Sensor, kit and method for detecting biomarkers in saliva

Abstract

The present invention provides a sensor for measuring biomarker or metabolite levels, such as glucose levels, in a fluid sample, such as saliva, comprising a sensing element comprising a redox reporter linked to at least one aptamer that binds to a saliva metabolite, and a solid surface; the sensing element is attached to a solid surface wherein the redox reporter undergoes a binding-induced electron transfer change when the biomarker or metabolite binds to the aptamer. Measurement of biomarker or metabolite levels in a patient's saliva allows for an increased risk of monitoring a patient's state of health, such as a patient's gingival disease or diabetes.

Description

[0001] Cross - Reference to Related Applications

[0002] This application claims priority to U.S. Provisional Application No. 63 / 375,802, filed on September 15, 2022, the content of which is incorporated herein by reference in its entirety. Technical Field

[0003] The present invention provides a sensor, a kit, and a method for measuring the level of a biomarker in saliva.

[0004] Background

[0005] Monitoring of various medical conditions and health states requires tests that are regularly or frequently performed by patients to monitor their condition or health.

[0006] Sensors based on electrochemical processes can be used to detect biological substances by using a transduction element to convert a detection event into a signal for processing and / or display. A biosensor can use a biological material as a biorecognition component, such as an enzyme, an antibody, or a nucleic acid. A biosensor typically uses a transduction element to convert an analyte signal detected from the biorecognition component into a different signal that can be processed by optical, electronic, or other means.

[0007] An example of a target analyte is measuring the glucose level in diabetic patients. Diabetes mellitus is a metabolic disease associated with hyperglycemia due to insufficient production of insulin in the human body or inadequate response of cells to the produced insulin. Good blood glucose control is also considered important for preventing the development of gingivitis. Diabetic patients have a higher risk of developing gum diseases.

[0008] Therefore, there is still a need to provide new means for measuring the level of biomarkers in body fluids.

[0009] Summary

[0010] The present invention provides a sensor for measuring a target molecule in a saliva sample, comprising a sensing element and a solid surface, the sensing element comprising a redox reporter connected to at least one aptamer that binds to a saliva metabolite; the sensing element is connected to the solid surface, wherein when the saliva metabolite binds to the aptamer, the redox reporter undergoes a binding - induced change in electron transfer.

[0011] Specifically, the present invention provides a sensor for measuring the level of a target molecule in a fluid sample, comprising a sensing element and a solid surface, the sensing element comprising a redox reporter connected to at least one aptamer that binds to glucose, the sensing element is connected to the solid surface, wherein when glucose binds to the aptamer, the redox reporter undergoes a binding - induced change in electron transfer.

[0012] In one embodiment, the aptamer is attached to the solid surface via a thiol molecule.

[0013] In another embodiment, the redox reporter is a methylene blue molecule.

[0014] In a further embodiment, at least one aptamer comprises a nucleotide sequence as shown in SEQ ID NO:3 or SEQ ID NO:4, or a combination thereof.

[0015] In a further embodiment, at least one aptamer comprises SEQ ID NO:7.

[0016] In an additional embodiment, the sensor covered herein further comprises at least one aptamer that binds to a different target molecule.

[0017] In one embodiment, the target molecule is glucose.

[0018] In another embodiment, the analyte is vancomycin.

[0019] In a further embodiment, the analyte is adenosine triphosphate (ATP), carnitine, lactate, lactic acid, malate, malic acid, maltose or adenosine monophosphate (AMP).

[0020] In one embodiment, the carnitine is 2-methylbutyrylcarnitine, butyrylcarnitine, isobutyrylcarnitine or propionylcarnitine.

[0021] In one embodiment, at least one aptamer that binds to a different target molecule comprises a nucleotide sequence as shown in SEQ ID NO:1.

[0022] In a further embodiment, the solid surface is an electrode.

[0023] In another embodiment, the electrode is a gold electrode.

[0024] In a complementary embodiment, the fluid sample is saliva.

[0025] In a further embodiment, the sensor element is used for a long time.

[0026] In one embodiment, the sensor element is used for more than three days.

[0027] In another embodiment, the sensor element is regenerated and used for multiple measurements.

[0028] The present invention also provides a kit, which comprises a sensor element as defined herein and instructions for use.

[0029] In one embodiment, the kit covered herein further comprises a monitoring device.

[0030] In another embodiment, the monitoring device monitors changes in electron transfer induced by the binding of a redox reporter.

[0031] In a further embodiment, electron transfer is monitored by changes in the peak height of a square wave voltammogram.

[0032] In another embodiment, the monitoring device monitors changes in binding-induced electron transfer by voltammetry or impedance measurement.

[0033] The present invention also provides a method for measuring at least one metabolite (such as glucose level) in a patient's fluid sample, which comprises contacting the fluid sample with a sensor as defined herein and monitoring changes in binding-induced electron transfer of a redox reporter to indicate the presence of at least one metabolite in the fluid sample.

[0034] In one embodiment, the method described herein further comprises the step of measuring the presence of at least one other analyte.

[0035] In one embodiment, measuring the level of at least one metabolite allows monitoring of the patient's health status.

[0036] In a further embodiment, measuring the glucose level indicates the presence or progression of diabetes in the patient or an increased risk of gum disease.

[0037] In one embodiment, the method covered is used to measure at least one metabolite using the sensor for a long time.

[0038] In one embodiment, the method covered is used to measure at least one metabolite using the sensor for more than three days.

[0039] In another embodiment, the method covered herein further comprises the steps of cleaning and storing the sensor and using the sensor again for multiple measurements.

[0040] In a further embodiment, the metabolite is measured before and / or after the patient eats.

[0041] Brief Description of the Drawings

[0042] Reference will now be made to the drawings.

[0043] Figure 1It is illustrated that (A) the electrochemical aptamer-based biosensor consists of an aptamer sequence modified with methylene blue that is linked to the surface of a monolayer-coated gold electrode by thiol chemistry; and (B) upon target recognition, the methylene blue redox reporter undergoes binding-induced electron transfer changes that can be easily monitored, for example, by changes in the peak height of square wave voltammograms.

[0044] Figure 2 It is illustrated that the signal response of the E-AB sensor depends on the square wave frequency, as shown for (A) the ATP aptamer and (B) the parental glucose aptamer observed over a wide square wave frequency range in working buffer and undiluted human saliva. Error bars represent the standard deviation observed for at least three independently fabricated and tested biosensors. The pH of the saliva samples was 7.25 in (A) and 7.74 in (B).

[0045] Figure 3 It is illustrated that when used (deploy) in undiluted saliva, the analytical performance (i.e., limit of detection and signal change) of the E-AB sensor is improved, in (A) for sensors that bind to aptamers that bind adenosine monophosphate; in (B) for sensors that bind glucose in saliva; and in (C) for the improved glucose aptamer signal response of three new variants.

[0046] Figure 4 It is illustrated that the E-AB sensors that bind to adenosine monophosphate (A) and glucose (B) respond rapidly within 20 seconds when attacked with 100 nM adenosine monophosphate and 5 pM glucose. Error bars represent the standard deviation of the mean measured for at least three independently fabricated sensors.

[0047] Figure 5 It is illustrated that the aptamer that binds glucose shows long-term stability when continuously interrogated in undiluted saliva for more than 3 days. The shading represents the standard deviation of the mean for eight continuously detected and independently fabricated sensors.

[0048] Figure 6 It is illustrated that the E-AB biosensor is able to withstand cycling of potentially more than 10 rinse / target attack steps continuously directly in undiluted saliva. The shading represents the standard deviation of the mean for at least 4 continuously detected and independently fabricated sensors.

[0049] Figure 7 It is illustrated that the E-AB sensor shows enhanced performance when used in undiluted saliva compared to when in buffer, due to the unique combination of ionic species at a given concentration (upper figure), as by exposure to increasing amounts of the divalent cations Ca 2+ or Mg 2+as shown by the E-AB sensor that binds adenosine monophosphate, and CD spectroscopic studies confirmed that the mechanism of enhanced sensor signal in saliva is due to divalent ions such as Ca 2+ and Mg 2+ in the presence (see figure below).

[0050] Figure 8 Shows that the curve response of the E-AB sensor is square wave frequency-dependent, where the shading represents the standard deviation observed for at least three independently fabricated and tested biosensors.

[0051] Figure 9 Shows that the E-AB sensor exhibits improved analytical performance when used in undiluted saliva. Error bars represent the standard deviation of the mean for at least three independently fabricated samples (too small to be shown in some cases).

[0052] Figure 10 A Shows the change in glucose concentration over time in saliva samples collected from the same individual at different times during the same day using either a standardized test colorimetric measurement or by using the E-AB sensor. As determined by paired sample T-tests, the glucose sensor responses at 4 and 6 hours post-meal were significantly different from those at 30 minutes after lunch (*p = 0.017 and ***p = 5.42e-4, n = 4 sensors, where *p < 0.05, **p < 0.01, ***p < 0.001, ns indicates not significant). The shading represents the standard deviation observed for at least three independently fabricated and tested biosensors.

[0053] Figure 10 B Shows the change in adenosine monophosphate concentration over time in saliva samples collected from the same individual at different times during the same day using the E-AB sensor. The shading represents the standard deviation observed for at least three independently fabricated and tested biosensors.

[0054] DETAILED DESCRIPTION

[0055] According to the present specification, there is provided a sensor for measuring a biomarker (such as but not limited to glucose level) in a fluid sample (such as but not limited to saliva), comprising a sensing element and a solid surface, the sensing element comprising a redox reporter, the redox reporter being connected to at least one aptamer that binds to glucose; the sensing element being connected to the solid surface, wherein when glucose binds to the aptamer, the redox reporter undergoes a binding-induced change in electron transfer.

[0056] Electrochemical aptamer-based biosensors (E-ABs) have the potential to provide point-of-need and convenient measurements of biomarkers directly in undiluted complex matrices. The E-AB sensors covered in this article consist of redox reporters (such as but not limited to methylene blue) conjugated to aptamers. More specifically, an E-AB containing methylene blue and a thiol-modified aptamer sequence was demonstrated. The aptamers covered in this article can detect molecules or biomarkers such as glucose, carnitine (such as 2-methylbutyrylcarnitine, butyrylcarnitine, isobutyrylcarnitine, propionylcarnitine), lactate / lactic acid, malate / malic acid, adenosine 5'-monophosphate (AMP), or maltose. Aptamers are short single-stranded DNA (ssDNA) or RNA sequences that have been used as recognition receptors to replace antibodies in biosensing devices. Aptamers can be selected in vitro against various targets (including small molecules, metal ions, and proteins) through a process called SELEX.

[0057] In addition, the E-AB also includes a monolayer solid surface for aptamer attachment, such as a coated gold electrode surface (see Figure 1 ). Electrodes made of other known materials are covered, such as copper, graphite, titanium, brass, silver, as well as platinum, carbon, and silicon.

[0058] As provided in this article, the aptamer is attached to the solid surface through a thiol molecule. A large library of thiol-modified molecules can be used to passivate the gold electrode. The gold electrode can be subjected to different electrode treatments to change its surface area and porosity.

[0059] When the target binds, the conjugated redox reporter undergoes a binding-induced change in electron transfer, thus allowing direct quantification of the target concentration ( Figure 1 B).

[0060] As covered in this article, a monitoring device can be coupled to the E-AB to monitor the binding-induced change in electron transfer of the redox reporter. Various electrochemical techniques (i.e., voltammetry or impedance spectroscopy) can be used to monitor this change. As covered in this article, the electron transfer is monitored through the change in the peak height of the square wave voltammogram.

[0061] Because E-AB sensors rely on an electrochemical signal transduction mechanism, they have shown use in undiluted whole blood for antibiotics, chemotherapeutic drugs, abused drugs, and proteins. Although blood is considered the gold standard in most health checks, in contrast, saliva, as a clinically informative biofluid, represents an attractive and non-invasive diagnostic means. To date, except for SARS-CoV-2, E-AB sensors have not demonstrated usability in undiluted saliva.

[0062] Because E-AB sensors rely on an electrochemical signal transduction mechanism, they can be powered by a micro portable electronic device.

[0063] Previously reported aptamers that bind glucose and adenosine triphosphate (and also adenosine monophosphate) were tested in buffer to evaluate their ability to be used easily in undiluted, unstimulated saliva. ATP- and AMP-binding aptamers are well known in the art and were further adapted for the E-AB sensor. However, the applicability of such E-AB sensors in undiluted saliva has not been demonstrated. The "parent" glucose aptamer previously disclosed and incorporated in a field effect transistor platform did not demonstrate measurements in undiluted complex matrices because it relied on a signal transduction scheme that was prone to electrode contamination. In this way, the analytical merits (sensitivity, dynamic range, limit of quantification, limit of detection) of the biosensor in such complex matrices were determined.

[0064] The aptamer sequences used were purchased from Bio Basic Inc. and purified by HPLC. These sequences were not further purified, simply resuspended in deionized water to 100 μM, and then aliquoted into 2 μL tubes. Bio Basic Inc. chemically modified the 5′ end of the aptamer with thiol chemistry on a 6-carbon linker and the 3′ end with carboxyl-modified methylene blue that was attached to the DNA by forming an amide bond with a primary amine on the 6-carbon linker.

[0065] -ATP aptamer: 5’-HS-SH-(CH 2 ) 6 -ACC TGG GGG AGT ATT GCG GAG GAA GGT-O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:1)

[0066] -Glucose parent aptamer (previously disclosed and modified with methylene blue and thiol): 5’-HS-SH-(CH 2 ) 6 - ACG ACC GTG TGT GTT GCT CTG TAA CAG TGT CCA T TG TCG T -O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:2)

[0067] -Glucose 3-trunc aptamer (new, unreported aptamer): 5’-HS-SH-(CH 2 ) 6 - ACC GTG TGTGTT GCT CTG TAA CAG TGT CCA T TG T -O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:3)

[0068] -Glucose 6-trunc aptamer (new, unreported aptamer): 5’-HS-SH-(CH 2 ) 6 -GTG TGT GTTGCT CTG TAA CAG TGT CCA T–O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:4)

[0069] -T26 sequence (control sequence of ATP aptamer): 5’-HS-SH-(CH 2 ) 6 -ATT ATT TTT TAT TTA TTTTTA TTT TAT-O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:5)

[0070] -T40 sequence (control sequence of original glucose aptamer): 5’-HS-SH-(CH 2 ) 6 -ATT ATT TTT TATTTA TTT TTA TTT TAT TTT ATT TTT TAT T-O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:6)

[0071] Before fabricating the sensors, the electrodes (2.0 mm in diameter, CH Instruments) were mechanically cleaned by sequentially polishing with 1 μm diamond suspension oil slurry and then an aqueous solution of 0.05 μm alumina powder. After each polishing step, the electrodes were sonicated in ethanol or distilled water for 5 minutes respectively.

[0072] Then, the electrodes were electrochemically cleaned by performing successive cathodic and anodic scans in NaOH and H 2 SO 4 solutions. First, in 0.5 M NaOH, the electrodes were cleaned by performing repetitive cyclic voltammetric scans within a potential window of -1 to -1.6 V (all potentials relative to Ag|AgCl) at a scan rate of 1 V s -1 for 300 cycles. Next, the electrodes were transferred into 0.5 M H 2 SO 4 solution, and an oxidative chronoamperometric step of 2 V was performed for 5 s. Subsequently, a reduction potential of -0.35 V was applied for 10 s, and then 10 voltammetric cycles were performed from -0.35 to 1.5 V at a scan rate of 4 V s -1 , followed by two cycles using the same potential range at a scan rate of 0.1 V s -1 .

[0073] Finally, the electrodes were washed in deionized water and their surface areas were determined by obtaining cyclic voltammograms from -0.35 to 1.5 V at a scan rate of 0.1 V s 2 SO 4 in 0.05 M H -1 solution.

[0074] The E-AB sensors were fabricated using the aforementioned target-assisted immobilization method (Liu et al., 2021, ACS Appl. Mater. Interfaces, 13:9491 - 9499). For this purpose, first, 4 μL of a 10 mM aqueous solution of tris(2-carboxyethyl)phosphine was added to reduce 2 μL of the as-received aptamer to free thiol (100 μM), and the tube was stored in the dark at room temperature for 1 hour. Then the reduced aptamer was diluted in the binding buffer (for the ATP aptamer, 10 mM HEPES pH 7.6, 100 mM NaCl, and 10 mM CaCl 2 ; for the glucose aptamer, 20 mM HEPES pH 7.6, 1 mM NaCl, 5 mM KCl, 10 mM MgCl 2 and 5 mM CaCl 2) to obtain a final concentration of 200 nM. The total aptamer concentration was kept constant at 200 nM in all experiments to control the DNA surface coverage. The target of the aptamer was added at a concentration of 0.01 mM before the electrochemically cleaned gold electrode was immersed in an aliquot of 150 μL of the aptamer - target mixture and incubated overnight at room temperature in the dark.

[0075] After that, the electrode was rinsed with 1X PBS solution (137 mM NaCl, 2.7 mM KCl, 10 mM Na 2 HPO 4 and 1.8 mM KH 2 PO 4 ) to remove any non - specifically adsorbed aptamers. Then the aptamer - functionalized electrode was immersed in the respective binding buffer containing 5 mM 6 - mercaptohexanol and the corresponding target at the same concentration used in the aptamer immobilization step and incubated overnight at room temperature in the dark. Finally, the electrode was thoroughly rinsed with 1X PBS to remove any non - specifically adsorbed target molecules and stored in the binding buffer at room temperature until use.

[0076] Saliva samples were collected according to the protocol of the ethics institution. Briefly, volunteers anonymously collected saliva in 50 mL Falcon tubes after natural salivation. The collected saliva was aliquoted into 5 mL volumes and stored at - 20 °C to avoid freeze - thaw cycles.

[0077] Since the signal of the E - AB sensor depends on the detection square - wave frequency, the response frequency was determined by measuring the function of the methylene blue peak current versus the square - wave frequency. For this purpose, at least three sensors were used to detect frequencies in the range of 2 to 1000 Hz in the working buffer and saliva in the absence and presence of the target. The collected peak currents were converted to signal change (%) using the following

[0078] Equation 1:

[0079]

[0080] Calibration curves were made in 10 mL of the respective working buffer for each aptamer and in 5 mL of undiluted saliva spiked with the target. For this purpose, square - wave voltammograms were collected from - 0.10 to - 0.45 V with an amplitude of 25 mV using the previously determined signal - off and signal - on frequencies.

[0081] First, in the absence of the target, the sensor was scanned 40 - 50 times until the peak current was stable. Then the sensor was exposed to gradually increasing amounts of the target (i.e., from 10 -20 to 10 -3M) Attack and record the methylene blue peak current. In the same electrochemical cell, a "control" electrode functionalized with a DNA sequence having the same number of nucleotides as the aptamer but no secondary structure was also attacked to evaluate the potential cross-reaction of the target with methylene blue. A Python-based script was used to extract the peak current at each target concentration, and then the resulting binding curve was non-linearly fitted to the Hill equation (Equation 2):

[0082]

[0083] where [X] is the target concentration, is the signal change at a given target concentration, is the baseline current observed in the absence of the target, is the signal change at the saturated target concentration, and K D is the aptamer-target dissociation constant.

[0084] Using the previously determined signal turn-on frequency described above, the sensor response time was determined by repeating the detection every 20 s in saliva without the target. After recording the sensor signal for 200 s, the target was injected into the saliva, and the electrochemical signal was continuously monitored for an additional 200 s.

[0085] The response of the E-AB sensor is square-wave frequency-dependent. This can be observed when the sensor is detected over a wide square-wave frequency range (i.e., 2 - 1000 Hz) in the presence or absence of a saturated target concentration. Doing so results in two different sensor responses, where one increases and the other decreases in the presence of the target ( Figure 2 ). The experiment was repeated in the presence of a saturating amount of the target (100 μM adenosine monophosphate and 100 μM glucose), and the relative signal change of the sensor was calculated using Equation 1 at each detection frequency. Doing so results in two frequency regions within which the sensor produces a lower or higher response to the addition of the target, referred to as "signal-off" or "signal-on", respectively. It was observed that using the sensor in undiluted saliva produced a greater signal change over the entire detection frequency range.

[0086] When attacking the ATP aptamer sensor with a saturating amount of adenosine monophosphate, the maximum signal-off and signal-on responses were obtained at 12 and 750 Hz, respectively. In contrast, the glucose sensor produced signal-off and signal-on responses at 10 and 150 Hz, respectively.

[0087] The E-AB sensor shows enhanced response in undiluted saliva. This was confirmed when the detection of sensors binding glucose and adenosine monophosphate was repeated over a wide frequency range ( Figure 2) This also resulted in a frequency-dependent sensor response, where the identified response frequency remained relatively unchanged compared to the frequency measured in buffer. When used in undiluted saliva, the relative signal change of both sensors increased. The signal ON response of the E-AB sensor binding glucose increased by more than 4-fold, while the signal change of the sensor binding adenosine monophosphate increased by approximately 25%.

[0088] When used in undiluted saliva, the detection limit of the E-AB sensor was enhanced. This was confirmed when detecting in saliva at the identified response square wave frequency and attacking the sensors binding adenosine monophosphate and glucose with increasing amounts of the corresponding target ( Figure 3 A and 3B). The sensor binding adenosine monophosphate showed a response in the dynamic picomolar-nanomolar range, with K D of 0.1 nM, while the sensor binding glucose responded in the femtomolar-nanomolar range, with K D of 300 nM. Figure 3 Figure A shows the results of testing the response signal ON frequency against increasing amounts of the target when using the aptamer binding adenosine monophosphate in undiluted saliva. The result of this was a maximum signal change of approximately 40%, with K D unprecedentedly at 0.1 nM. In contrast, when using the sensor in buffer, no response was measured, highlighting that the ionic conditions of saliva provide unique E-AB sensor performance. Considering that these results showed an unprecedented detection limit of the E-AB sensor for small molecules, control experiments were conducted where the titration experiment was repeated, in which the electrodes were functionalized with DNA sequences of the same length as the aptamer binding adenosine monophosphate (i.e., T26) and the aptamer binding glucose (i.e., T40) but without any secondary structure. In both cases, this did not result in a relative S-shaped binding, demonstrating the specificity of the sensor for adenosine monophosphate and glucose ( Figure 3 A and 3B). To verify that the response did not originate from other buffer components, saliva was titrated by providing increasing amounts of buffer to the sensor binding glucose ( Figure 3 Figure B). When used in saliva, the E-AB sensor binding glucose also showed enhanced analytical performance. When attacking the sensor with increasing amounts of glucose, a maximum signal change of approximately 45% was measured, with K D unprecedentedly at 300 nM. When attacking the sensor with increasing amounts of dilution buffer in saliva without the presence of glucose, no signal change was measured, supporting the specific binding of the aptamer. Similarly, no sensor response change indicating specific binding of the aptamer to its target was measured.

[0089] To further enhance the analytical performance of the glucose sensor, the parental aptamer sequence was redesigned by truncating and modifying the position of methylene blue. Specifically, base pairs were removed from both ends of the aptamer, generating two new variants (referred to herein as 3-trunc and 6-trunc) to reduce the distance separating the glucose-binding site from the electrode surface. This would allow for enhanced flexibility of the glucose aptamer, thus permitting an increased rate of methylene blue electron transfer and, in turn, improving the E-AB sensor signal. In addition to aptamer truncation, a glucose-binding variant was generated in which the original sequence was modified and methylene blue was conjugated internally within the sequence such that it could be closer to the glucose-binding site of the identified aptamer. When the variants of the glucose aptamer were assailed with increasing amounts of the target in undiluted saliva, all returned signal changes were higher compared to the parental variant. Although the affinity of the variants was greater than that of the parental variant (K D 3-trunc = 300 nM, K D 6-trunc = 15 nM, and K D MB内部位置 = 26 μM), the largest increases were observed in the responses of the 3-trunc and 6-trunc variants, which are better candidate variants for clinically relevant environments as their dynamic ranges are closer to the relevant concentrations.

[0090] As Figure 4 demonstrated, the E-AB sensor responds rapidly in the presence of the target.

[0091] As Figure 5 further demonstrated, as shown, the E-AB sensor remained stable over a long period of time (>3 days) when continuously detected in undiluted saliva. This test was conducted by immersing the glucose-binding E-AB sensor into undiluted saliva and recording square wave voltammograms at approximately 20-second intervals. In so doing, it was observed that the E-AB sensor response initially declined through an "exponential" phase, followed by a "linear" decay, after which the signal stabilized. As previously suggested in the literature, it was speculated that the former was due to the non-specific accumulation of saliva components at the E-AB sensor interface, resulting in a ~40% loss of the sensor. In contrast, the decay during the linear phase was attributed to the electrochemical detection at the sensor interface. Thiol adsorption on the electrode during cathodic detection has limited stability, which may trigger their desorption. Once past this stage, it was observed that the sensor maintained 50% of its original electrochemical signal for up to at least 3 days. Thus, the E-AB sensor was able to last longer and withstand the harsh conditions of undiluted saliva. It should be noted that the peaks in the decay corresponded to issues associated with the potentiostat, which was accidentally disconnected from the biosensor and should therefore be ignored.

[0092] Figure 6It is demonstrated that the E-AB biosensor can withstand cycles of direct and continuous exposure in undiluted saliva for potentially more than 10 rinse / target attack steps. First, the E-AB sensors conjugated with adenosine monophosphate (left panel) and conjugated with glucose (right panel) were used by exposure in undiluted saliva. Then the sensors were transferred to undiluted saliva containing 1 nM of the corresponding target, and the increase in signal was measured. Finally, the sensors were rinsed with 1X phosphate buffer and these steps were repeated 9 more times. It was found that when using the signal-on detection square wave frequency, the sensors produced a consistent 15 - 20% response over 10 cycles. This repetition can continue for more rinse / attack steps and demonstrates that the E-AB sensors are renewable. When looking at the signal-on response after each rinse step (bottom panel), the signal-on square wave frequency produced a consistent 20 ± 5% response for the E-AB sensor conjugated with adenosine monophosphate and a consistent 15 ± 5% response for the sensor conjugated with glucose over these 10 cycles. Thus, the E-AB sensors can be easily rinsed and produce a consistent response over multiple rinse / attack cycles.

[0093] As provided herein, the E-AB sensors allow for the measurement of glucose levels in saliva. Measuring glucose levels with the provided E-AB sensors allows for the monitoring of an increased risk of diabetes and gum disease. Higher glucose levels promote bacterial growth and increase the risk of the gums becoming damaged. Similarly, several other metabolites that can be used to monitor health status are also found in saliva, such as carnitine (e.g., 2-methylbutyrylcarnitine, butyrylcarnitine, isobutyrylcarnitine, propionylcarnitine), lactate / lactic acid, malate / malic acid, adenosine 5'-monophosphate (AMP), maltose, etc.

[0094] It is also provided that the described E-AB sensor can be modified to measure biomarkers or target molecules other than glucose, ATP, and / or AMP. As provided, the sensor is modified to measure vancomycin in saliva. Vancomycin is known as a glycopeptide antibiotic drug used to treat bacterial infections. Specifically, vancomycin is also taken orally to treat severe Clostridium difficile colitis. Measuring the level of vancomycin in saliva allows for proper dosing and therapeutic monitoring. As provided herein, the performance of a previously reported vancomycin-binding aptamer in buffer was tested and its ease of use in undiluted, unstimulated saliva was evaluated. Vancomycin-binding aptamers and their 4-trunc variants are known in the art and have been further modified in the E-AB sensor (see Dauphin-Ducharme et al., 2019, ACS Sens., 4, 10:2832–2837). However, the feasibility of applying such an E-AB sensor in undiluted saliva has not been demonstrated. Through this article, the analytical performance metrics (sensitivity, dynamic range, limit of quantification, limit of detection) of this biosensor in this complex matrix are provided.

[0095] The response of the E-AB sensor is square-wave frequency-dependent. This can be observed when the sensor is detected within a wide square-wave frequency range (i.e., 2 - 1000 Hz) in the presence or absence of a saturating target concentration. Doing so produces two different sensor responses, one of which increases and the other decreases in the presence of the target ( Figure 8 ). These response frequency ranges are respectively identified as "signal on" and "signal off". When attacking the vancomycin aptamer sensor with a saturating amount of vancomycin, the maximum signal-off and signal-on responses were observed to be obtained at 10 Hz and 100 Hz, respectively.

[0096] The response of the E-AB sensor is enhanced in undiluted saliva. This was observed when repeatedly detecting the vancomycin-binding sensor within a wide frequency range ( Figure 8 ). Doing so also produces a frequency-dependent sensor response, where the identified response frequencies remain relatively unchanged compared to those measured in buffer. However, it was observed that when the sensor is used in undiluted saliva, the relative signal change increases. The signal-on response of the E-AB sensor increased by 3 - 4 times.

[0097] When used in undiluted saliva, the limit of detection of the E-AB sensor is enhanced. This was observed when detecting the vancomycin-binding sensor in saliva at the above-identified response square-wave frequencies and attacking it with increasing amounts of the corresponding target ( Figure 9)。In contrast, the vancomycin E-AB sensor produced a signal gain of approximately 25%, and the dissociation constant (K D ) was approximately 17 μM. In saliva, compared to the observations in 1X PBS, the vancomycin E-AB sensor produced an improved sensor gain of approximately 107%, and the affinity K D value was approximately 12 μM. Control experiments were performed in which the titration experiment was repeated, and the electrodes were functionalized with a DNA sequence of the same length as the aptamer that binds vancomycin (i.e., T40) but without any secondary structure. This did not produce a relative sigmoidal binding, demonstrating the specificity of the sensor for vancomycin( Figure 9 ).

[0098] As further provided herein, by using the E-AB sensor to monitor the changes in the concentrations of glucose and adenosine monophosphate over time in saliva samples collected from the same individual at different times during the same day (see Figure 10 A and B). By doing so, when saliva samples were collected 30 minutes after lunch, changes in the responses of both sensors were observed. As time elapsed after lunch, the sensor responses of the sensors binding glucose and adenosine monophosphate decreased and increased, respectively, reflecting the concentrations of these molecules—high after digestion and then changing over time. The signal changes measured by the E-AB sensor were verified using a colorimetric test to correspond to changes in the endogenous concentration of glucose. Thus, the provided E-AB sensor allows monitoring of molecular level changes throughout the day. The sensor can be regenerated in 7M urea solution, then rinsed with deionized water, and stored in 1X PBS until the next use. Calibrating the signal of fasting saliva may be useful for measuring endogenous molecules.

[0099] Example I

[0100] Further Characterization of the E-AB Sensor

[0101] Since the E-AB sensor detected in undiluted saliva produced a greater signal change than when used in buffer, it was hypothesized that part of this difference in response was due to differences in the ionic composition of saliva. The E-AB sensor binding adenosine monophosphate was attacked with 1 mM adenosine monophosphate while increasing the concentration of divalent cations (Ca 2+ or Mg 2+ ) present in the phosphate buffer, and the resulting sensor response was recorded. In doing so, it was determined that the response of the sensor changed monotonically with the amount of divalent cation.

[0102] Then the circular dichroism technique was used because this technique is sensitive to changes in oligonucleotide chirality and, notably, has been used to monitor G-quadruplexes and other secondary structures.

[0103] Circular dichroism experiments were performed using a J-810 spectrophotometer (JASCO, MD). For this purpose, quartz cuvettes (optical path length of 1 cm) were used, in which the aptamer solution was resuspended in 1X phosphate buffer at pH 7.4. 1 mM adenosine monophosphate was added, and then divalent cations were added at different concentrations using a stock solution containing the same aptamer and buffer concentrations, and stirred for 1 minute to equilibrate. Each concentration was scanned for wavelength between 230 and 310 nm at a rate of 50 nm / min, with a bandwidth of 4 nm and a digital integration time of 4 seconds, and a binding curve was constructed using the ellipticity at 265 nm.

[0104] When the aptamer that binds adenosine monophosphate was placed in the presence or absence of adenosine monophosphate (1 mM), while increasing the concentration of divalent cations (Ca 2+ or Mg 2+ ), it was found that the ellipticity measured at 265 nm for the 10 μM aptamer solution increased monotonically (see Figure 7 ). Thus these results further confirmed that the addition of divalent cations (Ca 2+ or Mg 2+ ) that are not normally included in phosphate buffer solutions plays a key role in forming the tertiary structure of the aptamer that is thought to allow binding to the target. The unique combination of these ions in saliva further enhanced the response of the E-AB biosensor relative to the buffer, as more aptamers formed their binding-competing structures.

[0105] Example II

[0106] E-AB sensor for detecting vancomycin

[0107] The aptamer sequences used were purchased from Bio Basic Inc and purified by HPLC. These sequences were used without further purification and resuspended in deionized water to 100 μM, and then aliquoted into 2 μL tubes. Bio Basic Inc chemically modified the 5′ end of the aptamer with thiol chemistry on a 6-carbon linker, and chemically modified the 3′ end with methylene blue modified with a carboxyl group that was attached to the DNA by forming an amide bond with a primary amine on a 6-carbon linker.

[0108] Vancomycin 4-trunc aptamer: 5’HO-(CH 2 ) 6 -S-S-(CH 2 ) 6 -O-CGA GGG TAC CGC AAT AGTACT TAT TGT TCG CCT ATT GTG GGT CGG-O-(CH 2 ) 6 -NH-CO-(CH 2) 2 -MB-3’(SEQ ID NO:7)

[0109] T40 control sequence: 5’HO-(CH 2 ) 6 -S-S-(CH 2 ) 6 -O-ATT ATT TTT TAT TTA TTT TTA TTTTAT TTT ATT TTT TAT-O-(CH 2 ) 6 -NH-CO-(CH 2 ) 2 -MB-3’(SEQ ID NO:8)

[0110] Before fabricating the sensors, the electrodes (2.0 mm in diameter, CH Instruments) were mechanically cleaned by sequentially polishing with 1 μm diamond suspension oil slurry and then 0.05 μm alumina powder aqueous solution. After each polishing step, the electrodes were sonicated for 5 minutes in ethanol or distilled water, respectively.

[0111] Then the electrodes were electrochemically cleaned by performing successive cathodic and anodic scans in NaOH and H 2 SO 4 solutions. First, the electrodes were cleaned by performing repetitive cyclic voltammetry scans in 0.5 M NaOH within a potential window of -1 to -1.6 V (all potentials relative to Ag|AgCl) at a scan rate of 1 V s -1 for 300 cycles. Next, the electrodes were transferred into 0.5 M H 2 SO 4 solution, and a 2 V oxidative chronoamperometry step was performed for 5 s. Subsequently, a reduction potential of -0.35 V was applied for 10 s, and then 10 voltammetric cycles were performed from -0.35 to 1.5 V at a scan rate of 4 V s -1 , followed by two cycles using the same potential range at a scan rate of 0.1 V s -1 . Finally, the electrodes were rinsed in deionized water and the cyclic voltammogram was obtained from -0.35 to 1.5 V at a scan rate of 0.1 V s 2 SO 4 in 0.05 M H -1 solution to determine their surface area.

[0112] The E-AB sensor was fabricated using the aforementioned target-assisted immobilization method (see Liu et al., 2021, ACS Appl. Mater. Interfaces, 13, 8: 9491-9499). To this end, first, 4 μL of 10 mM aqueous solution of tris(2-carboxyethyl)phosphine was added to reduce 2 μL of the as-received aptamer to free thiol (100 μM), and the tube was stored in the dark at room temperature for 1 hour. Then the reduced aptamer was diluted in binding buffer (10 mM HEPES pH 7.6, 25 mM NaCl, 12.5 mM KCl, 1 mM MgCl 2 and 2.5 mM CaCl 2 ) to obtain a final concentration of 200 nM. The total aptamer concentration was kept constant at 200 nM in all experiments to control the DNA surface coverage. Before immersing the electrochemically cleaned gold electrode into an aliquot of 150 μL of the aptamer-target mixture for overnight incubation in the dark at room temperature, the target of the aptamer was added at a concentration of 0.25 mM.

[0113] After that, the electrode was rinsed with 1X PBS solution (137 mM NaCl, 2.7 mM KCl, 10 mM Na 2 HPO 4 and 1.8 mM KH 2 PO 4 ) to remove any non-specifically adsorbed aptamers. Then the aptamer-functionalized electrode was immersed in the respective binding buffer containing 5 mM 6-mercaptohexanol and the corresponding target at the same concentration used in the aptamer immobilization step for overnight incubation in the dark at room temperature. Finally, the electrode was thoroughly rinsed with 1X PBS to remove any non-specifically adsorbed target molecules and stored in the binding buffer at room temperature until use.

[0114] Saliva samples were collected according to the protocol of the ethics institution. Briefly, volunteers anonymously collected saliva in 50 mL Falcon tubes after natural salivation. The collected saliva was aliquoted into 5 mL volumes and stored at -20 °C to avoid freeze-thaw cycles.

[0115] Since the signal of the E-AB sensor depends on the detection square wave frequency, the response frequency was determined by measuring the function relationship between the methylene blue peak current and the square wave frequency. To this end, in the absence and presence of the target, at least three sensors were used to detect the frequencies in the range of 2 to 1000 Hz in the working buffer and saliva. The collected peak currents were converted to signal changes (%) using Equation 1.

[0116] Calibration curves were performed in 10 mL of the respective working buffer of each aptamer and 5 mL of undiluted saliva spiked with the target. To this end, square wave voltammograms were collected from -0.10 to -0.45 V with an amplitude of 25 mV using the previously determined signal-off and signal-on frequencies.

[0117] First, in the absence of the target, the sensor was scanned 40 - 50 times until the peak current stabilized. Then the sensor was challenged with gradually increasing amounts of the target (i.e., from 10 -11 to 10 -4 M), and the methylene blue peak current was recorded. In the same electrochemical cell, a "control" electrode functionalized with a DNA sequence having the same number of nucleotides as the aptamer but no secondary structure was also challenged to evaluate potential cross-reactivity of the target with methylene blue. Peak currents were extracted at each target concentration using a Python-based script, and the resulting binding curves were then non-linearly fit to the Hill equation (Equation 2).

[0118] As Figure 8 shown, the signal change of the E-AB sensor (relative signal change after addition of saturating target) is a strong function of the detection square wave frequency. This was observed when detecting the vancomycin aptamer over a wide frequency range in buffer (black curve) and undiluted human saliva spiked with saturating target concentration (colored curve). The relative signal change of the sensor was calculated at each detection frequency using Equation 1. Doing so resulted in two frequency regions where the sensor produced a lower or higher response upon addition of the target, termed "signal-off" or "signal-on", respectively. It was found that deploying the sensor in undiluted saliva produced a greater signal change across the entire detection frequency range.

[0119] Example III

[0120] Characterization of the E-AB Sensor

[0121] Saliva samples were collected according to the protocol of the ethics institution (2021 - 3080). Briefly, volunteers anonymously collected saliva in 50 mL Falcon tubes after natural salivation. Fasting saliva samples were collected after waking up in the morning and before brushing teeth or using any oral care products such as mouthwash. Saliva samples were subjected to other measurements as soon as possible after collection.

[0122] Saliva samples were centrifuged at 14000 rpm for 5 minutes at room temperature. The supernatant was collected and the glucose concentration in the saliva was determined using a colorimetric EnzymChrom TM glucose assay kit (EBGL - 100) according to the described protocol.

[0123] When immersed in different saliva samples, the sensor binding adenosine monophosphate and glucose was interrogated at the signal-off and signal-on frequencies (10 and 400 Hz) once. After each measurement, the sensor was immersed in 7 M urea solution for 10 min, then rinsed thoroughly with DI water and then stored in 1X PBS until the next use. We used Equation 3 to derive the kinetic differential measurements from the peak currents obtained at the two frequencies.

[0124]

[0125] The sensor signals at different time points were normalized against the response in fasting saliva. In the same electrochemical cell, we fabricated and interrogated electrodes functionalized with a DNA sequence with the same number and ratio of nucleotides as the aptamer binding glucose or adenosine monophosphate but with a scrambled sequence (labeled as “scrambled sensor” in Figure 10 as a negative control.

[0126] Although this specification is described in connection with its particular embodiments, it is to be understood that it is capable of further modification, and this application is intended to cover any variations, uses, or adaptations and include such departures from the present disclosure as come within known or customary practice and such departures may apply to the essential features hereinbefore set forth and fall within the scope of the appended claims.

Claims

1. A sensor for measuring the level of a target molecule in a fluid sample, which comprises: a sensing element comprising a redox reporter, the redox reporter being linked to at least one aptamer that binds to a salivary metabolite; and a solid surface, the sensing element being linked to the solid surface, wherein when the salivary metabolite binds to the aptamer, the redox reporter undergoes a binding-induced change in electron transfer.

2. The sensor element according to claim 1, wherein the aptamer is linked to the solid surface via a thiol molecule.

3. The sensor element according to claim 1 or 2, wherein the redox reporter is a methylene blue molecule.

4. The sensor element according to any one of claims 1-3, wherein the salivary metabolite is glucose.

5. The sensor element according to any one of claims 1-3, wherein the salivary metabolite is vancomycin.

6. The sensor element according to any one of claims 1-5, which further comprises at least one aptamer that binds to a different target molecule.

7. The sensor element according to any one of claims 1-6, wherein the target molecule is adenosine triphosphate (ATP), carnitine, lactate, lactic acid, malate, malic acid, maltose or adenosine monophosphate (AMP).

8. The sensor element according to claim 7, wherein the carnitine is 2-methylbutyrylcarnitine, butyrylcarnitine, isobutyrylcarnitine or propionylcarnitine.

9. The sensor element according to any one of claims 1-8, wherein the solid surface is an electrode.

10. The sensor element according to any one of claims 1-9, wherein the electrode is a gold electrode.

11. The sensor element according to any one of claims 1-10, wherein the fluid sample is saliva.

12. The sensor element according to any one of claims 1-11, wherein the sensor element is regenerated for multiple measurements.

13. A kit comprising the sensor element according to any one of claims 1-12 and instructions for use.

14. The kit according to claim 13, which further comprises a monitoring device.

15. The kit according to claim 14, wherein the monitoring device monitors the binding-induced change in electron transfer of the redox reporter.

16. The kit according to claim 15, wherein the electron transfer is monitored by a change in the peak height of a square wave voltammogram.

17. The kit according to claim 14, wherein the monitoring device monitors the binding-induced change in electron transfer by voltammetry or impedance spectroscopy.

18. A method for measuring the level of at least one metabolite in a patient's fluid sample, which comprises the steps of: contacting the fluid sample with the sensor according to any one of claims 1-12; and monitoring the binding-induced change in electron transfer of the redox reporter to indicate the presence of at least one metabolite in the fluid sample.

19. The method according to claim 18, wherein the sample is saliva.

20. The method according to claim 18 or 19, wherein the electron transfer is monitored by a change in the peak height of a square wave voltammogram.

21. A method according to any one of claims 18 - 20, wherein at least one metabolite is glucose, vancomycin, adenosine triphosphate (ATP), carnitine, lactate, lactic acid, malate, malic acid, maltose or adenosine monophosphate (AMP).

22. A method according to claim 21, wherein the carnitine is 2 - methylbutyrylcarnitine, butyrylcarnitine, isobutyrylcarnitine or propionylcarnitine.

23. A method according to any one of claims 18 - 22, wherein measuring the level of at least one metabolite allows monitoring of the health status of a patient.

24. A method according to claim 23, wherein measuring the glucose level indicates the presence or progression of diabetes in the patient or an increased risk of gum disease.

25. A method according to any one of claims 18 - 24, which is for measuring at least one metabolite over a long period of time using a sensor.

26. A method according to claim 25, which is for measuring at least one metabolite using a sensor for more than three days.

27. A method according to claim 25, wherein the metabolite is measured before and / or after the patient eats.

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