Electrochemical sensor systems and methods with reduced signal loss

Abstract

Electrochemical aptamer (EAB)-based sensors support high frequency real-time monitoring of analytes in vivo. Methods of reducing sensor drift during in vivo placement are described. EAB sensors with non-natural nucleic acids can be used to reduce sensor signal drift. The EAB sensor improves the stability and reduces the signal loss.

Description

[0001] Cross-references of related citations This application claims priority to U.S. Provisional Patent Application No. 63 / 506,560, filed June 6, 2023, entitled "Electrochemical Sensor with Reduced Signal Loss". The entire disclosure of U.S. Provisional Patent Application No. 63 / 506,560 is incorporated herein by reference for all purposes.

[0002] Federally Funded Research Statement This invention was completed with government support under licenses AI164483 and R01AI145206 granted by the National Institutes of Health in the United States. The government holds certain rights to this invention.

[0003] sequence list This application contains a sequence of documents submitted electronically in XML format, all of which are hereby incorporated by reference. The XML copy was created on May 20, 2024, named 08507PCT.xml, and is 7 kilobytes in size. Technical Field

[0004] This invention relates to an electrochemical sensor incorporating nucleic acids, configured to specifically detect analytes in biological fluids. The sensor output signal is configured to resist unwanted signal loss, thus enabling the sensor to monitor analytes over extended periods. Background Technology

[0005] Electrochemical aptamer (EAB) based sensors are known for detecting target analytes in biological fluids. The aptamer portion of the sensor is an oligonucleotide with a defined base sequence that selectively interacts with the target analyte. In one design, the aptamer is surface-coupled to the working electrode, and a redox reporter is coupled to the free end of the aptamer. The binding of the analyte to the aptamer causes a conformational change in the aptamer, bringing the redox reporter closer to the electrode. This movement, in turn, increases the electron transfer rate (ket) between the redox reporter and the electrode. This change in ket reflects the target analyte concentration in real time without the need for the addition of exogenous reagents.

[0006] EAB sensors can perform second- and / or sub-second resolution measurements of a variety of drugs and metabolites in situ within the body, enabling closed-loop, feedback-controlled drug delivery. For in vivo applications, the sensor may include microneedles or wires coated with aptamers as working electrodes. These microneedles or wires can be inserted through the skin surface, allowing the aptamer-coated portion to contact the biological fluid in the subcutaneous tissue. Alternatively, the sensor may be placed in another internal cavity (e.g., a vein). The sensor may also include counter and reference electrodes. These electrodes may also be in the form of microneedles or wires similarly inserted under the skin.

[0007] Sensor electrodes can remain in situ for minutes, hours, or even days, providing clinically valuable information about the amount of analytes in bodily fluids during this time. The total amount of analytes in the atmosphere can be estimated based on reasonable inference. In this way, EAB sensors can provide clinically relevant information about the amount of exogenous analytes (such as drugs) or endogenous analytes (such as hormones) in a subject's body. This information can be used for disease diagnosis, treatment, and / or monitoring. For example, when the analyte is a drug, the information provided by the sensor can be used to optimize drug therapy. For instance, the amount of drug can be monitored to ensure its level remains above the minimum effective concentration but below concentrations that may cause toxicity. Drug dosage can be adjusted to ensure that appropriate levels are maintained.

[0008] Interrogating an EAB sensor for target analyte detection requires the application of a potential waveform, and the current output generated by the working electrode is used to determine the amount of target analyte in solution. Square wave voltammetry (SWV) can be used for in vivo detection of analytes because it can correct for losses (also known as “drift”) in the current output. To explain this, the amount of analyte reported by an EAB sensor tends to drift downwards over time. This drift can be corrected by measuring consecutive square wave voltammetry at two different frequencies using a method called kinetic differential measurement (KDM). KDM uses the difference between the relative SWV measurements obtained at the two frequencies to subtract and correct for the drift.

[0009] While drift correction methods are effective, after a period of time, the sensor's current output will inevitably decrease to or below the noise floor. At this point, there is no detectable signal above the system's inherent noise level, and the sensor can no longer function in target analyte detection. It is generally accepted in the art that the potential transferred to the sensor during interrogation can cause the aptamer to detach from the working electrode surface. Contamination of the working electrode by biological materials (such as proteins, lipids, and whole cells) in the analyzed biological fluids is also considered a contributing factor.

[0010] Signal drift, leading to eventual signal loss, negatively impacts the usability of EAB sensors, particularly for in vivo clinical applications where accurate measurement of analyte amounts over extended periods is crucial. As an example of such an application, a sensor might be configured to detect antibiotics in situ within a subject via an electrode that contacts plasma. Prolonged antibiotic monitoring may be necessary to ensure the dosage regimen maintains drug concentrations above the minimum inhibitory concentration (MIC) while avoiding toxic concentrations. Typically, antibiotic monitoring lasts at least 24 hours, and current sensors cannot maintain signals for such a long period without significant signal loss and a corresponding decrease in signal-to-noise ratio. Signal loss can be mitigated by continuous sensor replacement, but this approach is costly. Repeated sensor replacement requires constant interruption of the subject. Furthermore, sensor replacement (and calibration) takes time and may result in data loss during transitions.

[0011] The discussion of documents, laws, materials, devices, articles, etc., included in this specification is for the purpose of providing background information for the invention. This is not to imply or represent that any or all of these matters constitute part of the prior art, nor is it common general knowledge in the field relating to the invention prior to the priority date of each provisional claim of this application. Summary of the Invention

[0012] This section summarizes and describes in detail below EAB sensors with improved performance, such as (but not limited to) sensors with reduced signal loss. Some embodiments include a working electrode for an electrochemical sensor, the working electrode including a conductive element and an associated analyte sensing element configured to specifically interact with a target analyte, wherein the sensing element includes a non-natural nucleic acid coupled to a redox reporter.

[0013] In some embodiments, non-natural nucleic acids cannot be read and / or replicated by any natural mammalian cell.

[0014] In some embodiments, non-natural nucleic acids are more resistant to degradation of biological fluid components compared to similar natural nucleic acids.

[0015] In some embodiments, similarity relates to any of the following: length, base sequence, secondary structure, tertiary structure, and ability to interact with the target analyte.

[0016] In some embodiments, the biological fluid includes nucleases.

[0017] In some embodiments, non-natural nucleic acids are formed from single strands.

[0018] In some embodiments, non-natural nucleic acids are polymers.

[0019] In some embodiments, non-natural nucleic acids comprise 10 to 100 subunits. In some embodiments, non-natural nucleic acids are natural deoxyribonucleic acid or chemical variants of natural ribonucleic acid.

[0020] In some embodiments, the chemical variant is a sugar backbone and / or one or more bases.

[0021] In some embodiments, non-natural nucleic acids are prepared by humans or with human assistance.

[0022] In some embodiments, the non-natural nucleic acid is a heterologous nucleic acid (XNA) or a peptide nucleic acid (PNA).

[0023] In some embodiments, non-natural nucleic acids are DNA or RNA aptamers with altered chemical structures.

[0024] In some embodiments, the non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the interaction capability of the non-natural nucleic acid is at least 50%, 60%, 70%, 80%, 90%, 100%, 110%, 120%, 130%, 140%, or 150% of that of the natural nucleic acid. In some embodiments, the non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the non-natural nucleic acid has a sensitivity, precision, or specificity for the recognition of the target analyte that is at least 50%, 60%, 70%, 80%, 90%, or 100% that of the natural nucleic acid.

[0025] In some embodiments, non-natural nucleic acids are altered forms of natural nucleic acids capable of specifically interacting with the target analyte, and the non-natural nucleic acids have an uncorrected signal loss of less than 90%, 80%, 70%, 60%, or 50% that of natural nucleic acids. In some embodiments, the working electrode is part of an electrochemical sensor having an uncorrected signal drift rate, and the uncorrected signal drift rate is determined or averaged over a period of at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

[0026] In some embodiments, a non-natural nucleic acid binds to a conductive element at the first end.

[0027] In some embodiments, redox compounds bind to the second end of non-natural nucleic acids. In some embodiments, the conductive element includes a skin-penetrating portion having a non-natural nucleic acid bound thereto.

[0028] In some embodiments, the conductive element is a needle, microneedle, or wire.

[0029] Some embodiments include an electrochemical sensor device comprising a working electrode and a counter electrode as described in any one of claims 1 to 21.

[0030] In some embodiments, the electrochemical sensor device includes a reference electrode.

[0031] In some embodiments, the electrochemical sensor device has an associated retainer configured to keep the working electrode in contact with the subject's bodily fluids.

[0032] In some embodiments, the retainer is configured to keep the working electrode in contact with the subject's bodily fluids for at least 10, 20, 30, 40, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

[0033] In some embodiments, the electrochemical sensor device has an associated housing configured to enclose the power supply and / or electronic equipment used for the operation of the sensor.

[0034] Some embodiments include a method for monitoring a target analyte in a subject's biological fluids, comprising contacting the working electrode of any of claims 1 to 21 with the biological fluids for a period of time.

[0035] In some embodiments, the working electrode is in contact with biological fluids that remain in situ within the subject's body during the duration of the method.

[0036] In some embodiments, the biological fluid is blood or tissue fluid.

[0037] In some embodiments, the biological fluids have not been removed from the subject's body.

[0038] In some embodiments, the time period is at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

[0039] Additional embodiments and features are set forth in part in the description which follows, and will become apparent in part to those skilled in the art upon examination of the specification, or may be learned by practice of the disclosure. The nature and advantages of this disclosure can be further understood by referring to the remainder of the specification and drawings, which form a part of this disclosure. Attached Figure Description

[0040] This description will be more fully understood with reference to the following accompanying drawings, which are presented as embodiments of the invention and should not be construed as a complete statement of the scope of the invention, wherein: Figure 1A This is a schematic diagram of the response of an EAB sensor based on existing technology to the presence of a target analyte.

[0041] Figure 1B It shows 37 according to the prior art C. The EAB sensor, which uses DNA sequences, will be immersed in whole blood (in vitro).

[0042] Figure 1C The arrangement of an EAB sensor employing a DNA sequence in the jugular vein (in vivo) of a rat, according to the prior art, is shown.

[0043] Figure 1D and Figure 1E The graphs showing the changes in normalized in vitro and in vivo current over time obtained from SWV collected at two frequencies according to the prior art are shown.

[0044] Figure 1F and Figure 1G The graphs showing the changes of KDM signals over time in vitro and in vivo according to the prior art are shown.

[0045] Figure 2A The EAB sensor employing a natural DNA molecule, according to one embodiment, and its destruction by reagents in biological fluids are illustrated schematically and chemically. Figure 2B The illustration schematically and chemically shows a configuration according to one embodiment having a... Figure 2A EAB sensors, whose aptamers exhibit similar non-natural nucleic acid functions in their interaction with target analytes, although they are resistant to... Figure 2A The degradation of the reagent. Figure 2C A graph showing the change of associated sensor output signal over time according to one embodiment is shown, comparing the output signals of a sensor with natural nucleic acid as a sensing element and a sensor with non-natural nucleic acid as a sensing element in the presence of a nuclease.

[0046] Figure 3A The variation of KDM signal with tobramycin according to one embodiment is shown, wherein natural and non-natural nucleic acids serve as sensing elements.

[0047] Figure 3B The diagram shows the normalized current versus time curve of a sensor with separate natural and non-natural nucleic acids as sensing elements when the sensor is placed on the left and right jugular veins of a single rat excited with an aminoglycoside antibiotic, according to one embodiment. Figure 3C An embodiment is shown. Figure 3B The KDM signal of the sensor used changes over time. Figure 3D An embodiment is shown. Figure 3B The concentration of aminoglycosides in the sensor used changes over time. Figure 3E The signal of a DNA and Ome-RNA aptamer sensor according to one embodiment is shown.

[0048] Figure 3F and Figure 3G Charge transfer kinetics of a DNA aptamer sensor and an OMe-RNA aptamer sensor according to one embodiment are shown respectively. Figure 3H Normalized signals of a DNA aptamer sensor and an OMe-RNA aptamer sensor according to one embodiment are shown.

[0049] Figure 4A The illustration schematically depicts whole blood in vitro according to one embodiment, in which an EAB sensor is immersed, the EAB sensor having natural or non-natural nucleic acid as a sensing element. The graph on the right shows the change in relative sensor signal strength over time for each sensor type.

[0050] Figure 4B The diagram schematically illustrates an EAB sensor according to one embodiment, which has natural or non-natural nucleic acids as sensing elements disposed in the jugular vein of a rat. A non-natural nucleic acid sensor is inserted prior to the natural nucleic acid sensor. The graph on the right shows the change in relative sensor signal intensity over time for each sensor type and for two treated rats. Figure 4C A sensor according to one embodiment is schematically illustrated, having natural or non-natural nucleic acid as a sensing element disposed in the jugular vein of a rat. A natural nucleic acid sensor is inserted before the non-natural nucleic acid sensor. The graph on the right shows the change in relative sensor signal intensity over time for each sensor type and for two treated rats.

[0051] Figures 5A to 5F A series of charge transfer versus frequency plots (Lovric plots) over 5 hours according to one embodiment are shown, illustrating the temporal evolution of electron transfer dynamics of an EAB sensor placed under various conditions. Figure 5A The in vitro Lovri sensor, made from natural nucleic acids and excited by DNase, is shown. picture. Figure 5B The Lovri sensor, made from non-natural nucleic acids and excited by DNase, was demonstrated in vitro in PBS buffer. picture. Figure 5C This demonstrates the effect of a sensor made from natural nucleic acids and excited by DNase in whole blood in vitro on Lovri. picture. Figure 5D This demonstrates the Lovri sensor, made from non-natural nucleic acids and excited by DNase, in whole blood in vitro. picture. Figure 5E Lovri demonstrates a sensor made from natural nucleic acids and used in a whole blood environment in vivo. picture. Figure 5F Lovri demonstrates a sensor made from non-natural nucleic acids and used in a whole blood environment in vivo. picture.

[0052] Figure 6 A top perspective view of a microneedle embedding device according to one embodiment is shown. This embodiment relies on the user providing the impetus to insert the microneedles into the skin. The arm is shown in the first position presented to the user before the microneedles are inserted into the skin.

[0053] Figure 7A An embodiment is shown. Figure 6 The lower perspective view.

[0054] Figure 7B An embodiment is shown. Figure 6 The upper perspective view.

[0055] Figure 8 An embodiment is shown. Figure 6 The lower perspective view shows more fully the removable flexible layer, which has been removed to expose the dermatologically acceptable adhesive. Figure 9The following perspective view illustrates an embodiment. Figure 6 The microneedle embedding device has a removable flexible layer that exposes a dermatologically acceptable adhesive. Figure 10 The following perspective view illustrates an embodiment. Figure 9 A microneedle embedding device in which the microneedles are in an extended position for embedding in the subject's skin.

[0056] Unless otherwise stated herein, when used in different figures, features in the figures labeled with the same numbers are considered to be the same features, or at least features that are functionally similar.

[0057] The accompanying drawings are not prepared to any particular scale or size and are not intended to be a completely accurate representation of the various embodiments. Detailed Implementation

[0058] Having considered this description, those skilled in the art will understand how the invention can be implemented in various alternative embodiments and applications. However, although various embodiments of the invention will be described herein, it should be understood that these embodiments are by way of example only and not limitation. Therefore, the description of various alternative embodiments should not be construed as limiting the scope or breadth of the invention. Furthermore, statements of advantages or other aspects apply to specific exemplary embodiments and not necessarily to all embodiments, or indeed to any embodiment covered by the claims. In the description and claims of this specification, the word "comprising" and its variations, such as "having" and "including", does not imply the exclusion of other additives, ingredients, integers or steps.

[0059] Throughout this specification, the terms "an embodiment" or "an embodiment" refer to a specific feature, structure, or characteristic described in connection with that embodiment, which is included in at least one embodiment of the invention. Therefore, the phrases "in one embodiment" or "in an embodiment" appearing in various places throughout the specification do not necessarily refer to the same embodiment, but may refer to the same embodiment.

[0060] The term "subject" is used to refer to animals (including human and non-human animals) to which this invention can be applied.

[0061] As used herein, “body fluid” can be any biological body fluid of the subject, including but not limited to tissue fluid (ISF), blood, saliva, tears, lactation, nasal secretions, tracheal secretions, bronchial secretions, alveolar secretions, gastric juice, gastric contents, glandular secretions, vaginal secretions, uterine secretions, prostatic secretions, semen, urine, sweat, cerebrospinal fluid, glomerular filtrate, liver secretions, bile, or exudate, any of which come into contact with the electrodes of the present invention during use.

[0062] This invention is based, at least in part, on the inventors' discovery that non-natural nucleic acids can be used as detection elements. Non-natural nucleic acids possess similar functions to natural nucleic acid aptamers (such as DNA and RNA), but exhibit significantly lower signal loss rates when used in in vivo EAB sensors. As a result, the sensor can provide signal output above the noise floor over extended periods.

[0063] Many embodiments suggest that the degradation of aptamers in EAB sensors by reagents present in biological fluids is a significant factor and may be a major cause of signal loss in vivo. In several embodiments, this reagent is a nuclease. These findings reveal a previously unknown root cause of the signal drift problem. Based on this discovery, the inventors found that non-natural nucleic acids are resistant to signal drift in the context of EAB sensors, while being similar to (and in some cases better than) natural nucleic acid aptamers in detecting analytes.

[0064] These findings (i.e., signal loss is due to aptamer loss caused by interrogation current and contamination of biological materials) are inconsistent with the generally accepted dogma in the field.

[0065] Significant advantages are offered in EAB sensors for applications where aptamer degradation is problematic. In addition to the advantages described above for in vivo applications, certain in vitro applications may also benefit. In some embodiments, electrochemical sensors for monitoring the presence of analytes in industrial processes involving biological materials can resist degradation during process operation. Some embodiments provide in vitro testing of clinical samples in pathology facilities. Although the sensor may only be in contact with the clinical sample temporarily, significant degradation can still occur, leading to unstable output signals. Furthermore, the sensor can be cleaned and used with multiple samples, in which case the sensor's lifespan will be extended due to its resistance to degradation.

[0066] As used herein, the term "non-natural nucleic acid" is intended to include polymers that have biosimilarity to natural nucleic acid polymers (such as DNA or RNA) but with altered chemical structures not found in nature. Due to these structural alterations, non-natural nucleic acids may be more resistant to degradation (such as the breaking of chemical bonds) caused by agents in biological fluids such as blood and ISF than natural nucleic acids. These agents can be nucleases, such as DNase or RNase. Non-natural nucleic acids may originate from naturally occurring nucleic acids, but their chemical structure has been altered, thus the chemical structure is considered non-natural. More typically, non-natural nucleic acids are synthesized de novo in a modified form.

[0067] The term "non-natural nucleic acid" is not intended to include nucleic acids synthesized by humans or with human assistance that still possess a natural chemical structure. Although these molecules are not products of nature, they have the same chemical structure as nucleic acids found in nature.

[0068] Non-natural nucleic acid molecules that can be used in the context of this invention (i.e., as sensing elements associated with the working electrode of an EAB sensor) can be modified forms of aptamers. For example, DNA aptamers for known analytes (such as antibiotic drugs) can be chemically modified to retain the ability to detect the analyte while resisting degradation of the reagent in biological fluids. Non-natural nucleic acids can be oligomers with a non-natural backbone, which are molecular analogs of DNA or RNA. Examples of oligomers with a non-natural backbone include, but are not limited to, 2'-fluoroarabinonucleotide (FANA), 2'-O-methylRNA, locked nucleic acid (LNA), and threononucleotide (TNA). In general, these oligomers with non-natural backbones are referred to as heteronucleotides (XNAs). Because XNAs are not produced in nature, they are generally highly resistant to enzymatic degradation. In some embodiments, non-natural nucleic acids can be peptide nucleic acids (PNAs).

[0069] In addition to confirming the altered chemical structure of stability in biological fluids, the non-natural nucleic acids of this invention may possess one or more characteristics of prior art aptamers. These characteristics include, but are not limited to, length, base sequence (primary structure), secondary structure, and tertiary structure. Aptamers are small (typically 20 to 60 nucleotides) RNA or DNA oligonucleotides formed from single strands capable of binding to target analytes with high affinity and specificity. Aptamers can be considered nucleotide analogs of antibodies, but aptamer production is an in vitro, cell-free process, which is easier and cheaper than producing antibodies through cell culture or in vivo methods.

[0070] In the context of this invention, one method for identifying aptamers is to first identify natural DNA or RNA aptamers using methods such as SELEX, and then modify the identified aptamers to have a non-natural chemical structure. Alternatively, methods such as SELEX can be modified first using enzymes configured for the synthesis and amplification of non-natural nucleic acids. Aptamers can be selected from combinatorial libraries containing a large number (up to 10¹⁸) of different oligonucleotides. Although RNA aptamers offer greater structural diversity compared to DNA aptamers, their application is complicated by stability issues in the presence of RNases, high temperatures, and unfavorable pH values.

[0071] The selection of aptamers selective for a given analyte can be facilitated through a process called SELEX (Spiritual Evolution via Exponential Enrichment of Ligands). This process can be viewed as two alternating phases. In the first phase, a library of oligonucleotides is amplified to the desired concentration via polymerase chain reaction (PCR). For RNA aptamers, single-stranded oligonucleotides are generated by in vitro transcription of double-stranded DNA using T7 RNA polymerase. For DNA aptamers, a pool of oligodeoxyribonucleotides is generated by strand separation of the double-stranded PCR product. In the second phase, the amplified product is incubated with the target analyte and the oligonucleotides, which bind to the analyte used in the next round of SELEX.

[0072] Through intense competition for binding sites, the separation of oligonucleotides with higher affinity for the target drug and the removal of unbound oligonucleotides are achieved. The selection pressure increases with each round of SELEX. After approximately 5 to 15 rounds, the oligonucleotide pool can be sufficiently enriched using aptamers with the strongest affinity for the target analyte. Once determined to be useful, non-natural nucleic acids can be bound to the working electrode of the EAB sensor. The working electrode may have at least one associated counter electrode and at least one associated reference electrode. Each working electrode may have a dedicated counter electrode; however, in some embodiments, the counter electrode is shared among some or all assembled working electrodes. Each working electrode may have a dedicated reference electrode; however, in some embodiments, the reference electrode is shared among some or all assembled working electrodes.

[0073] In many embodiments, a useful EAB sensor within the context of this disclosure can be an volt-ampere sensor, a timing current sensor, or an impedance sensor. In a volt-ampere sensor, a potential waveform is applied to the sensor interface, and the resulting current response is recorded. In a timing sensor, a step potential is applied, and the resulting time-evolution current response is recorded. In an impedance sensor, a sinusoidal potential waveform is applied, and the resulting sinusoidal current response is recorded. EAB sensors are typically voltammetric, with an aptamer (or, according to many embodiments, a non-natural nucleic acid) bound to the working electrode. Gold can be used as the probe surface of the working electrode. The aptamer has an associated redox-active species that serves as a reporter. The redox reporter can be, but is not limited to, methylene blue. Upon binding to a target (e.g., a drug), the aptamer undergoes a conformational change, bringing the redox reporter closer to the working electrode surface. This increased proximity increases electron transfer from the redox reporter to the electrode. This increased electron transfer rate helps the potentiostat detect changes in Faraday current. EAB sensors can be incorporated into circuits with a reference electrode. The reference electrode is the site of a known chemical reaction with a known redox potential. For example, a reference electrode based on the silver-silver chloride (Ag / AgCl) redox pair has a fixed and known potential, forming a reference point for measuring the redox potential of the working electrode. The circuit also typically includes a counter electrode, which serves as the cathode or anode of the working electrode. Since no current flows through the reference electrode (due to the impedance of the potentiostat), any current generated is attributable to both the working and counter electrodes. The current is measured as a function of the potential of the interrogating electrode relative to the reference electrode. The potential difference generates a current in the circuit, thus producing an output signal. This signal quantifies target binding based on the stoichiometric electron transfer that ideally binds to the target.

[0074] This device, once assembled, is particularly well-suited for use as a wearable device, thereby allowing measurements to be taken while the subject is engaged in normal activities and / or for extended periods. In several embodiments, the wearable device may be a collar, bracelet, or other suitable jewelry, watch, clothing, strap, adhesive, or patch. Those skilled in the art will understand that devices (e.g., microanchors) may be provided to assist in adhering to and / or securing the wearable device to the subject during use. In some embodiments, the wearable device may include a housing structure that includes one or more other components, such as (but not limited to) an electronic processing unit. The electronic processing unit is configured to communicate directly or indirectly with at least one conductive element (such as an electrode) and will typically include any one or more of a power supply, a data processing unit, an analog front end, and a wireless transmitter. In several embodiments, the housing structure may be configured to at least partially enclose the device, with electrodes (such as microneedles) exposed from the plane of the housing structure. The electrodes may be protected by a protective cover that can be removed before use to expose the protruding electrodes.

[0075] In many embodiments, the device may also include means for monitoring the temperature or pH of bodily fluids, wherein the validity of the output depends on this, or the operation or output may be regulated.

[0076] In several embodiments, the housing structure can be configured to surround and couple to the device via any suitable mechanism. Examples of mechanisms include, but are not limited to, electromagnetic coupling, mechanical coupling, adhesive coupling, magnetic coupling, etc. In some embodiments, the coupling mechanism allows the device and housing structure to be connected and disconnected, which allows the housing structure and other components to be reused, while the device can be discarded and replaced with another device as needed.

[0077] In many embodiments, the wearable device may also include a computer program product executable as a software application residing on a mobile communication device in communication with an electronic processing unit, wherein the computer program product is capable of controlling one or more of the following: (i) detection of electrochemical measurements performed on an electrode-based platform, (ii) data analysis, (iii) data transmission, (iv) device configuration, and (v) device power management. Examples of suitable mobile communication devices include, but are not limited to, smartphones, smartwatches, tablets, smart glasses, laptops, or other personal computers. In some embodiments, the device itself includes a processor with program instructions configured to drive onboard functions (such as voltammetry measurements) and transmit output to a remote device via a wireless module (such as a Bluetooth™ module).

[0078] In some embodiments, the working electrode or any other electrode may be a wire, needle, microneedle, electrode array, or microneedle array that contacts the subject's ISF, blood, or any other relevant body material. Microneedles and / or microneedle arrays are preferably used for transdermal applications (requiring skin puncture to contact the ISF).

[0079] While the specific geometry of the electrodes for transdermal applications is optimized to penetrate the stratum corneum for reliable skin penetration, electrodes according to many embodiments can be fabricated in a variety of shapes and geometries. In some embodiments, the device can be configured to be pushed into the subject's skin to facilitate the electrodes' penetration of the stratum corneum and into the skin layers. For non-human applications, the stratum corneum can be replaced by similar or even dissimilar layers on the subject's surface. Typically, each electrode has the shape of a protruding tip structure extending from the base. The electrode typically extends roughly vertically from the base. The protruding structure of each electrode can be any needle-like shape. In some embodiments, the protruding structure may taper smoothly from a base to form a tip (e.g., tapered), may have multiple sides extending from the base (these sides converge to form the tip (e.g., pyramidal or triangular prism)), the protruding structure may be tapered only in one dimension, or may have a base including curved sides of a relatively constant diameter, which is segmented to form the tip (e.g., a portion of a cylinder). Typically, the tip will be sharp. The electrode may or may not include shape variations along its length. Furthermore, any edges or sides of the shape may be beveled, curved, or rounded. In some embodiments, the electrode is cone-shaped or pyramidal, such as a triangular pyramid, square pyramid, or hexagonal pyramid. In other embodiments, the electrode is tetrahedral or triangular prism-shaped. In further embodiments, it can take the shape of a rocket, turret, arrowhead, spike, or spear.

[0080] It should be understood that a range of other shapes may be used. In some embodiments, the shape may be a truncated circular or elliptical cylinder. Any other shape described herein may or may not be truncated. The term "truncated" as used in this context may refer to a shape cut in a plane parallel to the base, which may be called a parallel truncated shape, or more specifically, a truncated body, or a shape cut at an angle relative to the electrode axis. For angularly truncated shapes, the truncation angle relative to the shape axis is at least about 50° and does not exceed about 75°. It should be understood that the same or different shapes may be provided on a single EAB sensor.

[0081] The outer wall of the microneedle can have a smooth or rough surface and can include surface features such as (but not limited to) protrusions, etchings, serrations, anchors, barbs, etc., which may aid in the adhesion of biological tissue once the electrode penetrates the stratum corneum to anchor it within the subject's body. It should be understood that the ability of the EAB sensor to remain in situ is particularly beneficial, as this ensures continuous measurements at the same site within the subject's body over extended periods. Furthermore, limiting the location of measurements ensures more accurate longitudinal monitoring. In some embodiments, the EAB sensor is configured to remain in situ for at least 1 minute, or at least 1 hour, or at least about 8 hours, or at least about 18 hours, or at least 1 day (about 24 hours), or at least about 3 days, or at least about 4 days, or at least one week. In some applications, it may be necessary or desirable to remain in situ for a month or longer.

[0082] It should be understood that the size of the electrodes and their arrangement on the base can vary depending on the intended application.

[0083] In many embodiments, the length of the electrode can be at least greater than the thickness of the stratum corneum and penetrate the skin layer to a depth of at least 100 µm to place the electrode in biological tissue to contact the subject's bodily fluids. In some embodiments, the length will be at least about 10%, or at least about 20%, or at least about 50%, or at least about 75%, or at least about 100% greater than the thickness of the stratum corneum. In some embodiments, the length is less than about 1500 µm, or less than about 1000 µm, or less than about 500 µm, or greater than about 100 µm, or greater than about 50 µm, or greater than about 20 µm, or greater than about 10 µm. In some embodiments, the length is between about 100 µm and about 1000 µm.

[0084] In some embodiments, the electrodes are arranged in layers, so that not all electrodes have the same length.

[0085] The substrate width of the electrode may be at least less than about 50% of its length, or less than about 25% of its length, or less than about 20% of its length, or less than about 15% of its length, or less than about 10% of its length, or less than about 5% of its length. In some embodiments, the substrate width is at least about 100 µm but not more than about 400 µm. In some embodiments, the diameter is about 200 µm or about 300 µm.

[0086] The diameter of the electrode can be at least less than about 50% of the length, or less than about 25% of the length, or less than about 20% of the length, or less than about 15% of the length, or less than about 10% of the length, or less than about 5% of the length. In some embodiments, the diameter is between about 0.5 mm and about 1 mm.

[0087] Compared to another microneedle, a microneedle may penetrate the skin more deeply. Therefore, the two microneedles may terminate at different distances from the skin surface or at different distances from the electrode mount. In some embodiments, the two microneedles are of different lengths. In some embodiments, the microneedles are of the same length, and the mount is configured to allow one microneedle to be axially displaced relative to the other. In some embodiments, the mount may be multi-layered, with a first electrode extending from a first layer and a second electrode extending from a second layer. In many embodiments, the electrodes can be arranged in pairs, groups, or matrices. Paired arrangements will include an even number of electrodes. Grouped arrangements can include one to five groups, each group comprising about four to eight electrodes. Matrix arrangements can include an even or odd number of electrodes, as such arrangements may or may not have the same number of rows and / or columns. In some embodiments, the electrodes are arranged in a matrix selected from combinations of 2x2, 2x3, 2x4, 2x5, 2x6, 3x2, 3x3, 3x4, 3x5, 3x6, 4x2, 4x3, 4x4, 4x5, 4x6, 5x2, 5x3, 5x4, 5x5, 5x6, 6x2, 6x3, 6x4, 6x5, and 6x6. In any arrangement, the electrodes may be spaced less than about 5 mm, less than about 4 mm, less than about 3 mm, less than about 2 mm, less than about 1 mm, or less than about 0.5 mm apart, and the electrodes may be spaced more than about 0.1 mm apart. Space can be measured from the center point of each corresponding electrode to the center point.

[0088] The invention will now be described more fully with reference to the following non-limiting examples.

[0089] Exemplary embodiments The following examples are provided to show those skilled in the art how to make and use the invention in its entirety, and are not intended to limit the scope of what the inventors consider their invention, nor to represent that the following experiments are all or only those experiments performed. Efforts have been made to ensure the accuracy of the figures used (e.g., quantities, temperatures, etc.), but some experimental errors and biases should be taken into account.

[0090] Example 1: Materials and Sources The following materials were used in Examples 2 through 5.

[0091] The in vitro sensors are made of 0.2 mm diameter gold wire (99.99%) insulated with polyolefin heat-shrink tubing (0.05 inch, 0.017 inch, 0.007 inch). For in vitro testing, commercially available Ag|AgCl(s) and commercially available platinum reference electrodes are used. The sensors for in vivo measurements consist of 0.2 mm diameter gold wire, 0.005 inch diameter platinum wire (99.99% purity), and 0.005 mm diameter silver wire (99.9% purity). The insulation used for these sensors is polytetrafluoroethylene heat-shrinkable material (PTFE, 0.02 inch, 0.005 inch, 0.003 inch, black).

[0092] Use sodium hydroxide, 6-mercapto-1-hexanol, tris(2-carboxyethyl)phosphine, sulfuric acid, and DNase. Dilute phosphate-buffered saline (PBS) 20-fold from the stock solution. Use heparinized whole blood. Use tobramycin sulfate. Use T37 methylene blue and HO-C6S-S-C6 modified DNA sequences, as well as other methylene blue and HO-C6S-S-C6 modified sequences, including 2'-O-methyl sequences.

[0093] Example 2: Fabrication of EAB Sensors The in vitro sensors were made by exposing and insulating 3 mm long polyolefin electrodes (see, for example, Leung et al.; ACS Sensors 2021, 6(9), 3340–3347; this disclosure is incorporated herein by reference). After fabrication, these sensors were electrochemically cleaned. The in vivo sensors were fabricated by bundling gold (working electrode), platinum (counter electrode), and silver (reference electrode) wires together in parallel. These wires were insulated from each other using polytetrafluoroethylene heat shrink tubing and bundled together in an interlaced manner. The gold wire was at the bottom with an exposed length of 3 mm, followed by platinum with an exposed length of 6 mm, and finally silver with an exposed length of 1 cm. Once bundled together, this three-electrode sensor was immersed overnight in household bleach with the silver chloride electrode. Subsequently, the three electrodes were rinsed with microporous water prior to electrochemical cleaning. The gold portion of the in vitro and in vivo sensors was electrochemically cleaned in NaOH, followed by roughening in H2SO4 using a CH1040C potentiostat. The cleaning consisted of 1000 cycles of potential cycling between -1.0 V and -2 V at a rate of 2 V / s in 0.5 M NaOH. Subsequently, the gold wires were roughened 32,000 times in 0.5 M H2SO4 with 20 ms pulses at 0 V and 2.2 V to increase the microsurface area of ​​the electrode. The gold electrode was then analyzed by cyclic voltammetry in 0.5 M H2SO4 (1.5 V to -0.35 V, 1 V / s) to determine its electroactive surface area. The electrode was then thoroughly rinsed with microporous water before DNA deposition. At this point, the in vivo sensor was inserted into a 20G catheter. Before depositing DNA onto gold, the HO-C6S-S-C6 modified aptamer was deprotected by mixing 14 μL of 10 mM tris(2-carboxyethyl)phosphine with 2 μL of 100 µM DNA in the dark for 1 hour. After deprotection, the DNA was diluted to 500 nM in PBS. The electrochemically cleaned sensor was then immersed in the DNA solution in the dark and rinsed. After a 1-hour incubation period, the sensor was transferred to a 10 mM 6-mercapto-1-hexanol solution prepared in PBS and stored overnight before use.

[0094]

[0095] Example 3: Electrochemical Measurement In vitro electrochemical measurements were typically performed using a SWV (e.g., see Leung et al.; ACS Sensors 2021, 6(9), 3340–3347; this disclosure is incorporated herein by reference). All sensors were first tested in glass vials at room temperature with 1x PBS before being used for quality checks. They were then transferred to whole blood at 37°C or to PBS containing 2.5 mM CaCl2 and 1.5 mM MgCl2, and electrochemical measurements were immediately initiated. For DNase experiments, measurements were paused after a stable baseline was obtained, the sensor was quickly removed, 5 µg / mL DNase was added, the solution was mixed, and measurements were resumed. The sensor was repeatedly interrogated at multiple square wave frequencies (5, 7, 10, 15, 30, 50, 70, 100, 200, 250, 300, 600, and 1000 Hz), with the square wave frequencies close to the methylene blue electron transfer rate used to monitor sensor degradation.

[0096] An in vivo electrochemical sensor was deployed in the jugular vein of live rats. For experiments involving the monitoring of DNA and 2'-O-methylRNA sensor analog degradation (using the T37 and OMe37 sequences), the sensor was repeatedly interrogated at multiple square wave frequencies, and analysis was performed using square wave frequencies close to the methylene blue transfer rate. For experiments involving the tobramycin detection sensor, SWV was performed alternately at frequencies of 100 Hz and 200 Hz. Signal drift was corrected using these two frequencies to calculate the KDM signal, as shown below:

[0097] in, and These are the peak heights at their respective frequencies. Peak height , t=0 and n=0 indicates the peak height measured before injection. Tobramycin concentrations found in the vein were inversely calculated using Langmuir fitting of the calibration curve.

[0098] Example 4: In vivo surgery and dosage All in vivo experiments were conducted on male Sprague-Dawley rats (4-5 months old). These rats, weighing between 350g and 500g, were kept in pairs in a standard photoperiod chamber (a standard 12:12 photoperiod, with lights turned on at 8:00 AM). They were allowed free access to food and water. Prior to measurements, rats were anesthetized in a plexiglass anesthesia room using 4% isoflurane, and then anesthesia was maintained throughout the experiment using a 2%–2.5% isoflurane / oxygen mixture via a nasal cone. The neck was shaved and dissected to surgically isolate the left and right jugular veins. A small incision was made in each vein using a spring-loaded microscope to allow insertion of an in vivo sensor and a catheter with an infusion line (for heparin delivery). Both the sensor and the infusion line were secured with two sterile 6-0 sutures. Before any recording, 30 units of heparin were infused through the infusion line (jugular vein delivery for tobramycin experiments and femoral vein delivery for aminoglycosides). To administer tobramycin (30 mg / kg) intravenously to rats, a pre-calculated volume of 0.1 M tobramycin sulfate diluted with PBS was injected via the femoral vein using an infusion pump.

[0099] Example 5: Experiment and Discussion of Results EAB sensors enable high-frequency, real-time, in-situ measurements of specific molecules in vivo. A signal is generated in these sensors when a binding-induced conformational change in the aptamer alters the electron transfer rate of the attached redox reporter. This signal transduction mechanism makes EAB sensors independent of the specific chemical structure or reactivity of their targets, as aptamers can bind to a wide variety of small molecule analytes. Therefore, EAB sensors represent the first in vivo molecular monitoring method that can be generalized across a wide range of analytes. Consistent with this, EAB sensors can perform second- and / or sub-second resolution measurements of a variety of drugs and metabolites in situ in the veins, brain, and subcutaneous ISF of live animal subjects, supporting closed-loop, feedback-controlled drug delivery.

[0100] Figures 1A to 1G This demonstrates that, according to existing technology, an electrochemical aptamer (EAB)-based sensor enables continuous, real-time molecular monitoring of untreated biological fluids in vivo and in situ. Figure 1A EAB sensors are shown, each with a gold electrode. A redox reporter-modified aptamer is attached to the gold electrode by forming an alkane thiol on a gold self-assembled monolayer. Upon exposure to the target, the aptamer undergoes a conformational change that alters the electrochemical behavior of the redox reporter (here, methylene blue), producing an easily measurable electrochemical signal. When the EAB sensor is in whole blood at 37°C (… Figure 1B In vitro (commercially sourced bovine blood) or ( Figure 1C When stimulated in vivo (in situ in the rat jugular vein), Figure 1D and Figure 1E The electrochemical peak current it produces is shown to drift downward over time (data collected at the indicated square wave frequency is shown). Figure 1F and Figure 1GThis study demonstrates that this drift can be corrected by the difference between the relative currents observed at square wave frequency pairs where the signal is uniformly drifting; this drift correction method is called kinetic differential measurement (KDM). Good baseline recovery when the sensor is returned to drug-free blood, in vitro, or when the drug is removed from the body, indicates the accuracy of this drift correction method. However, while KDM drift correction ensures that EAB sensors remain accurate, the peak current loss associated with drift reduces their signal-to-noise ratio, ultimately reducing accuracy (e.g., ...). Figure 1G The standard deviation of the signal increased from 0.008 before drug injection to 0.020 four hours after in vivo. The data shown here are from vancomycin detection EAB sensors, which were activated by vancomycin at specified times. Recovery to baseline when the sensors were returned to drug-free blood (in vitro) or when the drug was expelled from the body (in vivo) indicates the accuracy of the KDM correction for drift.

[0101] The limitation of EAB sensors is that when deployed in bodily fluids, they suffer from signal loss over time, an effect that is more pronounced outside the body. Figure 1D ) and in the body ( Figure 1E This will all happen. Because this "drift" can be addressed using the dual-frequency square wave method (… Figure 1F , Figure 1G Precise calibration can be performed using methods such as KDM, or it can be circumvented by interrogation methods such as chronoamperometry or electrochemical impedance spectroscopy, so it does not affect measurement accuracy. However, the loss of signal current reduces the signal-to-noise ratio (regardless of whether it is corrected or circumvented), which ultimately impairs measurement accuracy (see, for example, [reference needed]). Figure 1G (Added noise on the right). Previous researchers have characterized the in vitro drift mechanism of EAB sensors in undiluted whole blood at 37°C. Under these conditions, drift is primarily caused by electrochemically induced loss of the target recognition aptamer (likely due to loss of thiols attached to the gold monolayer) and protein and cellular contamination of the electrode surface, the former of which can be easily avoided by judiciously selecting the potential window used in its interrogation. However, the following experiments demonstrated that the in vivo drift behavior of EAB sensors differs from that in in vitro whole blood drift. By revealing the origin of this difference, numerous embodiments provide hardware and methods that can significantly reduce in vivo drift of EAB sensors. In in vitro body fluids, enzymatic degradation of the aptamers used by the EAB sensor for target recognition is only a minor factor in drift. For example, it has been previously shown that in undiluted whole blood at 37°C, approximately 80% of the significant signal loss of the EAB sensor can be recovered after washing the sensor with concentrated urea, suggesting that the contribution of physical removal of DNA from the surface (if possible during enzymatic degradation) to the drift observed under these conditions is relatively small. Some prior art has also shown that, despite the excellent DNase resistance of non-natural L-enantiomer DNA oligonucleotides (“spiegelmers”), the in vitro blood and serum drift characteristics of sensors made using this polymer are similar to those of natural D-DNA-based sensors deployed under the same conditions. However, it is noteworthy that commercially sourced bovine blood and plasma used in these studies, taken several days prior, may differ significantly from in situ blood from a live animal vein. For example, DNase is known to be active in circulating blood. Given the potentially significant differences between these two experimental conditions, numerous embodiments have provided mechanisms for the drift of EAB sensors placed in the cervical canal of live rats.

[0102] The blood-driven drift of the EAB sensor in vivo differs significantly from that observed in vitro. In several embodiments, a sensor-like construct consisting of unstructured oligonucleotide sequences (i.e., lacking significant internal complementarity) was first employed, made using DNA or 2'-methoxyribonucleic acid (OMeRNA), a relatively enzyme-resistant non-natural heteronucleotide (XNA). Devices using unstructured oligomers of XNA 2' methoxyRNA (OMe RNA) are less susceptible to DNase degradation than devices using equivalent unstructured DNA oligonucleotides. Figures 2A to 2C The invention illustrates a significant reduction in the enzymatic degradation of oligonucleotides by introducing non-natural oligonucleotide XNAs according to one embodiment. To confirm that this applies to the 2'-methoxyribonucleotide backbone (OMe RNA) XNA used herein, some embodiments use DNA (…). Figure 2A ) or OMe RNA ( Figure 2B Oligonucleotides were used to create EAB sensor analogs, with their 3' ends modified with methylene blue and their 5' ends modified with thiols for surface attachment. The sequences used lacked any significant self-complementarity to avoid any influence that might arise from secondary structure differences. Figure 2C In the study, when excited with DNase in PBS (containing the necessary divalent cations and maintained at 37°C), the signal from the device using DNA showed a monotonically decreasing trend. In contrast, the signal from the device using OMe RNA stabilized after a small initial decrease. The shaded areas represent the standard deviation of independently fabricated devices (n=4); no significant difference was observed in the reproducibility of devices fabricated using DNA or XNA. Given this observation and the fact that DNase activity in fresh blood in vivo may be higher than in aged blood in vitro, several examples explored whether using OMe RNA could reduce in vivo drift of the EAB sensor during in vivo placement. However, to test this, the availability of an XNA aptamer is likely required, which binds to a molecular target suitable for in vivo use. That is, the target is non-toxic at concentrations well above the sensor's detection limit and is rapidly cleared from the blood relative to the several-hour timeframe provided by anesthetized animal experiments. While the development of aptamer selection schemes has been applied to the selection of XNA aptamers, no XNA aptamer can simultaneously meet both requirements. For example, most XNA aptamers reported to date bind to proteins, but these are targets that are typically cleared from plasma rather slowly. One small-molecule-binding XNA aptamer can bind to mycotoxins or anthrax toxins, but its toxicity prevents its use in in vivo studies.

[0103] Many embodiments yield novel small-molecule binding XNA aptamers for desired in vivo applications. Several embodiments utilize nucleic acid receptors that bind aminoglycoside antibiotics. Aminoglycoside drugs (including tobramycin, gentamicin, and kanamycin) bind to prokaryotic ribosomal RNA, disrupting translation and inhibiting bacterial growth. Previous reports have included the in vitro folding of 14-base RNA hairpins cleaved from prokaryotic ribosomal RNA into a conformation that binds to these drugs with micromolar affinity. Due to the clinical importance of measuring aminoglycoside drugs (the therapeutic windows of these nephrotoxic and ototoxic drugs are very narrow), the practicality of applying this RNA to an EAB sensor was investigated. (See, for example, A.A.Rowe et al., Anal. Chem. 2010, 827090-7095; the disclosure of which is incorporated herein by reference). Unfortunately, although the resulting sensor reaches the clinically relevant limits of detection, precision, and specificity, the RNA “aptamer” degrades rapidly upon excitation in untreated biological fluids. An unexpected solution to this problem became apparent upon examining the atomic details of the RNA / drug complex. Specifically, (1) a hairpin forms a structure between the B-type of the DNA helix and the A-type of the RNA helix, and (2) aminoglycosides bind in the main groove of the hairpin, on the opposite side of the double helix opposite to the 2' hydroxyl group that distinguishes RNA from DNA. Given these observations, it might be worthwhile to test whether, in this unusual and special case, the naive approach of simply replacing the RNA sequence with DNA would not significantly alter the structure of the nucleic acid or the drug-binding properties. Tests showed that the DNA construct could also bind aminoglycosides, although its affinity was significantly reduced (by a factor of 4 relative to the equivalent RNA construct). Several embodiments suggest that the same sequence synthesized from OMe RNA may also bind to aminoglycoside antibiotics, since methoxyRNA is likely closer to RNA than DNA. In several embodiments, sensors employing equivalent OMe RNA sequences not only bind to the antibiotic tobramycin but also exhibit higher affinity than sensors employing equivalent DNA aptamers. This increased affinity significantly improves the in vivo detection limit of the OMe RNA aptamer. Specifically, when placed in vivo, the root mean square noise observed before drug stimulation corresponds to detection limits (coefficient of variation of 3) of 1.2 µM and 9.9 µM for sensors employing OMe RNA or DNA aptamers, respectively.

[0104] Figures 3A to 3D This study demonstrates that, according to one embodiment, the use of an aminoglycoside-bound OMe RNA sequence significantly reduces drift of an EAB sensor deployed in situ in the jugular vein of a live rat. Figure 3A This paper illustrates an aminoglycoside-based EAB detection sensor using an OMe RNA or DNA sequence that binds to an aminoglycoside antibiotic, and characterizes its in vitro binding to tobramycin in whole blood at 37°C. Error bars represent the standard deviation (n=4) of independently fabricated sensors. Figure 3B In this study, a DNA-based sensor and an OMeRNA-based sensor were placed in the left and right jugular veins of rats, respectively. Three hours later, the animals were injected with tobramycin at a dose of 30 mg / kg via a femoral artery catheter. Under these conditions, the signal drift generated by the OMeRNA-based sensor was relatively small (peak currents at 100 Hz and 200 Hz, normalized to the first voltammogram, are shown). Specifically, during the 5-hour in vivo experiment, the drift of the OMeRNA-based sensor was only 7%, nearly 7 times lower than the 48% signal loss of the DNA-based sensor. Figure 3C The results show that while KDM can be used to correct drift in both sensors, signal loss associated with the DNA construct ultimately reduces the signal-to-noise ratio, thus decreasing accuracy. Note that for sensors using OMeRNA, the observed peak KDM signal (at drug excitation) is due to the high tobramycin binding affinity of that sequence. Figure 3A ). Figure 3D The conversion of these KDM signals into reproducible estimates of plasma drug levels is shown. The slightly lower Cmax of the OMe RNA sensor was associated with a slower rise in Cmax, suggesting that limitations such as thrombosis or sensor invasion of the vein wall may have slowed drug transport to the sensor. The duration of these experiments was limited by animal welfare concerns, which restricted the time we could maintain the animals under anesthesia. Figure 3E The signals of the DNA and Ome RNA aptamer sensors according to an embodiment are shown. Figure 3E A magnified view of signal changes before and after intravenous injection of 30 mg / kg tobramycin into the femoral vein of rats is shown. The injection was delivered separately to DNA and OMeRNA aptamer sensors in the left and right necks of individual rats. The slight delay (approximately 2.5 minutes) before the OMeRNA sensor reaches Cmax suggests that drug delivery to the sensor may be restricted due to thrombosis or invasion of the working electrode into the vein wall. Either effect would slow drug delivery to the sensor, resulting in a slightly lower observed Cmax.

[0105] To determine whether the in vivo drift performance of the OMeRNA-based sensor was improved compared to the DNA-based sensor, each sensor was placed in the left and right jugular veins of anesthetized rats, respectively. Under these conditions, the signal of the DNA-based device decreased by 48% after 5 hours, while the signal of the equivalent OMeRNA device decreased by only 7%. Figure 3B This difference exists across the square wave frequency range. Although KDM drift correction accurately corrects the drift of both ( Figure 3C However, just a few hours later, the greater signal loss of the DNA-based sensor significantly reduced its signal-to-noise ratio. Crucially, both sensors functioned correctly in animals, as demonstrated by intravenous excitation with tobramycin (30 mg / kg). Due to the higher affinity of the OMe RNA aptamer (…),… Figure 3A The resulting KDM signal change is significantly larger for the OMe RNA sensor. Figure 3C However, when these signals were converted to concentrations, the two datasets produced very similar concentration-time curves. Figure 3D ). Figure 3F and 3G The charge transfer kinetics of the DNA aptamer sensor and the OMeRNA aptamer sensor according to the embodiments are shown respectively. When aminoglycosides are used in vivo to bind DNA ( Figure 3F ) or OMe RNA ( Figure 3G When the sensors of the construct are excited, the drift of both is largely independent of the square wave frequency employed. The charge transfer kinetics shown here were recorded prior to tobramycin administration. Figure 3H Normalized signals for the DNA aptamer sensor and the OMe RNA aptamer sensor according to embodiments are shown. Inconsistent in vivo drift was observed between aminoglycoside-binding constructs at 70 Hz, a frequency that best matches the electron transfer rate.

[0106] The use of OMeRNA significantly improved the in vivo drift characteristics of aminoglycoside-based EAB sensors, a stark contrast to previous observations that this framework only slightly reduced drift when used in vitro in undiluted whole blood at 37°C (see, e.g., KKLeung et al., ACS Sensors 2021, 63340-3347; the disclosure of which is incorporated herein by reference). To elucidate the origin of this difference, several examples characterized the in vivo performance of devices employing unstructured constructs of DNA and OMeRNA to eliminate any complexity that might arise due to differences in secondary or tertiary structures. Previously, devices using these simple constructs exhibited significant initial drift when placed in whole blood at 37°C in vitro, although the OMeRNA-based devices showed less drift. Following this rapid decline, the signals produced by both devices subsequently “plateaued,” suggesting that this was due to some saturation effect, such as dirt. Building on this, several examples replicated these experiments in vivo in the left and right necks of four live rats. Under these conditions, the devices employing DNA and OMeRNA again exhibited a rapid initial loss of signal. Similar to in vitro experiments, the signal generated by the OMeRNA-based device subsequently plateaued and remained relatively stable. However, during our experiments, the signal from the DNA-based device continued to decline, a behavior that contrasted sharply with the behavior of the same device in blood in vitro. These observations suggest that while both in vivo and in vivo drift of the OMeRNA device are dominated by an eventual saturation mechanism (such as dirt), in vivo drift of the DNA-based device is primarily due to some unsaturated process (such as enzymatic degradation).

[0107] Figures 4A to 4C The behavior of the EAB sensor in in vitro and in vivo tests according to embodiments is illustrated. Figure 4A In the study, when excited in vitro in whole blood at 37°C, sensors made using DNA and OMe RNA both exhibited rapid signal loss, after which their signals stabilized. The shaded areas represent the standard deviations (n=4) of independently fabricated sensors. Figure 4B and Figure 4C In this study, when the OMe RNA analog was placed in the jugular vein of rats, the OMe RNA analog behaved similarly to OMe RNA. In contrast, the DNA analog continued to lose signal during these in vivo experiments. The surgical procedure of inserting each probe into the jugular vein lumen takes approximately 3 to 4 minutes, resulting in a delay between the sensor's exposure to the in vivo environment and the start of measurement. Therefore, this is illustrated here. Figure 4B An example of an OMe RNA sensor inserted before a DNA sensor, and Figure 4C An example of reversing the order is shown in the image. Several embodiments provide characterization of the electron transfer kinetics of the EAB sensor under various in vitro and in vivo conditions. These characterizations support the view that DNase-driven degradation dominates the in vivo drift of DNA-embedded devices. To see this, charge transfer for each device can be measured as a function of square wave frequency, first exploring DNA-embedded devices excited with DNase in buffer. Figures 5A to 5F The temporal evolution of electron transfer dynamics of EAB sensors placed under various conditions according to embodiments is illustrated. These measurements provide insights into the origin of signal loss when they are deployed in the blood both in vitro and in vivo. Figures 5A to 5F A Lovech plot of charge transfer versus frequency is shown over a 5-hour measurement duration. Arrows indicate the magnitude of signal loss over time and whether it corresponds to a decrease in the electron transfer rate indicated by a shift in the peak of the charge transfer distribution. When excited in vitro with DNA in phosphate-buffered saline at 37°C, the signal from unstructured (… Figure 5A DNA or ( Figure 5B The sensor signal of the OMe RNA construct decreased at all frequencies, indicating that while the number of methylene blue molecules on the surface decreased, the transfer kinetics of the reporter remaining on the surface remained unchanged otherwise. Notably, the decrease in the OMe RNA construct was relatively small, reflecting its relative nuclease resistance. In contrast, Figure 5C and 5D The results showed that when immersed in whole blood in vitro at 37°C, both devices exhibited a rapid initial decrease in charge transfer, accompanied by a shift to lower frequencies. Subsequently, the electron transfer characteristics of both devices tended to stabilize. Figure 5E and 5F The results showed that the two devices behaved fundamentally differently when excited in situ in the rat jugular vein. Specifically, the DNA-based device exhibited a sharp decrease in charge transfer at all frequencies, an effect very similar to that observed during in vitro enzymatic degradation. Figure 5A This contrasts sharply with the charge transfer behavior of the OMe RNA sensor, which mimics the charge transfer behavior of the OMe ribonucleic acid sensor when excited in whole blood in vitro. Figure 5D This indicates that dirt dominated the sensor's in vivo drift.

[0108] Under these conditions, the magnitude of charge transfer decreased sharply and continuously throughout the experiment, while the transfer rate remained constant. Figure 5AThis observation is consistent with methylene blue loss due to DNA cleavage, where the amount of methylene blue redox reporter is reduced, while the charge-retention transfer kinetics remain unchanged. Under these same conditions, the charge transfer reduction is much smaller with devices using OMe RNA. Similar to the case with DNA, the same decrease is observed at all frequencies. Figure 5B This indicates that this is also due to the more limited nuclease-driven degradation of this oligonucleotide. When both devices were excited in vitro in whole blood at 37°C, it had previously been shown that under dirt-dominated drift conditions, their charge transfer initially decreased and shifted to lower frequencies before eventually stabilizing. Figure 5C and 5D This may occur when proteins adsorb onto a surface, inhibiting the entry of redox reporters and slowing electron transfer. In contrast, the two devices exhibit more significant behavioral differences when excited in the bloodstream in vivo. Figure 5E and 5F Specifically, the behavior of the DNA-based device in vivo was similar to its behavior in vitro when excited in a buffer containing DNase: the signal drifted downwards throughout our experiment, and the transfer rate did not change significantly (comparison). Figure 5A and 5E In contrast, the device using OMe RNA behaved in vivo similarly to its behavior during in vitro excitation in whole blood: the magnitude of charge transfer initially decreased rapidly, along with the transfer rate, before both effects plateaued (comparison). Figure 5D and 5F These observations suggest that the in vivo behavior of devices employing DNA and OMeRNA differs at an important mechanistic level, and indicate that while in vivo drift of OMeRNA devices is primarily caused by fouling (which eventually saturates), drift of DNA-based devices is caused by DNase-driven oligonucleotide loss.

[0109] In many embodiments, the use of DNase-resistant OMe RNA aptamers significantly reduced the drift observed when EAB sensors using aminoglycosides were administered intravenously. In several embodiments, after 5 hours in the jugular vein of live rats, the in vivo drift of aminoglycoside sensors using OMe RNA aptamers was 7-fold less than that of sensors using equivalent DNA aptamers. Subsequent mechanistic studies based on several embodiments using simple model oligonucleotides (i.e., lacking secondary or tertiary structures) indicated that this was because enzymatic degradation was the dominant factor in the drift of in vivo deployed EAB sensors, a finding that contrasts sharply with previous studies that showed dirt dominated the drift observed in vitro in blood at 37°C. Several embodiments provide the application of XNA aptamers in the in vivo duration of EAB sensors and other aptamer-based technologies.

[0110] Example 6: Sensor devices based on wearable microneedles The working electrodes (and any other electrodes) described herein can be configured as microneedles and incorporated into wearable sensor devices, exemplary types such as Figure 6 , Figure 7A , Figure 7B , Figure 8 , Figure 9 and Figure 10 As shown. The device includes an upper housing portion (25) and a skin contact portion (30). A removable flexible layer (90) that can be gripped by tabs (95) is also provided, the removal of which exposes a dermatologically acceptable adhesive on the skin contact surface (35). The adhesive is intended to hold the device in place on the subject's skin for an extended period. The flexible layer (90) functions to prevent the adhesive from curing or drying, to prevent contamination of the adhesive layer before use, and / or premature adhesion of the adhesive to packaging or other surfaces. In a particularly preferred embodiment, in addition to covering the adhesive layer, the flexible layer (90) extends in space (45) to prevent contamination of the microneedles (15) and also helps prevent accidental needlestick injuries to the user. The device may have a retaining portion for holding the device on the skin, such that the protruding portion remains in contact with the subject's bodily fluids. The retaining portion may be dedicated to this function or may perform other functions. In many embodiments, having or including a retaining portion made of a dermatologically acceptable adhesive would be useful. The adhesive allows the user to easily apply the device, typically requiring only the removal of the protective backing to expose the adhesive, which is then applied to the skin. This application method is similar to the application of a plaster and is therefore a familiar process to the user.

[0111] As an alternative to using adhesives, the retaining component can be a mechanical device for holding the device in the desired position on the skin. For example, the device could include an adjustable, specialized strap around the limb attachment to keep the device securely applied to the subject. Alternatively, the device could be integrated into wearable items such as gloves or shirts, or into jewelry such as rings to hold the device in place. The device can be configured to engage with a standalone wearable item (e.g., via complementary hook-and-loop mechanisms) or the wearable item can be integrally molded to it. In some embodiments, the device is held in place solely by a wearable item abutting against the housing. For example, the holding element could be a close-fitting elastic glove worn on the device.

[0112] In some embodiments, the retaining portion is any surface or part of the device that comes into contact with the subject's skin, and the subject's characteristics are at least partially responsible for holding the device in the appropriate position on the subject's body. For example, the device may be configured to be held between two parts of the body, typically in a close fit, or to be embedded within existing anatomical structures. The shape and / or size of the device may allow it to be held between the toes, buttocks, groin, mouth, nostrils, ear canals, or navel. In some embodiments, the shape and / or size of the device housing are designed to fit snugly onto, for example, fingers, toes, or ears. The device housing may be elastically deformable, for example, made of rubber material, and configured to stretch over any anatomical location, such as a finger.

[0113] In the context of this invention, each embodiment can be a holding part. The device also includes a release member (100) having a gripping portion (105) and a wedging portion (110), the function of which will be described more fully below. Now go to Figure 7A and Figure 7B The exploded view shows that the components shown are similar to those in the previous diagrams, which is immediately apparent. In some embodiments, the power responsible for moving the arm (205) to push the microneedles (15) into the underlying skin is provided by the user. In use, the user places their fingers on the upper housing (25) and pushes downwards. Furthermore, the arm (205) can be moved via a hinge mechanism.

[0114] The hinge mechanism is provided by opposing lugs (115) extending from the skin contact portion (30), each lug including a hole. The arm (205) includes opposing laterally extending discs (120), each disc located in the hole of the lug (115). It is evident that the arm (205) is hinged relative to the skin contact portion (30) to allow movement from a first position to a second position.

[0115] The arm (205) is presented to the user with the arm in a first position. The arm (205) is held in the first position by a wedge portion (110) of the release member (100). The wedge portion is inserted between the skin contact portion (30) and the arm (205) before the release member (100) is removed, thereby holding the microneedle within the device.

[0116] When the device is to be applied to the subject's skin, the user removes the flexible layer (90) by pulling the tab (95) to expose the adhesive layer on the skin contact surface (35). The device is then applied to the skin, where the adhesive keeps it in place for an extended period.

[0117] Once the device is applied to the skin, the user grasps the gripping part (105) and pulls it laterally to the left (as shown) to completely remove the release member (100). The release member (100) has no other function and is discarded at this point. By removing the release member (100), the arm (205) is released from the first position and allowed to move (under a downward force applied by the user) to the second position, whereby the lower surface (205) of the arm contacts the upper surface of the skin contact part (30). In the second position, the microneedles (15) extend through the space (45) and into the underlying skin.

[0118] As will be understood, the release member (100) may be configured to prevent the upper housing (25) of the device from closing onto the skin contact portion (30) without the user's intentional action. The release member (100) is inserted or otherwise juxtaposed between the upper housing (25) and the skin contact portion (30) to prevent closure of the upper housing (24) toward the skin contact portion sufficiently to allow the tip of the microneedle (i.e., the protrusion) to protrude from the base of the hole in the skin contact portion. Preventing closure also prevents movement of the arm (205) from the first position to the second position. Thus, when the release member (100) is in place, the tip of the microneedle will not be accidentally touched, resulting in microneedle contamination or injury. When using the device, the user removes the release member (100) as a step in the process of use. In a preferred embodiment of device use, the user first adheres the device to the subject's skin before pressing the upper housing (25) to insert the microneedle into the skin, and then removes the release member (100). Before being removed by the user, the release member (100) can be held in place by any of a variety of features. In one example, the release member (100) includes a protrusion that fits into a recess in the upper housing (25), the skin contact portion (30), or both the upper housing (25) and the skin contact portion (30) to help hold it in place until it is intentionally removed. In another example, the release member (100) is designed to slidably assemble onto the skin contact portion (30) or the upper housing (25) such that friction between the release member (100) and the upper housing (25) or the skin contact portion (30) helps hold it in place until it is intentionally removed. In yet another example, magnetic force can be used to help hold the release member (100) in place. In one embodiment of the invention, when the release member (100) is in place, a magnet mounted within the release member (100) is positioned close to a Hall effect sensor located in the upper housing (25) or the skin contact portion (30). According to this embodiment, when the user removes the release member (100), the Hall effect sensor detects the removal of the magnet and causes the device to take some action, such as energizing the electronic circuitry ready for use, switching it from sleep mode to active mode. It should be understood that the above are examples of possible methods to help hold the release member (100) in place before intentional removal; these methods can be used alone or in combination, and other methods known in the art can also be used alone or in combination with the given examples.

[0119] In some embodiments of the invention, the release member (100) may also serve as a covering element for covering the microneedles after the device has been removed from the subject. In a preferred example of this embodiment, a locking element is located on the upper housing (25) and extends downward toward the skin contact portion (30). The release member (100) includes a groove that allows the release member (200) to slide past the locking element as it is withdrawn from the device, while maintaining continuity of the surface of the release member (300) facing the upper surface of the skin contact portion (30). In use, the release member (100) according to this preferred embodiment is removed by the user before the upper housing (25) is pressed to insert the microneedles into the subject's skin and held by the user. After the device has been removed from the subject following use, the user is instructed to adhere the release member (100) to an adhesive layer on the lower surface of the skin contact portion (30) to cover the protruding microneedles. In another example of this embodiment, the release member (100) is flexibly attached to the device such that the release member (100) remains attached to the device after the user withdraws it, and is then repositioned to cover the protruding microneedles after the device is removed from the subject following use. In another example of this embodiment, the release member (100) and the upper housing (25) are designed such that the release member (100) can slidably or otherwise engage with the upper housing (15) once the release member (10) is removed, wherein the release member is stored during device use and removed as a covering element after the device is removed from the subject.

[0120] In some embodiments, the device is configured to facilitate removal from the subject by a user. As will be understood, the use of an adhesive layer may result in difficulty in removing the device from the skin. Examples of such configurations include leaving a portion of the skin-contact surface (35) uncoated with adhesive, creating a gap between the subject's skin and the surface (35), whereby the user uses this gap as a lever point to help pull the device off the skin by breaking the adhesive. In another example, a lever mechanism not located on the skin-contact surface is incorporated to allow a gap greater than that created by leaving a portion of the skin-contact surface uncoated. In yet another example, a tab extending beyond at least one edge of the skin-contact portion (30) and attached to the adhesive layer may be incorporated, whereby the user pulls the tab with sufficient force to stretch and yield the adhesive layer, further delaminating the adhesive from the skin-contact surface (35) and the skin.

[0121] In some embodiments, the device is designed such that the release member (100) is locked in its position before the device is used unless pressure is applied to the upper housing (25). This embodiment aims to further reduce the risk of premature retraction of the release member (100). In this example embodiment, when the upper housing (25) is not pressed, the release member (100) and at least one of the upper housing (25) and the skin contact portion (30) have lockably engaging features. When the upper housing (25) is pressed down, features on at least one of the upper housing (24) and the skin contact portion (30) are deformed to disengage the release member (100) and allow it to be withdrawn.

[0122] In other embodiments, the release member (100) does not need to be removed from the device by the user. According to these embodiments, the release member (100) comprises a flexible element with sufficient stiffness such that it does not substantially deflect when subjected to closing forces that may exist on the device and in the user's hand during manufacturing, storage, and before application to a subject, but is flexible enough that it deflects when the user intentionally applies a closing force to the device while applying it to the subject's skin. In this bending, the release member (100) is deflected, allowing the upper housing (25) to close toward the skin contact portion (30). In these embodiments, the release member (100) may also be used as a locking element, or the release member (100) may be separate from the locking portion. In some of these embodiments, such as Figure 7A , Figure 7B The feature marked (220) forms the release component (100).

[0123] Each space (45) of the device is sized such that a microneedle can pass clearly through it without at least one tapered portion of the microneedle impacting the side of the aperture during insertion. In some embodiments, the aperture may have a sufficient cross-section such that no part of the microneedle contacts the side of the space during insertion. In several embodiments, at least a portion of the aperture along its length will have a cross-section such that a portion of the microneedle's length contacts the side of the aperture during insertion. According to this embodiment, the aperture serves to help support a portion of the microneedle's length to help prevent the microneedle from bending during insertion. In some embodiments, the skin contact portion (30) includes a further space or recess configured to receive a protrusion on the release member to help retain the release member until it is removed by the user. Alternatively, the skin contact portion (30) includes a protrusion designed to be received into a recess in the release member to help hold the release member in place until it is intentionally removed by the user.

[0124] The device includes a locking portion in the form of a latch (220) that permanently locks the arm (205) in a second position, preventing any articulated movement. In the illustrated embodiment, the latch (220) is a simple, integral component capable of deflecting in response to movement of the arm (205) toward a closed position, but returning to its original position when the arm (205) is in the second position (205b), thereby locking the arm (205) in place. The locking component can act on another part of the device, rather than on the arm (205), which in turn locks the arm in place. For example, the locking component can act on the upper housing (25), which in turn holds the arm (205) in a second position. In another alternative embodiment, the locking component can act on the PCB (65), which in turn holds the arm (205) in a second position.

[0125] In some embodiments, the locking portion includes a recess into which a protrusion on the upper housing (25) inserts to lock the upper housing (24) in a closed position (i.e., the arm (205) is in a second position). In one embodiment, the locking portion includes a flexible element designed to allow the locking portion to move upon impact with the upper housing (25), thereby closing the housing (25) relative to the skin-contact portion (30), such that once the upper housing (25) is closed, the locking portion can move to lock the upper housing (25) in the closed position. In one embodiment, the device includes a protrusion on the upper housing (25) designed to insert into the recess of the locking portion, the protrusion including a flexible element to allow the protrusion to move to close the upper housing (25) relative to the skin-contact portion (30), after which the housing (25) is closed relative to the skin-contact portion (30), at which point the protrusion moves and inserts into the recess of the locking portion, thereby locking the upper housing (25) in the closed position. The flexible element may include a shaft that is sufficiently deformable to allow the upper housing (25) to close without yielding, so that the flexible element will attempt to return to its original position after the upper housing (25) is closed. In a less preferred but still functional embodiment, the flexible element includes a helical spring.

[0126] The flexible element of the locking portion can be made of any suitable material with the necessary stiffness and yield point. Examples of suitable materials include amorphous plastics, crystalline plastics, spring steel, non-spring steel, stainless steel, or other materials known in the art with suitable mechanical properties. In several embodiments, the locking portion is made of the same material as the skin contact portion (30) to facilitate the fabrication of a skin contact portion having an integral locking portion. In some embodiments, the force required to deflect or otherwise move the flexible element is designed to be large enough that the pressure required from the user is sufficient to deform the flexible element, thereby closing the upper housing (25) toward the skin contact portion, sufficient to insert the microneedle into the skin. According to this embodiment, the flexible element of the locking portion is used to set the force required to close the device (thereby placing the arm in a second position) and ensure that the force is sufficient to insert the microneedle into its intended location embedded in the skin. In some embodiments, the locking portion includes at least one adhesive area located on at least one of the lower surface of the upper housing (25) and the upper surface of the skin contact surface (35). When the device is closed, one or more adhesive areas adhere the upper housing (25) to the skin contact portion (30), locking the device in the closed position.

[0127] In several embodiments, the locking portion can present three different stable states. In a first state, the locking portion is in a disengaged configuration before the upper housing (25) is pushed down toward the skin contact portion (30) to close the device. In a second state, the locking portion is in a first engaged position. When the locking portion is in the first engaged position, it is used to lock the microneedle (15) in the embedded position in the skin (i.e., the arm (205) is in the second position). In a third state, the locking portion is in a second engaged position. In this state, the locking portion locks the device in the open position (i.e., the arm (205) is in the first position) while retracting the microneedle into the device to reduce the possibility of microneedle protrusion leading to needlestick injury after device use. In several embodiments, the locking portion includes a user engagement portion that can be grasped by a user or otherwise engaged, for example by engaging a fingernail under a hanging flange so that the user can deflect a flexible portion of the locking portion. According to this embodiment, in order to close the device, the user presses the upper housing (25) and locks it in place (as in other embodiments disclosed herein). When it is necessary to remove the device from the subject, the user engages the locking portion and deflects it in a first direction to unlock the upper housing (25) from the skin contact portion (25), and then deflects the locking portion in a second direction to lock the device in the open position (i.e., the arm is in the first position) while the microneedles are in the retracted position. In a preferred embodiment of this example, the locking portion moves away from the device body in the first direction and toward the device body in the second direction. When fully deflected in the second direction, the locking portion is designed to engage, for example, stably in a recess to prevent the device from being accidentally closed.

[0128] In some embodiments, when inserted into the skin, the downward force on the microneedle is provided by a flexible element of the locking portion, which applies the downward force when the device is locked in the closed position (i.e., the movable arm is in the second position). In some embodiments, effective locking of the movable arm in the second position is provided by a dedicated spring or other suitable biasing device. In some embodiments, the plug or other biasing device is not dedicated to the locking function; for example, it may also serve as the driving force for moving the arm from the first position to the second position. For example, a torsion spring may apply a closing torque at the pivot point (if present). In another example, a flat, disc, or helical spring is mounted at the rear of the microneedle so that when the device is closed, the spring deforms or compresses to apply a downward force on the microneedle when the device is in the closed position.

[0129] While the PCB (65) is not an essential feature of the present invention, it is necessary in many applications where microneedles are used to conduct current into, from, or through the skin. In this regard, the PCB may carry a microprocessor and / or volatile electronic memory (such as RAM) and / or non-volatile electronic memory (such as ROM) and / or a wireless network module (such as a Bluetooth™ module). The device will, of course, include a power source (typically via a coin cell battery). Those skilled in the art will understand that the invention described herein is readily adaptable to further variations and modifications beyond the specific description. It should be understood that this invention includes all such variations and modifications falling within the spirit and scope of the invention.

[0130] Therefore, the spirit and scope of the invention are not limited to the examples described above, but should be understood in the broadest sense permitted by law.

[0131] Example Example 1: A working electrode for an electrochemical sensor, the working electrode including a conductive element and an associated analyte sensing element configured to specifically interact with a target analyte, wherein the sensing element includes a non-natural nucleic acid coupled to a redox reporter.

[0132] Example 2: The working electrode of Example 1, wherein non-natural nucleic acids cannot be read and / or replicated by any natural mammalian cell.

[0133] Example 3: The working electrode of Example 1 or 2, wherein non-natural nucleic acids are more resistant to degradation of biological fluid components compared to similar natural nucleic acids.

[0134] Example 4: The working electrode of Example 1, 2 or 3, wherein the similarity relates to one or more of the following: length, base sequence, secondary structure, tertiary structure and ability to interact with the target analyte.

[0135] Example 5: The working electrode of any one of Examples 1 to 4, wherein the biological fluid contains a nuclease.

[0136] Example 6: The working electrode of any one of Examples 1 to 5, wherein the non-natural nucleic acid is formed from a single strand.

[0137] Example 7: The working electrode of any one of Examples 1 to 6, wherein the non-natural nucleic acid is a polymer.

[0138] Example 8: The working electrode of any one of Examples 1 to 7, wherein the non-natural nucleic acid comprises 10 to 100 subunits. Example 9: The working electrode of any one of Examples 1 to 8, wherein the non-natural nucleic acid is natural deoxyribonucleic acid or a chemical variant of natural ribonucleic acid.

[0139] Example 10: The working electrode of any of Examples 1 to 9, wherein the chemical variant is a sugar backbone and / or one or more bases.

[0140] Example 11: A working electrode of any of Examples 1 to 10, wherein the non-natural nucleic acid is prepared by or with human assistance.

[0141] Example 12: The working electrode of any of Examples 1 to 11, wherein the non-natural nucleic acid is a heterologous nucleic acid (XNA) or a peptide nucleic acid (PNA).

[0142] Example 13: The working electrode of any of Examples 1 to 12, wherein the non-natural nucleic acid is a DNA or RNA aptamer with a modified chemical structure.

[0143] Example 14: The working electrode of any one of Examples 1 to 13, wherein the non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the interaction capability of the non-natural nucleic acid is at least 50%, 60%, 70%, 80%, 90%, 100%, 110%, 120%, 130%, 140%, or 150% of that of the natural nucleic acid. Example 15: The working electrode of any one of Examples 1 to 14, wherein the non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the non-natural nucleic acid has a recognition sensitivity, accuracy, or specificity for the target analyte that is at least 50%, 60%, 70%, 80%, 90%, or 100% of that of the natural nucleic acid. Example 16: The working electrode of any one of Examples 1 to 15, wherein the non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the uncorrected signal loss of the non-natural nucleic acid is less than 90%, 80%, 70%, 60%, or 50% of that of the natural nucleic acid. Example 17: The working electrode of any one of Examples 1 to 16, wherein the working electrode is part of an electrochemical sensor having an uncorrected signal drift rate, and for at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

[0144] Example 18: The working electrode of any of Examples 1 to 17, wherein a non-natural nucleic acid is bound to a conductive element at the first end.

[0145] Example 19: The working electrode of any of Examples 1 to 18, wherein a redox reporter is bound to the second end of a non-natural nucleic acid. Example 20: The working electrode of any of Examples 1 to 19, wherein the conductive element includes a skin-penetrating portion incorporating a non-natural nucleic acid.

[0146] Example 21: The working electrode of any of Examples 1 to 20, wherein the conductive element is a needle, microneedle, or wire.

[0147] Example 22: An electrochemical sensor device comprising a working electrode and a counter electrode of any one of Examples 1 to 21.

[0148] Example 23: The apparatus of Example 22 includes a reference electrode.

[0149] Example 24: The device of Example 22 or 23 has an associated retainer configured to keep the working electrode in contact with the subject's bodily fluids.

[0150] Example 25: The apparatus of Examples 22, 23 or 24, wherein the retainer is configured to keep the working electrode in contact with the subject's bodily fluids for at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours or 72 hours.

[0151] Example 26: A device of any one of Examples 22 to 25, having an associated housing configured to enclose the power supply and / or electronics used for sensor operation.

[0152] Example 27: A method for monitoring a target analyte in a subject's biological fluid, comprising: contacting a working electrode of any of Examples 1 to 21 with the biological fluid for a period of time.

[0153] Example 28: The method of Example 27, wherein the working electrode is in contact with a biological fluid that remains in situ within the subject for the duration of the method.

[0154] Example 29: The method of Example 27 or 28, where the biological fluid is blood or interstitial fluid.

[0155] Example 30: The method of Examples 27, 28 or 29, wherein the biological fluids have not yet been removed from the subject.

[0156] Example 31: The method of any one of Examples 27 to 30, wherein the time period is at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

Claims

1. A working electrode for an electrochemical sensor, the working electrode comprising a conductive element and an analyte sensing element associated with the conductive element, the analyte sensing element being configured to specifically interact with a target analyte, wherein, The sensing element includes a non-natural nucleic acid coupled with a redox reporter.

2. The working electrode according to claim 1, wherein, The non-natural nucleic acid cannot be read and / or replicated by any natural mammalian cell.

3. The working electrode according to claim 1 or 2, wherein, Compared to similar natural nucleic acids, the non-natural nucleic acids are more resistant to degradation by biological fluid components.

4. The working electrode according to claim 3, wherein, The similarity is related to any one of the following: length, base sequence, secondary structure, tertiary structure, and ability to interact with the target analyte.

5. The working electrode according to claim 3 or 4, wherein, The biological fluid contains nucleases.

6. The working electrode according to any one of claims 1 to 5, wherein, The non-natural nucleic acid is formed from a single strand.

7. The working electrode according to any one of claims 1 to 6, wherein, The non-natural nucleic acid is a polymer.

8. The working electrode according to claim 7, wherein, The non-natural nucleic acid contains 10 to 100 subunits.

9. The working electrode according to any one of claims 1 to 8, wherein, The non-natural nucleic acid is a natural deoxyribonucleic acid or a chemical variant of natural ribonucleic acid.

10. The working electrode according to claim 9, wherein, The chemical variants involve a sugar backbone and / or one or more bases.

11. The working electrode according to any one of claims 1 to 10, wherein, The non-natural nucleic acid was prepared by humans or with human assistance.

12. The working electrode according to any one of claims 1 to 11, wherein, The non-natural nucleic acid is a heterologous nucleic acid (XNA) or a peptide nucleic acid (PNA).

13. The working electrode according to any one of claims 1 to 12, wherein, The non-natural nucleic acid is a DNA or RNA aptamer with a modified chemical structure.

14. The working electrode according to any one of claims 1 to 13, wherein, The non-natural nucleic acid is a modified form of natural nucleic acid capable of specifically interacting with the target analyte, and the non-natural nucleic acid has at least 50%, 60%, 70%, 80%, 90%, 100%, 110%, 120%, 130%, 140%, or 150% of the interaction capability of the natural nucleic acid.

15. The working electrode according to any one of claims 1 to 14, wherein, The non-natural nucleic acid is a modified form of natural nucleic acid that can specifically interact with the target analyte, and the sensitivity, accuracy, or specificity of the non-natural nucleic acid for recognizing the target analyte is at least 50%, 60%, 70%, 80%, 90%, or 100% of that of the natural nucleic acid.

16. The working electrode according to any one of claims 1 to 15, wherein, The non-natural nucleic acid is a modified form of natural nucleic acid that can specifically interact with the target analyte, and the uncorrected signal loss of the non-natural nucleic acid is less than 90%, 80%, 70%, 60%, or 50% of that of the natural nucleic acid.

17. The working electrode according to claim 1, wherein, The working electrode is part of an electrochemical sensor with an uncorrected signal drift rate, and the uncorrected signal drift rate is averaged or determined over a period of at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

18. The working electrode according to any one of claims 1 to 17, wherein, The first end of the non-natural nucleic acid binds to the conductive element.

19. The working electrode according to any one of claims 1 to 18, wherein, The redox reporter binds to the second end of the non-natural nucleic acid.

20. The working electrode according to any one of claims 1 to 19, wherein, The conductive element includes a skin-penetrating portion having the non-natural nucleic acid bound thereto.

21. The working electrode according to claim 20, wherein, The conductive element is a needle, microneedle, or wire.

22. An electrochemical sensor device comprising a working electrode and a counter electrode as described in any one of claims 1 to 21.

23. The device of claim 22, wherein the device includes a reference electrode.

24. The device of claim 22 or 23, wherein the device has an associated retainer configured to maintain the working electrode in contact with bodily fluids of the subject.

25. The device according to claim 24, wherein, The retainer is configured to keep the working electrode in contact with the subject's bodily fluids for at least 10, 20, 30, 40, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

26. The device according to any one of claims 22 to 25, the device having an associated housing configured to enclose a power supply and / or electronics for operation of the sensor.

27. A method for monitoring a target analyte in the biological fluids of a subject, comprising: The working electrode according to any one of claims 1 to 21 is brought into contact with biological fluid for a period of time.

28. The method according to claim 27, wherein, During the duration of the method, the working electrode is in contact with biological fluids that remain in situ within the subject's body.

29. The method according to claim 27 or 28, wherein, The biological fluid is blood or tissue fluid.

30. The method according to any one of claims 27 to 29, wherein, The biological fluids have not yet been removed from the subject's body.

31. The method of any one of claims 27 to 30, wherein, The time period is at least 10 minutes, 20 minutes, 30 minutes, 40 minutes, 50 minutes, 1 hour, 2 hours, 3 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, 9 hours, 10 hours, 11 hours, 12 hours, 13 hours, 14 hours, 15 hours, 16 hours, 17 hours, 18 hours, 19 hours, 20 hours, 21 hours, 22 hours, 23 hours, 24 hours, 36 hours, 48 ​​hours, or 72 hours.

Images