelerated electron transfer that occurs within nanostructured electrodes.

[0138] Reversibility can be characterized by quantifying the extent to which the SWV signal returned to baseline after being challenged with 1 mM kanamycin in undiluted human serum and subsequently washing with kanamycin-free serum from the same batch. The unprotected PEG-free pAu control experienced considerable signal drift (13.69 ± 2.40%), and this poor reversibility is potentially due to nonspecific binding of serum proteins to the electrode surface, which in turn hinders the aptamers’ capacity to undergo target-binding induced conformational change. In contrast, the SENSBIT design (npAu plus hyperbranched PEG) enhanced reversibility and enabled the sensor to reset to near-baseline after washout with kanamycin-free serum (1.54 ± 1.35%). Together, these results indicate that this design is successful at preventing sensor contamination while still allowing target molecules to engage with aptamer switches on the electrode surface. C. Long-term, stable operation of SENSBIT in human serum

[0139] To evaluate the long-term stability of the sensors during extended exposure to complex biological matrices, the baseline response and sensitivity of unprotected sensors and the optimized SENSBIT system over the course of a month-long incubation in undiluted human serum can be quantified. Baseline stability can be quantified in terms of signaldeterioration over time in the absence of kanamycin, with SWV measurements collected on days 1, 3, 7, 14, and 28 and quantified relative to the sensors’ initial SWV peak current at day 0.

[0140] FIG.2A is a graph of the square wave voltammetry (SWV) assessment of electrochemical drift after exposing a PEG-free pAu (MCH-only) electrode to undiluted human serum for a month. The SWV frequency is 200 Hz, and the amplitude is 20 mV.

[0141] FIG.2B is a graph of the square wave voltammetry (SWV) assessment of electrochemical drift after exposing a SENSBIT electrode to undiluted human serum for a month.

[0142] FIG.2C is a graph of drift from baseline electrochemical signal over the course of a month-long incubation in human serum for SENSBIT versus the pAu sensor. The shaded area indicates the error bars of sensor signal changes (mean ± SD, n = 4).

[0143] The baseline signal from the PEG-free pAu electrode fell by >50% within seven days with progressively drifting redox-tag currents and unstable voltammograms, indicating considerable degradation of the sensor (FIG.2A). In contrast, SENSBIT preserved the reproducible voltammograms (FIG.2B) and maintained 90.27 ± 1.46% of its original baseline signal after one week—and more surprisingly, retained >70% baseline signal after 28 days of continuous exposure to undiluted human serum (FIGS.2B and 2C).

[0144] Long-term kanamycin sensitivity can be assessed by generating calibration curves based on signal gain from various concentrations of kanamycin at days 1, 3, 7, 14, and 28 of continuous exposure to undiluted human serum. Signal gain can be defined as the ratio of SWV peak current at a defined concentration relative to the target-free baseline. The calibration curves can be obtained by spiking fixed amounts of kanamycin up to a final concentration of 10 mM, while collecting measurements throughout. At the end of each measurement, the kanamycin can be washed out and replaced the sample with fresh, undiluted human serum with no kanamycin.

[0145] FIG.2D shows calibration curves obtained at multiple timepoints during a month- long incubation in undiluted human serum for pAu (mean ± SD, n = 4).

[0146] FIG.2E shows calibration curves obtained at multiple timepoints during a month- long incubation in undiluted human serum for SENSBIT (mean ± SD, n = 4).

[0147] For the PEG-free pAu control sensor, considerable loss of sensitivity over time was observed, with significantly reduced signal gain apparent by day 1 (FIG.2D). This indicates that aptamer-target binding was likely impaired either by non-specific binding of the sensor to plasma proteins or by degeneration of the aptamer switches. In contrast, SENSBIT exhibited minimal calibration curve drift over the course of a month in human serum (FIG. 2E), with reproducible signal gain observed at each target concentration. The signal gain data can be fitted to the Langmuir isotherm model, which assumes one target-binding site per aptamer switch: 100%in which is the fitting maximum signal gain of the sensor and is the equilibrium dissociation constant measured by the sensor.

[0148] FIG.2F shows a graph of the drift from fitted maximum signal gain of the sensor, B_max (R2 > 0.80) over the course of a month-long incubation in human serum for SENSBIT versus the pAu sensor. The shaded area indicates the 95% confidence intervals of Langmuir isotherm model fitting (n = 4).

[0149] Quantification of fitted Langmuir isotherm model revealed stable and of SENSBIT sensors maintained their initial levels with negligible drift along with month-long human serum challenge, while and of unprotected pAu sensors experience magnitudes of drift and progressively increasing errors (FIG.2F). These results indicate that the SENSBIT design effectively protects the sensor against fouling and degradation while enabling stable biochemical sensing operation for up to a month in undiluted serum. D. Protected biochemical interface in complex biological matrices

[0150] Sensing in whole blood introduces several additional challenges in comparison to serum. For example, despite advances in biomaterials, fouling on blood-interfacing devices remains a common problem, and intravenous devices often fail within a few hours after implantation due to extensive blood cell and clot deposition on the device surface. This contamination is initially activated by protein adsorption in which fibrinogen is typically the first plasma protein adhering to blood-interfacing devices, followed by platelet, leukocyte, and red blood cell adhesion and eventual material-induced thrombosis. For an electrochemical aptamer-based sensor, such non-specific fouling obstructs aptamer-target binding and sterically impedes the conformational change required for signal readout. Theblood cell coverage density on devices with different surface treatments can be examined after overnight (14 hour) incubation with human whole blood.

[0151] FIG.3A shows top-view SEM images of PEG-free pAu (left) and SENSBIT (right) sensor electrodes after 14 hour incubation with fresh human blood. Scale bars are 100 μm. Microscopic analysis revealed fewer blood cells adhering to the SENSBIT surface compared to the PEG-free pAu control

[0152] FIG.3B is a graph that shows the extent of blood cell fouling for PEG-free and PEG-coated pAu and npAu sensors (mean ± SD, n = 4). Even without PEG coating, the PEG- free npAu electrode showed only 2.83 ± 0.98% blood cell coverage relative to 28.47 ± 3.93% for the PEG-free pAu surface. The full SENSBIT design further suppressed blood cell adhesions to just 1.29 ± 0.27% surface coverage.

[0153] In principle, even low levels of blood cell fouling might still block analyte access to the sensor and undermine sensor sensitivity. In situ sectioning and visualization of the interface between single blood cells and the SENSBIT surface can be performed using focused ion beam scanning electron microscopy (FIB-SEM).

[0154] FIG.3C shows in situ focused ion beam scanning electron microscopy (FIB-SEM) cross-section image of a blood cell attached to SENSBIT. Scale bars are 1 μm. The magnification shows that direct interactions between SENSBIT and the cell membrane are limited to a small area, leaving most of the sensor’s biochemical interface intact. Scale bars are 200 nm. This cross-section revealed that blood cell coverage on SENSBIT only blocks limited regions of the apical surface, and that the complex porous structure of the SENSBIT electrodes should still support lateral target transport even if direct access from the apical surface is blocked.

[0155] We also assessed the potential for protein-induced blockage within the 3D structure of SENSBIT by characterizing the electrode interface after exposure to chicken plasma for 19 days in vitro. We chose serum rather than blood to eliminate potential interference from whole blood cells and pre-filtered the chicken plasma through a 0.2-μm filter to prevent protein aggregation and sedimentation prior to interaction with the electrode surface. Finally, we imaged the protein-fouled interface after washout.

[0156] FIG.3D shows SEM image of SENSBIT electrodes after incubation for 19 days in 0.2-μm filtered chicken plasma. Scale bars are 1 μm. Magnification shows a FIB-SEM cross-section revealing how the internal pore system remains unfouled even beneath apical surfaces exhibiting protein fouling. Scale bars are 200 nm. As expected, plasma protein fouled the SENSBIT apical surface and formed protein islands on top that could readily be visualized from 52° tilted SEM. By locally FIB-milling the fouled apical surface and in situ sectioning the protein-fouled SENSBIT surface regions, it can be determined that the internal porous infrastructure SENSBIT electrode remained unobstructed, with minimal fouling observed within. Small-molecule analytes can still freely access aptamers immobilized within the SENSBIT electrode via the bicontinuous 3D tunnels and neighboring openings, even for regions where the apical electrode surface is covered by a fouling ‘island’. This is in keeping with the robust sensitivity observed with SENSBIT even after a month-long incubation in human serum (e.g., FIG.2E).

[0157] Finally, the extent to which SENSBIT is protected against protein-surface interactions that arise during material-induced thrombosis can be assessed. In this state, fibrinogen anchors onto materials and subsequently serves as the backbone for further protein adsorption and thrombosis within minutes of surface contact. This could sterically affect the conformational change of aptamers for electrochemical reporting and hinder target transport within the sensor. To test this, PEG-free and PEG-coated pAu and npAu sensors can be incubated with human whole blood containing fluorescently-labeled fibrinogen (150 μg / mL) for 2 hours at 37 °C. The sensors can be washed and the electrodes can be imaged with an epifluorescence microscope.

[0158] FIG.3E shows fluorescence micrographs of sensor electrodes after 2 hours of incubation with fresh human blood containing fluorescently-labeled fibrinogen (150 g / ml). The micrographs show considerably more fibrin adhesion on the PEG-free pAu electrode (left) versus SENSBIT (right). Scale bars are 100 μm. Fibrinogen coverage was greatly reduced on the SENSBIT surface compared to PEG-free pAu.

[0159] FIG.3F is a graph of fibrin coverage for PEG-free and PEG-coated pAu and npAu electrodes (mean ± SD, n = 4). The use of either PEG-free npAu electrodes or PEG-coated pAu electrodes partially reduced fibrinogen coverage, but the bioinspired hierarchical configuration with npAu and hyperbranched PEG coating in the SENSBIT design reduced fibrinogen adhesion to just 0.27 ± 0.18% versus 8.53 ± 5.52% for PEG-free pAu. SENSBIT can effectively mitigate the impact of cellular and protein interferents in whole blood.E. Correction-free stable continuous intravenous biochemical monitoring with SENSBIT

[0160] SENSBIT’s intravenous real-time continuous monitoring performance over hours in anesthetized rats can be characterized. FIGS.4A-4E show in vivo real-time kanamycin monitoring over several hours with SENSBIT. FIG.4A illustrates the workflow of intravenous real-time continuous monitoring over hours in anesthetized rats. Kanamycin was administrated via retro-orbital injection.

[0161] The continuous biochemical monitoring is achieved with real-time electrochemical raw signal extraction followed by target estimation through pre-calibrated inverse-Langmuir standard curve to back-calculate contemporary target concentrations based on the fraction of target-bound switches read from electrochemical signal gain. However, for conventional pAu probes, the blood-device interactions including non-specific binding in blood and blood- induced degradation inevitably cause sensor electrochemical signal drifting along multi-hour recording, which contaminated the raw data for target estimation. To disentangle the target- induced electrochemical signal gain from the sensor electrochemical drift, there have been proposed numerous drift-correction algorithms including kinetic differential measurements (KDM) and dual-reporter drift corrections, which empirically normalized the aptamer loss on surface-modified biosensors along with biofluid exposure. Although empirically effective in short-term monitoring, the correction-based real-time monitoring has two caveats: (1) the blood-induced signal drift renders rapid absolute signal loss and causes indistinguishable SWV signal for data processing after hours, which prohibit potential signal normalization; (2) the existing baseline-correction algorithms were built on a strong assumption where the biosensor’s sensitivity based on aptamer switches remained on surface after blood interactions is identical to that from intact biochemical chemical interface with shared calibration curves, which were tested negative with continuously decaying electrochemical sensitivity for long-term biochemical in protein-rich human serum environment and post- intravenous-surgery calibration measurement (e.g., FIG.2D). Therefore, SENSBIT can achieve protected biochemical sensing and minimize the electrochemical signal drift during multi-hour interface within flowing blood for direct target-induced electrochemical signal extraction based on raw SWV data.

[0162] An implantable microwire probe with SENSBIT (or pAu as a control) bundled with an Ag / AgCl reference electrode and Pt counter electrode can be prepared. FIG.4B is anoptical image of SENSBIT intravenous probe compared to 22G clinical catheter. Scale bar is 10 mm.

[0163] During the surgical procedure, the probes can be directly implanted and fed into rat femoral veins with minimal tissue damage. FIG.4C illustrates the SENSBIT intravenous probe implanted within a femoral vein to achieve a direct interface with the blood in anesthetized rats. Insets are schematic and surgical images of implanted SENSBIT. The implanted probes can be secured onto the femoral vein by tissue glue and surgical sutures at the connection ends. Successful intravenous access was confirmed based on blood flow after retracting the implanted probes. The I / O port of the probe can be connected to an external potentiostat for real-time SWV data extraction and analysis with customized codes, while the three electrodes can be fully exposed to the blood within femoral veins. The raw electrochemical SWV signal can be continuously read out without administrating kanamycin. Following the collection of the raw electrochemical signal, target estimation can be performed by using a pre-calibrated inverse-Langmuir standard curve to back-calculate target concentrations based on the electrochemical signal gain:SENSBIT experienced minimal baseline signal drift during 2.5 hours of continuous monitoring, with no need for drift-correction algorithms, while pAu probes experienced progressive signal degradation with ~30% signal loss after 2.5-hour scanning. This allowed validation that the SENSBIT sensor maintains baseline operational stability over extended durations with minimal drift following intravenous implantation. SENSBIT’s capacity for intravenous monitoring of real-time changes in kanamycin pharmacokinetics can be characterized.

[0164] FIG.4D is a graph of real-time SWV scanning with intravenously implanted pAu over several hours after the injection of a single bolus of drug (time marked with dashed line). Inset is the pAu sensor calibration curve in buffer before implantation (3 measurements per concentration). The SWV frequency is 50 Hz. The amplitude is 20 mV. After 10-40 minof stable baseline visualization, kanamycin stock is dosed via retro-orbital sinus injection (400 mg / kg). While pAu probes exhibited negligible target-induced signal change coupled with naturally degrading SWV signals over continuous blood interface, SENSBIT probes monitored target-induced electrochemical signal gain with enhanced and minimal sensor electrochemical drift. The real-time SWV signal initially increased and reached a peak ~36 min after injection due to redistribution in flowing blood, after which the signal gradually decayed due to elimination and metabolism over the following three hours. To enable faithful target estimation with pre-calibrated calibration curves, the biosensors can be calibrated in vitro before and after the multi-hour intravenous interface and confirmed that there was minimal decay in the sensitivity of our implanted biosensors, with reproducible signal gain observed at each target concentration. This allowed calculation of intravenous kanamycin concentrations at each timepoint from the raw SWV signals by applying these pre-calibrated SENSBIT calibration curves.

[0165] FIG.4E shows real-time tracking of kanamycin concentration with SENSBIT over several hours after the injection of a single bolus of drug (time marked with dashed line) without the need for signal correction. The raw signal from SENSBIT provides an accurate measure of real-time intravenous kanamycin concentrations as derived from a pre-calibrated sensor binding curve (3 measurements per concentration). The SWV frequency is 50 Hz. The amplitude is 20 mV. F. Long-term stability after multi-day implantation in free-moving rats

[0166] The stability of the baseline response and sensitivity of SENSBIT can be characterized after an extended duration of intravenous implantation in free-moving rats. Given the difficulty of distinguishing the target-induced signal gain from the effects of drift during in vivo experiments, SENSBIT’s stability can be assessed by performing sensor calibration in PBS containing defined concentrations of kanamycin both before and after various durations of intravenous implantation. Implantation was performed as described herein, and after implantation and recovery, the rats were able to move freely while the implanted electrode probes were constantly subjected to blood flow and mechanical strain.

[0167] FIG.5A shows the workflow for the evaluation of long-term SENSBIT implantation in free-moving rats. Sensors were surgically implanted within rat femoral veins for multiple days before being extracted for further evaluation.

[0168] To begin, the baseline stability of the sensors after multiday intravenous interface was quantified by calibrating and comparing sensor performance in kanamycin-free PBS buffer with external reference and counter electrodes before and after implantation.

[0169] FIG.5B shows signal degradation after two days of continuous implantation of a PEG-free pAU sensor. FIG.5C shows signal degradation after two days of continuous implantation of a SENSBIT electrode. The pAu sensor had lost its baseline aptamer-MB signal with indiscernible SWV peaks, while SENSBIT exhibited robust aptamer-MB currents and maintained comparable SWV peak levels according to the voltammograms after 2-day intravenous implantation.

[0170] Next, it was studied how baseline stability was affected by both surgical implantation and protracted sensor interface with the intravenous environment. The former is associated with acute mechanical damage to the probe during implantation as well as fast- activation blood interactions (e.g., fast adsorption of plasma proteins, and adhesion of nonactivated platelets), whereas the latter involves progressive contributions of factors such as continuous blood shearing, nuclease degradation, and deposition of proteins and cells onto the probe surface. Surgical baseline stability was quantified as the ratio of the baseline SWV signal in buffer after 20 min of intravenous implantation to the baseline pre-implantation SWV signal of the same probe in buffer. SENSBIT probes exhibited superior surgical baseline stability of 82.58 ± 13.16%, whereas the PEG-free pAu probes retained just 14.99 ± 13.75% of their baseline signal level post-surgery. The longer-term baseline stability was quantified as the ratio of the extracted baseline SWV peak signal in buffer after multiple days of intravenous implantation in free-moving rats relative to the baseline obtained prior to implantation. This value was then normalized by dividing by the ratio that we calculated for the surgical baseline stability, as described above. This allowed quantification of the progressive intravenous effects that reduce biosensor performance over a multi-day interface with the circulation.

[0171] FIG.5D shows baseline electrochemical signals of the two sensors after multi-day implantation. The data were normalized to the signal loss after surgical damage. The pAu probes underwent rapid signal decay during intravenous interface, with just 6.67 ± 3.91% baseline signal remaining after one day of implantation after normalization to the acute surgical baseline loss. The individual SWV peaks were indistinguishable after even a single day of implantation.

[0172] FIG.5E is a graph summarizing papers on device longevity of existing aptamer- based electrochemical sensors interfacing blood derivatives and whole blood. The different shapes and patterns of the symbols were used as visual aid to distinguish this work from existing strategies. The poor chronic baseline stability of pAu probe was also aligned with the limited blood-interfacing aptamer-based biosensor longevity (defined as the period that the sensor can maintain >50% baseline signal) that remained less than 11 hours developed by existing protecting strategies, including physically shielded biosensor, SAM-engineered biosensor, nanostructured biosensor, aptamer-engineering biosensor, drift-corrected biosensor. Importantly, SENSBIT retained 83.04 ± 11.92% of its baseline signal after one day of intravenous implantation and 78.50 ± 10.23% of its baseline signal after four days. The SENSBIT technology leveraging the hierarchical nanostructure has significantly extended the device longevity in blood for days (from 12 hours to 96 hours).

[0173] In order to quantify sensitivity stability, SWV signal gains after exposing sensors to defined concentrations of kanamycin in PBS can be compared before and after intravenous implantation. After just 1 day of implantation, the pAu sensor cannot faithfully read out aptamer-MB SWV signals.

[0174] FIG.5F shows calibration curves of pAu before intravenous implantation and after 1 day implant. By manually extracting the SWV signals, pAu sensors exhibited severe deviation in its calibration curve and greatly distorted sensitivity, such that the post- implantation sensor could no longer faithfully report defined kanamycin concentrations.

[0175] FIG.5H shows calibration curves of SENSBIT after one day, four days, or seven days of chronic implantation. SENSBIT effectively overcame issues that faced pAu, with near-identical calibration curves at days 0, 1, 2, 4, and 7 post-implantation. This unprecedented stability indicates that the SENSBIT design durably protects the electrochemical aptamer-based sensor against fouling and degradation even after multiple days of continuous implantation and retains the capacity to accurately report analyte concentrations.

[0176] SENSBIT’s capacity to respond to changing concentrations of kanamycin can be evaluated before and after multiple days of implantation in freely moving rats. The biosensor probes were continuously measured in buffer (10 scans) with sampling rate of ~3 min-1and then subsequently challenged with 1 mM (5 scans), 0 mM (5 scans), 1 mM (5 scans), 2 mM (5 scans), and 0 mM (5 scans) kanamycin.

[0177] FIG.5G shows continuous SWV scanning to extract real-time kanamycin concentrations in buffer with pAu before and after 1 day implantation. The pAu biosensor failed to report these dynamic changes in kanamycin concentration after just 1 day of intravenous interface, with noisy signal changes and a signal readout that severely deviated from pre-implantation measurements. FIG.5G shows that the pAu biosensor cannot provide useful data in this experiment.

[0178] FIG.5I shows continuous kanamycin monitoring with SENSBIT before and after 1-, 4-, and 7-day implantation in femoral veins of free-moving rats. For SWV data, frequency is 50 Hz, and the amplitude is 20 mV. In FIGS.5F and 5G, * indicates that SWV peaks were inadequately resolved for quantitative data analysis.

[0179] In striking contrast to pAu, SENSBIT retained real-time sensing capabilities with reproducible signal gains and comparable kinetics even after multi-day intravenous implantation. Even after four days of implantation, the retracted SENSBIT probes could dynamically track kanamycin concentration fluctuations in real time, with a ~3 min-1reading frequency and a response profile that closely matched the readings collected pre-implantation on day 0. Importantly, the retracted SENSBIT consistently returned to baseline level in the absence of kanamycin with minimal hysteresis.

[0180] FIG.12 shows a graph of intravenous real-time monitoring of kanamycin pharmacokinetics during anesthesia according to embodiments of the present invention. In some embodiments, FIG.12 may illustrate intravenous real-time monitoring of kanamycin pharmacokinetics during anesthesia after one day, two days, or four days of implantation in freely-moving rats with SENSBIT. Real-time SWV readings with SENSBIT sensors are from three different rats one day (a), two days (b), or four days (c) after i.v. implantation. Readings were collected over several hours after the injection of a single bolus of drug (400 mg / kg, injection time marked with vertical dotted line). Insets show sensor calibration curves in buffer before implantation (three measurements per concentration).

[0181] In some embodiments, SENSBIT is capable of intravenous real-time monitoring of kanamycin pharmacokinetics during anesthesia after multi-day of implantation in freely- moving rats. The implantable SENSBIT microwire probes bundled with Ag / AgCl reference electrode and Pt counter electrode were implanted in rat femoral veins. After establishing baseline SWV readings, the I / O port of the probe is disconnected from the external potentiostat and surgically covered with subcutaneous tissues, and the wound is suturedbefore waking the rats. The rats were able to move freely while the probes were constantly subjected to blood flow and mechanical strain. After multiple days, rats are put back under anesthesia and re-exposed the I / O port of the probe for SWV analysis. The raw voltammograms after 25-hour implantation maintained clear SWV peaks, while pAu control probes had indiscernible SWV peaks. SENSBIT’s capacity for intravenous monitoring of real-time changes in kanamycin pharmacokinetics after multi-day of continuous blood interface is characterized in FIG.12. After 15 to 60 min stable baseline visualization, kanamycin is administered via retro-orbital sinus injection (400 mg / kg), and observed corresponding SWV signal change with initial increase over the first 20 to 60 min followed by gradual decay due to elimination and metabolism. Calibration curves of the SENSBIT probe in vitro are generated before and after one-day, two-day, or four-day implantation and it can be confirmed that there was minimal sensitivity drift or decay, with reproducible signal gain observed at each target concentration. Therefore, the pre-implantation calibration curves for each SENSBIT probe can be used to derive intravenously-measured kanamycin concentrations with inverse-Langmuir standard curves in real time, without the need for correction or post-calibration. SENSBIT’s operational stability thus enables long-term tracking of kanamycin pharmacokinetics from individual animals, even four days after the sensor intravenous implantation.

[0182] In some embodiments, real-time measurements of kanamycin dose-in and wash-out in undiluted human serum may exhibit no hysteresis, and similar on / off response times for SENSBIT in kanamycin sensing and regeneration compared to pAu sensors over the course of 50 min. Kinetically, SENSBIT may show no delay compared to a pAu control in reading out kanamycin spike-in and wash-out, with a temporal resolution of 4 s. The SENSBIT architecture can maintain kanamycin transport across its nanostructures without compromising sensor performance in terms of sensitivity, reversibility, or temporal resolution. The same SENSBIT architecture can be generalized to sense a variety of small- molecule analytes. For example, the chemotherapy drug doxorubicin (DOX) and ATP can be sensed with SENSBIT by simply incorporating the appropriate aptamer switches. G. Discussion

[0183] In summary, a hierarchical sensor design is demonstrated that enables long-term stable biochemical monitoring in complex biological matrices with robust performance for at least a month. The sensor’s functional stability during long-term exposure is enhanced inundiluted serum or intravenous blood with a nanoscale physically restricted zone for protected molecular recognitions against interferents. As a result, SENSBIT retained >70% baseline signals and reproducible calibration curves after one-month incubation in direct contact with undiluted human serum in vitro, showing significantly better performance than existing electrochemical aptamer-based biosensors (FIG.5E).

[0184] The SENSBIT can be challenged for real-time pharmacokinetic monitoring of kanamycin in the bloodstream, and it can be validated that SENSBIT exhibited minimal electrochemical signal drift during multi-hour interface with flowing blood. This unprecedented baseline stability successfully disentangled target-induced signal gain from sensor signal drift during long-term biointerface, enabling real-time biochemical concentration extraction via pre-calibrated calibration curves. The sensing stability at an extended period (up to four days) of the intravenous interface in free-moving rats was validated. Except for superior baseline stability (maintaining >75% baseline signal after 4- day intravenous implantation) compared to any other published in vivo sensor described to date, SENSBIT also exhibited reproducible calibration curves and faithful real-time sensing capability after multi-day intravenous implantation, which has not been reported before. Collectively, the results suggest that the simple multi-component sensor design may serve as a general and robust platform for long-term monitoring of blood biomarkers matching clinical and non-clinical needs.

[0185] As there have been certain commercially available compact or wearable sensors that enable real-time measurement of circulating analytes over protracted periods of time, such as continuous glucose meters (CGM) that have revolutionized care for diabetic patients, it is noteworthy that none of them so far can continuously interface flowing blood, which would inevitably undermine the quality of real-time biochemical monitoring. For example, as the biofluid CGM is interfacing, interstitial fluid has a delay of ~15 minutes from actual glucose concentrations in the bloodstream. Especially for emergency medicine, surgery, and treatment of infectious diseases, it would be necessary to technically enable instant biomarker measurements through real-time continuous sensing directly from blood. As a powerful tool to prolong the biochemical sensor longevity in blood, SENSBIT could potentially give researchers and clinicians the capability to set up a chronically stable monitoring system for individual healthcare.

[0186] The results discussed herein demonstrate a step towards a solution to a critical biosensing challenge: achieving stable, accurate, continuous molecular detection directly in the bloodstream for extended periods of time. The SENSBIT platform can be generalized to experiments in which electrochemically-modified aptamer switches that respond to DOX or ATP are incorporated. H. Methods

[0187] Materials

[0188] All chemicals were purchased from Sigma-Aldrich, unless otherwise noted. Thiolated polyethylene glycol (PEG-SH), including linear methoxy PEG (mPEG-SH, MW 2k, 10k, 20k), branched PEG (4-arm PEG-SH, MW 10k, 8-arm PEG-SH, MW 20k), was ordered from Creative PEGWorks. Fluorescent-labelled fibrinogen from human plasma, S1 nuclease, nuclease-free water (NFW) were purchased from Thermo Fisher Scientific. Human whole blood was ordered from the Stanford Blood Center. Pooled human plasma (blood- derived) and chicken plasma were purchased from Innovative Research. Kanamycin monosulfate, USP Grade, was purchased from Gold Biotechnology. PFA-coated platinum wires (bare diameter 0.003”), PFA-coated tungsten wires (bare diameter 0.002”), and PFA- coated silver wires (bare diameter 0.008”) were purchased from A-M Systems. Gold wires (diameter 0.01”) and 10k gold wire (diameter 0.01”) were purchased from Alfa Aesar and Amazon, respectively. Miniature heat-shrink tubes (pre-shrinking inner diameters 0.006” or 0.01”) were purchased from Nordson Medical, and McMaster-Carr. The kanamycin aptamer was ordered from Integrated DNA Technologies, and the sequence used was / 5ThioMC6- D / GGGACTTGGTTTAGGTAATGAGTCCC / 3MeBlN / , where 5ThioMC6-D indicates a thiol modification at the 5’ end and 3MeBlN indicates methylene blue (MB) labeling at the 3’ end. Phosphate-buffered saline (PBS) buffer in this work was prepared by diluting 10X PBS and MgCl2 with NFW to 1X PBS and 2 mM Mg2+.

[0189] Planar gold and nanoporous gold preparation

[0190] For in vitro pAu sensors, a 90-nm-thick Au layer was patterned (1 mm × 1 mm for each working electrode) onto a pre-cleaned glass slide via shadow mask by electron beam evaporation (ATC-E, AJA International) after 10 nm Ti deposition. For in vitro SENSBIT sensors, nanoporous gold (npAu) film was patterned with shadow mask to produce 1 mm × 1 mm working electrodes (WE) on pre-cleaned glass slides with Ti / Au / Ag:Au (10 / 50 / 300 nm).The ratio of cosputtered Ag:Au film was about 2:1 (Lesker Sputter). Ag was selectively etched by nitric acid (70% v / v) for 3 hours, forming the npAu electrodes. For in vivo pAu microwire sensors, gold wires were cut into ~6 cm lengths and used as received. For in vivo SENSBIT microwire sensors, nanoporous gold wires were prepared by etching 10k yellow gold wire in nitric acid (70% v / v) for 15 min. The prepared pAu and npAu electrodes were sonicated with acetone, 70% isopropyl alcohol, and NFW for 2 min respectively. The cleaned pAu and npAu electrodes were stored in NFW before functionalization.

[0191] Microwire sensor preparation for in vivo applications

[0192] The pAu or npAu microwire WEs were insulated with heat-shrink tubes around the body of the wires, with an about 2–5-mm length of electrode exposed for aptamer functionalization. The reference electrode (RE) was prepared by stripping a ~1 cm length of PFA insulation off the PFA-coated silver wires, followed by bleach oxidation of an 8-mm length of the exposed silver wire. The counter electrode (CE) was prepared by stripping an about 5 mm length of PFA insulation off the PFA-coated platinum electrode. The three electrodes were assembled and bundled with heat-shrink tubing and washed with 70% isopropyl alcohol and NFW. Heat-shrink tubes with lengths of about 2–4 cm were used to strengthen the stem of the microwires for surgical manipulation.

[0193] pAu sensor and SENSBIT sensor preparation and functionalization

[0194] Aptamer switches were suspended in NFW at a final concentration of about 100 μM and stored at about -20 °C before usage. To produce free thiol groups for aptamer immobilization, a 1,000-fold molar excess of freshly prepared Tris(2-carboxyethyl)phosphine hydrochloride (TCEP) solution was added to the aptamer switch solution and reacted for about 1 hour at room temperature in dark. The mixture was then diluted to about 1 μM aptamer in PBS. pAu and npAu electrodes were briefly dried with N2gas. For in vitro experiments, a polydimethylsiloxane (PDMS) chamber was prepared by punching an about 6- mm-diameter hole in the 5-mm-thick PDMS film and subsequently putting the PDMS film onto the glass slide, such that the WEs were confined within the chamber. For in vivo experiments, the microwire electrodes or sensors were incubated in tubes. The samples were then immersed in the about 1 μM aptamer solution for 40 min. After immobilization, pAu and npAu electrodes were washed twice with PBS. For the SENSBIT sensors, polymer coating was achieved by incubating the aptamer-immobilized npAu electrode with PEG solutions at room temperature for about 3 hours. PEG was prepared by dissolving in 70%isopropyl alcohol (1 wt%) with 65 °C heating and sonication. The pAu and npAu electrodes were then washed with excess PBS twice and incubated with about 10 mM 6-mercapto-1- hexanol (MCH) for about 3 hours at room temperature to fully passivate the remaining electrode surface. The pAu sensors and SENSBIT sensors were washed in excess PBS twice and then stored in PBS buffer at room temperature before electrochemical measurement.

[0195] Electrochemical characterizations of aptamer-based biosensors

[0196] Electrochemical measurements were conducted with PalmSens 4 potentiostat (PalmSens) and EmStat4 (PalmSens) with multiplexer (MUX8-R2 or MUX8, PalmSens). For in vitro experiments, the measurements were performed in PDMS wells. Commercial Ag / AgCl reference electrodes and Pt wire counter electrodes were used as received from CH Instruments. SWV was carried out in buffers or corresponding complex biological matrices over the potential range of about -0.5 V to about 0.0 V with an amplitude of 20 mV, step size of 2 mV, and pulse frequencies ranging from 20 to 200 Hz. To measure biosensor sensitivity in vitro, we obtained dose-response curves by subjecting the sensors to kanamycin concentrations ranging from about 0.01–10 mM in various biological matrices by spiking in kanamycin stock solution (dilution < 1% v / v) at room temperature. Data processing and visualization were performed with custom MATLAB code to automatically extract the SWV peak current levels. The calibration curves were generated by fitting the dose-response curves to the Langmuir isotherm model (GraphPad Prism). Long-term incubation was achieved by loading about 100–200 μL undiluted biofluids into the PDMS chamber, which was temporarily sealed with a PDMS lid before next measurement. For chronic measurements, biofluids were aliquoted from the same batch of biological matrix and thawed 2 hours before the measurement followed by gently mixing.

[0197] For in vivo intravenous measurement, bundled WE, RE, and CE were connected to the multiplexer with a customized adaptor that minimizes the strain of the microwire electrodes. SWV was carried out after implantation over the potential range of about -0.5 V to about 0.0 V with an amplitude of 20 mV, step size of 2 mV, and pulse frequencies ranging from 20 to 100 Hz, adjusted according to the signal levels and signal noise. Real-time data processing and visualization were performed with custom MATLAB code to automatically extract the SWV peak current levels. For sensor calibration measurements before and after in vivo implantation, pAu or SENSBIT sensors were connected to the multiplexer with the customized adaptor and measured in a tube. To calibrate the biosensor sensitivity, the dose-response curve was obtained by sequentially adding kanamycin stock solution to the buffer, followed by about 10 times pipette mixing. For real-time measurements before and after in vivo implantation, pAu or SENSBIT sensors were continuously measured at a sampling rate of about 3 min-1. Kanamycin concentration changes were achieved by dipping the sensors into kanamycin solution of different concentrations.

[0198] SENSBIT optimization via nuclease-aided screening platform

[0199] Biosensors with immobilized aptamers and various surface conditions were prepared on glass slides and incubated in 200 μL S1 nuclease buffer with 100 mM Zn2+(Thermo Fisher Scientific) before electrochemical characterization. During screening, the biosensors were continuously measured with SWV scanning at a sampling rate of 0.5 min-1. After stabilizing for about 30 min in S1 nuclease buffer under continuous electrochemical characterization, S1 nuclease was added to the biosensor buffer to reach a final nuclease concentration of 60 U / mL. The continuous electrochemical signals were tracked and analyzed with custom MATLAB code to automatically extract the SWV peak current levels.

[0200] Whole blood adhesion assay

[0201] Human whole blood was freshly prepared by the Stanford Blood Center within one day and used as received at room temperature. After gentle mixing of the fresh blood, about 100 μL human whole blood was added to in vitro sensors within the PDMS chamber and incubated at room temperature in the dark with a sealing PDMS lid. For EM characterization, 20 μL 1 M CaCl2 was added to the PDMS chamber and incubated for 10 min, followed by a gentle PBS wash via pipette. After removing the PDMS chambers, the samples were immersed in fixation solution (4% paraformaldehyde, 0.4% glutaraldehyde in 0.2 M sodium cacodylate) at room temperature for 10 min. After three washes with NFW, the samples were immersed in 1% OsO4staining solution for 1 hour in the dark. After three washes with NFW, the samples were transferred to a series of ethanol baths: 30%, 50%, 70%, 80%, 90%, 96%, and absolute (each 10 min). The fixed samples were transferred to a critical point dryer (Tousimis 815, Series A) and dried. Finally, the samples were mounted and sputtered-coated with a thin layer of Au / Pd (Denton Desk II Sputter Coater). After imaging, pseudocolor was applied to provide contrast of blood cell adhesion on samples. According to the visually distinguishable cells, blood cell coverage was identified within four randomly selected regions and extracted by bandpass filtering, thresholding, and binarizing with Fiji ImageJ (four areas per sample were obtained), followed by MATLAB analysis. For the fibrinogenassay, fluorescently-labeled fibrinogen (150 g / ml) was added to the blood, which was then incubated in the sensor sample chamber for 2 hours at 37 °C. After washing with PBS, the samples were imaged with a Leica DM4 upright epifluorescence microscope. Quantification was performed by thresholding the fluorescent features, followed by binarization and ratio calculation with MATLAB. The threshold was manually set to 800 for all images, and the area covered by fibrin was extracted from the measurements (four areas per sample were obtained).

[0202] Surface and tomography characterization by SEM, FIB-SEM

[0203] Sensor surface characterization was carried out by SEM (Thermo Fisher Scientific Apreo S LoVac and FEI Magellan 400 XHR). The samples underwent plasma cleaning within the SEM chamber and were imaged with 2-5 keV. Tomography characterization was carried out using a FIB-SEM (FEI Helios NanoLab 600i DualBeam), which combines high- resolution SEM imaging and FIB milling. To characterize the cross-section of the nanoporous gold structure, a thin layer (about 300 nm) of Pt was deposited on the field of interest before FIB milling. For cell or protein adhesion characterization, no Pt was deposited within the SEM chamber in order to preserve the intact interface between sensor and biological materials. To expose the sensor interface, fouling components on the electrode surface were first located in top-view SEM, after which FIB was used in situ to carry out vertical dissection at the desired locations. The dissection cross-sections were visualized through SEM via 52° tilting.

[0204] Live animal studies

[0205] Live animal studies were performed with male Sprague–Dawley rats under Stanford APLAC protocol number 33226. All rats used in this work were of weight of 200–450 g. The rats were anesthetized using isoflurane gas (2.5%) during implantation procedures and monitored continuously. The femoral veins of the rats were isolated and exposed with a small incision, through which the microwire probes were implanted and pushed far into the vein. The microwire probes were secured in the place while maintaining the intravenous blood flow, after which the incision was closed and sutured with absorbable sutures and a small amount of tissue glue to anchor the suture wire. For the in vivo pharmacokinetic monitoring experiments, the probes were connected to the external potentiostat and continuous SWV scanning was performed under anesthesia.100 mM kanamycin was administrated via retro- orbital sinus injection to reach 400 mg / kg dosage. For multi-day sensor implantation in free-moving rats, the microwire probes were sterilized with a UV light sanitizer prior to implantation. The implantation wounds were sutured and secured with wound clips. The animals were monitored daily for post-surgery recovery and maintained for up to four days. At the end of the experiments, animals were anesthetized using isoflurane gas and the microwire probes were retracted for post-implantation analysis. Rats were then euthanized by exsanguination while under general anesthesia.

[0206] Any of the methods described herein may be totally or partially performed with a computer system including one or more processors, which can be configured to perform the steps. Thus, embodiments can be directed to computer systems configured to perform the steps of any of the methods described herein, potentially with different components performing a respective step or a respective group of steps. Although presented as numbered steps, steps of methods herein can be performed at a same time or at different times or in a different order that is logically possible. Additionally, portions of these steps may be used with portions of other steps from other methods. Also, all or portions of a step may be optional. Additionally, any of the steps of any of the methods can be performed with modules, units, circuits, or other means of a system for performing these steps.

[0207] As will be apparent to those of skill in the art upon reading this disclosure, each of the individual embodiments described and illustrated herein has discrete components and features which may be readily separated from or combined with the features of any of the other several embodiments without departing from the scope or spirit of the present disclosure.

[0208] The above description of example embodiments of the present disclosure has been presented for the purposes of illustration and description and are set forth so as to provide those of ordinary skill in the art with a complete disclosure and description of how to make and use embodiments of the present disclosure. It is not intended to be exhaustive or to limit the disclosure to the precise form described nor are they intended to represent that the experiments are all or the only experiments performed. Although the disclosure has been described in some detail by way of illustration and example for purposes of clarity of understanding, it is readily apparent to those of ordinary skill in the art in light of the teachings of this disclosure that certain changes and modifications may be made thereto without departing from the spirit or scope of the appended claims.

[0209] Accordingly, the preceding merely illustrates the principles of the invention. It will be appreciated that those skilled in the art will be able to devise various arrangements which, although not explicitly described or shown herein, embody the principles of the invention and are included within its spirit and scope. Furthermore, all examples and conditional language recited herein are principally intended to aid the reader in understanding the principles of the disclosure being without limitation to such specifically recited examples and conditions. Moreover, all statements herein reciting principles, aspects, and embodiments of the invention as well as specific examples thereof, are intended to encompass both structural and functional equivalents thereof. Additionally, it is intended that such equivalents include both currently known equivalents and equivalents developed in the future, i.e., any elements developed that perform the same function, regardless of structure. The scope of the present invention, therefore, is not intended to be limited to the exemplary embodiments shown and described herein. Rather, the scope and spirit of present invention is embodied by the appended claims.

[0210] A recitation of “a”, “an” or “the” is intended to mean “one or more” unless specifically indicated to the contrary. The use of “or” is intended to mean an “inclusive or,” and not an “exclusive or” unless specifically indicated to the contrary. Reference to a “first” component does not necessarily require that a second component be provided. Moreover, reference to a “first” or a “second” component does not limit the referenced component to a particular location unless expressly stated. The term “based on” is intended to mean “based at least in part on.”

[0211] The claims may be drafted to exclude any element which may be optional. As such, this statement is intended to serve as antecedent basis for use of such exclusive terminology as “solely”, “only”, and the like in connection with the recitation of claim elements, or the use of a “negative” limitation.

[0212] Where a range of values is provided, it is understood that each intervening value, to the tenth of the unit of the lower limit unless the context clearly dictates otherwise, between the upper and lower limits of that range is also specifically disclosed. Each smaller range between any stated value or intervening value in a stated range and any other stated or intervening value in that stated range is encompassed within embodiments of the present disclosure. The upper and lower limits of these smaller ranges may independently be included or excluded in the range, and each range where either, neither, or both limits are included inthe smaller ranges is also encompassed within the present disclosure, subject to any specifically excluded limit in the stated range. Where the stated range includes one or both of the limits, ranges excluding either or both of those included limits are also included in the present disclosure.

[0213] All patents, patent applications, publications, and descriptions mentioned herein are hereby incorporated by reference in their entirety for all purposes as if each individual publication or patent were specifically and individually indicated to be incorporated by reference and are incorporated herein by reference to disclose and describe the methods and / or materials in connection with which the publications are cited. None is admitted to be prior art.

Claims

WHAT IS CLAIMED IS:

1. A sensor for detecting an analyte, the sensor comprising: a sensing surface, the sensing surface comprising a nanoporous layer defining a plurality of nanopores, wherein: the nanoporous layer is functionalized with a plurality of recognition elements in contact with the nanoporous layer and disposed in the plurality of nanopores, and each recognition element of the plurality of recognition elements is configured such that upon binding of the recognition element to the analyte, a detectable signal from the recognition element is generated; and a polymer coating overlaying the nanoporous layer and the recognition elements disposed thereon, wherein: the nanopore and the polymer coating are configured to reduce an amount of interferents from contacting the recognition elements.

2. The sensor of claim 1, wherein the nanoporous layer comprises gold.The sensor of claim 1, wherein the nanoporous layer is selected from the group consisting of silver, platinum, titanium, copper, titanium oxide, silica, ceramic material, graphene, graphene-oxide, molybdenum disulfide nanosheets, indium tin oxide (ITO), and cobalt oxyhydroxide nanosheets.

4. The sensor of claim 1, wherein the nanoporous surface comprises nanopores having an average diameter in a range of 1 to 50 nm.

5. The sensor of claim 1, wherein the nanoporous surface comprises nanopores at a density in a range of 107-1013nanopores per cm2.

6. The sensor of claim 1, wherein the recognition element comprises an aptamer or analyte-responsive molecular switches.The sensor of claim 1, wherein the detectable signal is an electrical signal.

8. The sensor of claim 1, wherein the detectable signal is an optical signal.

9. The sensor of claim 1, wherein the polymer coating comprises a material selected from the group consisting of polyethylene glycol, a polysaccharide, a polyoxazoline, a polyglycerol fatty acid ester, a poly(thioglycidyl glycerol), a polycarbonate, a propylene glycol, a polyaminoacid, a polyvinylpyrrol, a zwitterionic polymer, a polybetaine, a polyester, and a polysarcosine.

10. The sensor of claim 9, wherein the polymer coating comprises polyethylene glycol.

11. The sensor of claim 1, wherein the polymer coating comprises polyacrylamide.

12. The sensor of claim 1, wherein the polymer is branched.

13. The sensor of claim 12, wherein the polymer comprises units having 6 to 8 branches.

14. The sensor of claim 12, wherein each branch of the polymer has a length of at least 1 nm.

15. The sensor of claim 1, wherein the polymer coating comprises a copolymer or homopolymer.

16. The sensor of claim 1, wherein the polymer coating comprises a hydrogel.

17. The sensor of claim 1, wherein the polymer coating comprises a polymer with a molecular weight from 2,000 to 20,000 Da.

18. The sensor of claim 1, wherein the polymer coating comprises terminal functional groups.

19. The sensor of claim 1, wherein the polymer coating comprises a polymer having a hydrodynamic size less than an average pore size of the nanopore.

20. The sensor of claim 1, wherein the polymer coating has an average height in a range from 5 to 10 nm.

21. The sensor of claim 1, wherein the nanopore has an effective diameter in a range from 10 to 50 nm.

22. The sensor of claim 1, wherein a channel free of polymer is present in the nanopores and has an effective diameter in a range from 5 to 10 nm.

23. The sensor of claim 1, wherein the nanoporous layer overlays a support or substrate.

24. The sensor of claim 23, wherein the substrate comprises polyimide.

25. The sensor of claim 23, wherein the support or substrate is characterized by a surface having an area of in a range from 5,000 to 10,000 μm2.

26. The sensor of claim 23, wherein: the nanoporous layer overlays and is in contact with a gold layer, the gold layer overlays and is in contact with a metal layer comprising titanium and gold, and the metal layer overlays and is in contact with a polymer substrate.

27. The sensor of claim 1, wherein the analyte is selected from the group consisting of a cell-free nucleic acid molecule, a drug or drug metabolite, a lipid, a protein, a peptide, a polysachharide, a pathogen, a cell, a nuclease, a pollutant, and a contaminant.

28. The sensor of claim 1, wherein the nanoporous layer is in electrical connection with elements for electrochemical interrogation of the redox state of the nanoporous layer.

29. The sensor of claim 28, further comprising a logic system configured to translate the detectable signal to an analyte concentration.

30. The sensor of claim 1, wherein the polymer coating comprises a layer covering openings of the plurality of nanopores.

31. The sensor of claim 1, wherein the polymer coating comprises a layer disposed conformally in the plurality of nanopores.

32. The sensor of claim 1, wherein the polymer coating comprises a portion filing the nanopore in one or more of the nanopores of the plurality of nanopores.

33. A sensor system comprising a sensor of any one of claims 1 to 32 and configured for deployment in vivo.

34. A method of making a sensor for detecting an analyte, the method comprising: functionalizing a nanoporous layer with a plurality of recognition elements; and depositing a polymer coating over the nanoporous layer.

35. The method of claim 34, further comprising forming the nanoporous layer by: depositing an alloy of a first metal and a second metal over a substrate; and removing the second metal to form a nanoporous layer comprising the first metal.

36. The method of claim 35, wherein removing the second metal comprises etching the second metal with nitric acid.

37. The method of claim 35, further comprising: depositing the alloy onto a first metal layer comprising titanium and gold.

38. The method of claim 37, further comprising: depositing the first metal layer onto a second metal layer comprising gold.

39. The method of claim 38, further comprising depositing the second metal layer onto a polymer substrate.

40. The method of claim 35, further comprising adjusting sizes of channels in the nanoporous layer by thermal engineering.

41. A method of detecting an analyte, the method comprising: contacting the sensor of any one of claims 1 to 32 to a sample comprising thebinding the analyte to a recognition element of the plurality of recognition elements to generate a detectable signal; and detecting the detectable signal to detect presence of the analyte.

42. The method of claim 41, wherein the sensor comprises an electrochemical sensor and the signal is an electrical signal.

43. The method of claim 41, wherein the sensor comprises an optical sensor and the signal is an optical signal.

44. The method of claim 41, wherein contacting the sensor to the sample is in vivo. 45 The method of claim 41, further comprising deploying the sensor is in a human subject.

46. The method of claim 41, further comprising releasing the analyte from the recognition element.

47. The method of claim 41, further comprising: binding a plurality of analytes to the plurality of recognition elements, wherein the detectable signal is generated from the binding of the plurality of analytes to the plurality of molecular switches, measuring the detectable signal, and determining an amount of analytes using the measured detectable signal.

48. The method of claim 41, further comprising: disposing the sensor in an environment, wherein the environment would degrade the molecular switch within seven days if the polymer coating were absent.

49. The method of claim 48, wherein the environment is a chemical reactor, an industrial reactor, a manufacturing facility, or a food production reactor.

50. The method of claim 48, further comprising: detecting the analyte at least seven days after disposing the sensor in the environment.

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