Electrochemical sensors

Abstract

The present disclosure relates to electrochemical sensors that use fluoroorganothiols or fluoroorganosilanes in the formation of self-assembled monolayers (SAMs) for use in diagnostic tests. Methods of testing patient samples using the disclosed electrochemical sensors are also provided.

Description

Technical Field

[0001] The present disclosure relates to an electrochemical sensor for use in diagnostic tests. A method for testing a patient sample using the disclosed electrochemical sensor is also provided.

Background Art

[0002] Electrochemical biosensors are a promising route to achieve rapid and sensitive detection of a wide range of pathogens and clinically important biomarkers. 1 The best-known example is the glucose biosensor (most commonly an amperometric sensor), which is widely used for home testing of blood glucose levels and is very useful for diabetic patients in routine monitoring of blood glucose levels. 2 Cyclic voltammetry (CV) 4 Linear sweep voltammetry (LSV) 5 Electrochemical impedance spectroscopy (EIS) 6 And differential pulse voltammetry (DPV) 7 Many other biosensors have been developed that operate by a wide range of principles, including 3 EIS involves a measurement setup in which a small AC excitation potential is imposed on the working electrode (often under open circuit potential), and the resulting current response of the electrochemical cell is measured. A variety of parameters associated with the cell and its response can be extracted from the EIS response, including solution resistance (R s ), double layer capacitance (C DL 0), charge transfer resistance (R CT) and Warburg impedance (W) 8 are included. The double-layer capacitance and charge transfer resistance have been shown to be particularly effective for label-free monitoring of binding at biologically functionalized electrode surfaces. These techniques enable highly sensitive and specific measurements of DNA and protein biomarkers, as has been repeatedly shown in the literature 9 .

[0003] For many electrochemical biosensors such as those described above, surface functionalization and attachment chemistry play important roles in sensor design and performance 10,11 . For gold sensors, the attachment of biological molecules is often carried out through the use of gold-thiol attachment, more specifically through the formation of self-assembled monolayers (SAMs) 12 . SAMs serve a dual purpose of blocking the electrode surface from non-specific binding of proteins, cells and other components in the sample medium and ensuring the correct orientation of bio-recognition elements (e.g., DNA sequences, antibodies or enzymes) 13 . Self-assembled monolayers are often formed by incubation of a gold surface with a solution of thiolated biomolecules and may contain single molecules (monolayers) or multi-component systems with additional complexity introduced to ensure proper orientation of the receptor and excellent resistance to surface fouling SUMMARY OF THE INVENTION

[0004] This disclosure relates to research conducted by the researchers on the development of sensors for diagnostic use, based on studies using fluoroorganothiol or fluoroorganosilane molecules for the formation of SAMs. In one embodiment, the SAM is formed using 1H,1H,2H,2H-perfluorodecanethiol (PFDT) on the sensor surface, which is then biofunctionalized with a biological agent. PFDT spontaneously forms a high-density hydrophobic SAM on a gold surface, reducing surface biofouling. 14 PFDTs were previously used to enhance the performance of organic transistors. 15 , reversibly structuring DNA on a micropatterning substrate 16 It has been used for that purpose.

[0005] In the first teaching, the present disclosure relates to an electrochemical biosensor for use in detecting a target analyte, The present invention provides an electrochemical biosensor comprising at least one detection electrode having a surface coated with a self-assembled monolayer (SAM), wherein the SAM contains, is essentially composed of, or consists of hydrofluorocarbon or fluorocarbon molecules bound to the electrode surface via reactive sulfur or silicon groups present on the hydrofluorocarbon or fluorocarbon.

[0006] Hydrofluorocarbons are organic compounds containing fluorine and hydrogen atoms. Fluorocarbons are compounds in which all CH bonds are replaced by CF bonds. Hydrofluorocarbons, or fluorocarbon molecules, can take the form of linear, branched, or cyclic alkanes, alkenes, or alkynes, having one or more reactive sulfur or silicon groups.

[0007] In one embodiment, the reactive sulfur group(s) may be a thiol, and in such an embodiment, the molecule may be a fluoroorganothiol. In one embodiment, the reactive silicon group(s) may be a silane, and in such an embodiment, the molecule may be a fluoroorganosilane.

[0008] In one embodiment, the fluorocarbon molecule is a linear fluoroalkanethiol or fluoroalkanesilane.

[0009] Examples of compounds suitable for use in this disclosure include: 1H,1H,2H,2H-perfluorodecanethiol 3,3,4,4,5,5,6,6,7,7,8,8,8-Tridecafluoro-1-Octanethiol 3,3,4,4,5,5,6,6,7,7,8,8,9,9,10,10,10-heptadecafluoro-1-decanethiol 3,3,4,4,5,5,6,6,6-nonafluoro-1-hexanethiol 2,2,2-trifluoroethanethiol 1H,1H,2H,2H-Perfluorooctyltriethoxysilane 1H,1H,2H,2H-perfluorodecyltriethoxysilane, and 1H,1H,2H,2H-Perfluorododecyltrichlorosilane It includes.

[0010] In one embodiment, the compound used to form the SAM is 1H,1H,2H,2H-perfluorodecanethiol.

[0011] The electrodes of this disclosure may be made of any suitable conductive material, which may be coated with hydrofluorocarbon or fluorocarbon molecules as described herein. Suitable materials include glassy carbon; metal oxides; conductive polymers; and precious metals, including gold, ruthenium, rhodium, palladium, platinum, and silver. In one particular embodiment, the electrode is gold.

[0012] Self-assembled monolayers (SAMs) are typically self-structured layers of amphiphilic molecules, in which one end of the molecule exhibits specific affinity to a substrate material. SAM molecules may include a head group that anchors the molecule to the substrate, as well as a terminal tail portion or functional group. SAM layers can be formed, for example, by chemiadsorption of head groups, such as reactive sulfur (e.g., thiols) and silicon (e.g., silanes), onto the substrate material from a vapor or liquid phase. The tail portion groups of SAMs according to this disclosure are known fluoresceins. 17 This material is inherently hydrophobic and, as described in more detail herein, enables the capture or physical adsorption of a variety of biological agents.

[0013] Preferably, the hydrophobic nature of the SAM not only allows suitable biological agents to be captured by the SAM, but also provides a barrier layer that works to reduce fouling and / or interference by other substances that may be present with the sample, such as proteins, cells, and other molecules. The sample may be obtained from a subject and may be blood, saliva, or any other suitable biological fluid, such as urine, semen, or tissue samples. Alternatively, the sample may be an environmental sample, such as a water sample, a soil sample, or even a plant sample.

[0014] In further teaching, the disclosure further provides an electrochemical biosensor as described in the first teaching and embodiments, further comprising a biological agent captured by a SAM layer coated on the electrode surface, for example, by physical adsorption.

[0015] Further teachings, the present disclosure provides a method for producing an electrochemical biosensor as described in the first or further teachings and embodiments described above: A step of forming a SAM on the surface of at least one detection electrode by bringing a solution containing an organic solvent and hydrofluorocarbon or fluorocarbon molecules as described herein into contact with the surface of at least one detection electrode, A step of evaporating the solvent to allow the SAM to form on the surface of at least one detection electrode; Optionally, the process then involves bringing the SAM-coated electrode into contact with a solution containing a biological agent, allowing the biological agent to be captured by the SAM layer coated on the electrode. This provides a method that includes [something].

[0016] In some embodiments of this disclosure, the electrochemical biosensor further includes at least one reference electrode and / or counter electrode, which may be electrically coupled to the at least one detection electrode, for example, through a measurement system or a coupling for connecting to a measurement system. The electrodes are typically on a substrate and may be provided in the form of, for example, screen-printed electrodes; microelectrodes; printed circuit boards; FETs / OFETs, etc. The electrochemical biosensor may include or be provided with a coupling for connecting to a measurement system. The measurement system may be configured to apply an electrical signal between at least one detection electrode and at least one reference and / or counter electrode. The measurement system may be configured to measure the electrical reaction resulting from the detection of an analyte by a biological agent. The measurement system may be configured to perform impedance, voltammetry, or amperometry measurements. The measurement system may be configured to perform electrochemical impedance spectroscopy (EIS).

[0017] An electrochemical biosensor in accordance with this disclosure may be used to detect a target analyte that is bound to, typically specifically, a biological agent captured by the sensor. The target analyte may be a chemical substance (e.g., hormone, narcotic, or contaminant) or a biological molecule, such as a peptide, protein, glycoprotein, enzyme, glycolipid, cell surface receptor, cytokine, antibody, nucleic acid, etc. The target analyte may be free within the sample to be analyzed, or it may remain as part of a cell, cell membrane, viral coat, etc., where the target analyte, such as a biological molecule, is typically found in situ. The binding between the biological agent and the target analyte involves an attractive binding (i.e., non-repulsive) of two or more species held together by attractive forces. Such binding involves interactions between the biological agent and the target analyte that pull (or attract) the biological agent and the target analyte together in the sample. Bonding includes, but is not limited to, covalent interactions; electrostatic interactions; ion-ion interactions, e.g., those between attractive or opposite charges; dipole interactions; ion-dipole interactions; hydrogen bonding interactions; van der Waals interactions; π-stacking interactions; electron density sharing or combinations thereof.

[0018] The electrochemical sensors of this disclosure provide a signal conversion method in which any binding of a target analyte to a biological agent can be converted into a signal for processing and / or display. The biological agent may be similar to the target analyte and may include proteins, enzymes, antibodies, nucleic acids, etc. The biosensor may be configured to use specific chemical interaction properties (e.g., enzymes and their substrates) or molecular recognition mechanisms (e.g., protein binding to a receptor or antibody binding to an antigen) to identify the target analyte. The biosensor may use electrodes to convert the electrical signal resulting from the detection of the analyte by the biological agent into a different signal that can be expressed optically, electronically, or by other means.

[0019] In one embodiment, impedance, voltammetry, or amperometry measurements may be performed by an electrochemical sensor to detect a signal or a change in a signal. In one embodiment, electrochemical detection of a target analyte is performed by electrochemical impedance spectroscopy (EIS). In this method, the electrochemical biosensor of the present disclosure converts changes in interfacial properties between electrodes and electrolytes induced by biological macromolecule binding to the target analyte into an electrical signal. Typically, a redox pair, e.g., K3[Fe(CN)6] / K4[Fe(CN)6], may be used as a redox index for electrode dynamics at the interface. Sensors based on the use of EIS detection are label-free and therefore have the advantages of being low-cost, simple, and easily miniaturized. EIS is sensitive to surface interactions and the interfacial charge transfer resistance (R) associated with the charged redox probe. CT It is particularly useful for quantifying ). CT This is strongly influenced by changes in the charge distribution near the electrode-solution interface in the sample solution, which gives rise to surface sensitivity.

[0020] The biosensors of this disclosure find applications in many different diagnostic applications, including, for example, the detection of infectious pathogens, such as bacteria and viruses (including influenza and SARS-CoV-2); molecules associated with sepsis, such as IL-16 and procalcitonin; cardiac biomarkers, such as troponin; and liver biomarkers. Preferably, the detection principle using the sensors of the present invention has broad applicability because it is relatively easy to manufacture and provides rapid detection.

[0021] In one non-limited embodiment, this disclosure relates to the detection of COVID-19. In this example, the biological macromolecule captured by the electrochemical biosensor is ACE-2, which has been identified as the entry receptor for SARS-CoV-2, the virus that causes COVID-19. Thus, in one embodiment, an electrochemical biosensor according to this disclosure, in which the biological macromolecule is ACE-2, can be used to detect SARS-CoV-2, for example, COVID-19, or a COVID-19 specific antibody. 18 For example, an antibody that can specifically bind to the spike protein of COVID-19 can be used as a biological macromolecule. In this approach, the target analyte to be detected is the SARS-CoV-2 virus or, more specifically, the COVID-19 virus or its coat protein.

[0022] The chemical terms used herein, unless otherwise explicitly stated, have their standard meanings as known in the art, in accordance with the IUPAC Goldbook. Throughout the specification and claims, unless the context clearly requires otherwise, the words “comprise,” “comprising,” etc., are considered to have an inclusive meaning, as opposed to an exclusive or exhaustive meaning, i.e., “includes, but not limited to.”

[0023] This disclosure is further described herein by reference to the following figures. [Brief explanation of the drawing]

[0024] [Figure 1] (A) Image of an 8x Au working electrode PCB including an on-chip Au counter electrode and a reference electrode. (B, C, D, E) Depictions of the Au sensor surface in the following states: clean (B), PFDT functionalized (C), and ACE2 functionalized (D), as well as with binding of SARS-CoV-2 spike protein or inactivated virus (E). (F) Example of a Nyquist plot showing the signal of an ACE2 functionalized sensor (black) and after exposure to recombinant SARS-CoV-2 spike protein (red). [Figure 2] (A) Examples of Nyquist plots from representative electrodes after purification (black), PFDT functionalization (red), and ACE2 incubation (blue). (B) Box plot showing Rct values ​​through the three stages of electrode functionalization (purification, SAM formation, and ACE2 fixation). (C) Protein structure of ACE2 (1R42)19. (D) Structural formula of PFDT. [Figure 3] (A) Box plot showing normalized Rct values ​​for ACE2-functionalized electrodes versus HRP-conjugated SARS-CoV-2 spike protein solution and HRP-conjugated streptavidin solution. (B) Nyquist plot showing the impedance measurement response to the increase in HRP-conjugated SARS-CoV-2 spike protein. (C) Bar graph showing the ΔRCT% change in response to the addition of HRP-conjugated SARS-CoV-2 spike and HRP-conjugated streptavidin protein. (D) Dose-response curve for HRP-conjugated SARS-CoV-2 spike protein. (E) Protein structures of SARS-CoV-2 spike protein (6XM4)20 and streptavidin (4BX5)21. [Figure 4] (A) Nyquist plot showing impedance measurement response to increase in HRP-conjugated SARS-CoV-2 spike protein. (B) Bar graph showing ΔRCT% change in response to addition of HRP-conjugated SARS-CoV-2 spike and IL-6. (C) Box plot showing normalized Rct values ​​for ACE2-functionalized electrode versus HRP-conjugated SARS-CoV-2 spike protein solution and IL-6 solution. (D) Dose-response curve for HRP-conjugated SARS-CoV-2 spike protein. (E) Protein structure of IL-6(2IL6)24. [Figure 5](A) Nyquist plot showing impedance measurement response to increasing inactivated SARS-CoV-2 virus concentration. (B) Bar graph showing ΔRCT% change in response to the addition of negative and positive samples of inactivated SARS-CoV-2. (C) Box plot showing normalized Rct values ​​for ACE2-functionalized electrodes versus positive and negative samples of inactivated SARS-CoV-2. (D) Dose-response curve for inactivated SARS-CoV-2. (E) SARS-CoV-2 structure (fitted from image by Maya Peters Kostman for the Innovative Genomics Institute. https: / / creativecommons.org / licenses / by-nc-sa / 4.0 / legalcode). [Figure 6] The impedance change characteristic of increased charge transfer resistance during spike protein binding, similar to that observed with PFDT-modified electrodes, is not observed when the underlying SAM layer is composed of 1-octanthiol or 1-undecanethiol. [Modes for carrying out the invention]

[0025] material and method

[0026] Abbreviation PFDT, 1H,1H,2H,2H-perfluorodecanethiol; ACE2, angiotensin-converting enzyme 2; IL-6, interleukin-6; SARS-CoV-2, severe acute respiratory syndrome coronavirus 2; HRP, horseradish peroxidase; EIS, electrochemical impedance spectroscopy; ARDS, acute respiratory distress syndrome; PCB, printed circuit board; PBS, phosphate-buffered saline; SAM, self-assembling monolayer; R ct , charge transfer resistance; OCP, open circuit potential; IQR, interquartile range.

[0027] Chemical substances: K3[Fe(CN)6], K4[Fe(CN)6], 1H,1H,2H,2H-perfluorodecanethiol, KOH, and H2O2 30% (v / v) were obtained from Sigma-Aldrich. Toluene was obtained from Fisher Scientific UK Ltd (Loughborough, UK). Deionized water (5.00 μS / cm, 25℃) was purchased from Scientific Laboratory Supplies Limited (Nottingham, UK). Inactivated SARS-CoV-2 and negative control were obtained from Randox laboratories Ltd (Crumlin, UK). ACE2 was purchased from Abcam (Cambridge, UK), HRP-conjugated spike protein from The Native Antigen Company (Oxford, UK), and HRP-conjugated streptavidin was purchased as part of the IL-6 diagnostic kit from Bio-techne (Abingdon, UK).

[0028] Preconditioning. The SEP1 BIOTIP multichannel electrode PCB platform (biotip ltd, Bath, UK) was cleaned according to the supplied protocol. This consisted of immersion in a 50 mM KOH solution in 30% (v / v) H2O2 for 15 minutes at room temperature. The PCB was then rinsed with DI water and dried using compressed air. The PCB was then electrochemically cleaned by immersion in 50 mM KOH (DI water as solvent) along with an external platinum counter electrode (Metrohm, Runcorn, UK) and a 3M NaCl Ag / AgCl reference electrode (IJ Cambria, Llanelli, UK). Cyclic voltammetry was performed on all working electrodes on the PCB using the following parameters: potential window of -1.2 to 0.6 V, scan rate of 0.1 V / s, and 15 scans per electrode. The PCB was then rinsed with DI water and dried again using compressed air. In both cases, all electrochemical measurements were performed using a PalmSens4 potentiostat and accompanying PSTrace software supplied by Palmsens BV (Houten, Netherlands).

[0029] Fluorocarbon SAM and ACE2 immobilization. SAM solutions were prepared by magnetically stirring toluene and adding PFDT until a 1 mM solution was formed. Stirring helps to disperse PFDT throughout the solution. Fluorocarbons may have low miscibility in organic solvents, tend to self-interact, and form separate phases through the fluorocarbon effect. 17 The PCBs were oriented horizontally in small glass petri dishes, and PFDT solution was added to cover the PCBs with an excess of the solution. Since toluene evaporates rapidly, an excess of the solution was used and covered with a film to reduce evaporation loss. The PCBs were incubated overnight at room temperature, then rinsed with DI water (10 seconds of water flow per electrode), and dried with compressed air. All operations with toluene were performed in a suitable fume hood with a proper halogenated solvent waste disposal route.

[0030] ACE2 was diluted from stock in 1x PBS to 1 μg / ml, and 10 μl of alicot was applied to each working electrode on the PCB and incubated at room temperature for 1 hour. After incubation, the PCB was rinsed with 1x PBS (water bottle flow for 10 seconds per electrode) and dried with compressed air.

[0031] Protein target detection. To investigate evidence of specific binding between ligand (ACE2) and protein (HRP-conjugated SARS-CoV-2 spike protein), a series of dilutions of positive control HRP-conjugated SARS-CoV-2 spike protein and negative controls of similar size (HRP-conjugated streptavidin and IL-6) were incubated on a PCB sensor array at room temperature for 30 minutes, rinsed with 1x PBS (water bottle flow for 10 seconds per electrode), and EIS measurements were performed between incubations for each concentration. The concentrations of HRP-conjugated SARS-CoV-2 spike protein and IL-6 used were 1, 10, 50, and 100 ng / ml (all dilutions in 1x PBS). HRP-conjugated streptavidin was obtained as part of the ELISA kit, and its concentration was not disclosed. The accompanying instructions recommended a 1:40 dilution for the ELISA assay. The series of dilutions used (1:100, 1:75, 1:50, 1:25, and 1:5) were distributed around the recommended 1:40 dilution.

[0032] Detection of inactivated virus. For the detection of inactivated virus, the SARS-CoV-2 Clinical Molecular Standard Kit (Qnostics) was purchased. The kit contained positive and negative samples of the virus present in a complex "transport medium" that served as a substitute for clinical samples. A series of dilutions of the positive control (inactivated virus + transport medium and human cells) were incubated on a PCB at room temperature for 30 minutes. The concentrations used were 10 2 , 10 3 , 10 4 , 10 5 and 10 6The value was dC / ml (digital copies per ml). Due to the small volume of solution provided, a negative control (transport medium + human cells) was incubated twice at room temperature for 30 minutes. Room temperature incubation was chosen to replicate the operating environment conditions likely to be required for the diagnostic device. The PCB was rinsed with 1xPBS (water bottle flow for 10 seconds per electrode), and EIS measurements were performed between each incubation.

[0033] EIS parameters. All EIS measurements used the following parameters: E ac = 0.01Vrms, E dc Measurements were performed against the open-circuit potential (OCP) at 0V, with a frequency range of 50 frequencies and a 9.8 / decade interval of 100kHz to 1Hz. All measurements were obtained using 5mM K3[Fe(CN)6] / K4[Fe(CN)6] in 1xPBS.

[0034] Results and Discussion Fluorocarbon SAM Functionalization. Conventional electrochemical biosensors would have probe molecules surrounded by a hydrocarbon-based SAM, directly attached to the sensor surface (through covalent bonding, physicoadsorption, and chemiadsorption). Less common is the first fixation of the hydrocarbon SAM, to which biomolecules adsorb via hydrophobic physicoadsorption interactions, which can achieve better orientation of the probe biomolecule and potentially increase receptor-target binding. However, these approaches are weaker fixation methods than covalent bonding and therefore have the disadvantage of being more likely to be removed during incubation and washing processes. These researchers sought to explore the use of fluorocarbons because they offer significantly increased amphiphilicity (hydrophobic and oleophobic properties) compared to hydrocarbons, providing stronger physicoadsorption and anti-biofouling properties. 17 The ability of fluorocarbons to form SAMs on PCB electrode surfaces was investigated. Overnight incubation with 1 mM PFDT affected the increase in the measured impedance of the electrode, which was greater than the clean impedance in the Nyquist plot. ctThis is evident as a semicircle (SAM) (Figure 2A). Quantitatively, this corresponds to the average R in Figures 2B and 2C. ct Clean electrode with Ω = 2.5kΩ and SAM stage R ct =13kΩ, 928% R ct This can be seen as an average percentage increase.

[0035]

number

[0036] The percentage change was calculated using Equation 1. In the equation, Δ% is the percentage change, and R ct前 This is the initial stage of R ct And R ct後 R in the incubation stage ct Significant differences were assessed from the box plots. A significant difference between groups is likely if the median of one group is outside the interquartile range (IQR) of the other group. T-test analysis was not performed because these experiments were not designed as hypothetical tests and could report erroneous results as such. The box plots (Figure 2C) showed that the clean and SAM stages could be significantly different because the IQRs of the two groups did not overlap, and therefore the PFDT layer formed during SAM formation caused a significant increase in electrode impedance, providing strong evidence for fixed PFDT layer formation. This was hypothesized to be due to a well-established process of SAM formation with SAM molecules adhering to the surface and ordered layer formation on the surface. Such layers are redox-active Fe(CN)6 in the measurement buffer. 3- / 4- This can limit the amount or rate at which ions may undergo redox reactions, leading to an increase in impedance measurements. In summary, these data demonstrate the successful formation of fluorocarbon SAMs on PCB electrodes.

[0037] ACE2 hydrophobic immobilization. A further advantage of strongly hydrophobic fluorescein SAMs is that they provide an ideal environment for promoting hydrophobic physicoadsorption of ACE2 biomolecules. To test this, ACE2 proteins were incubated in the presence of electrode SAMs. After incubation for 1 hour with a 1 μg / ml ACE2 solution on the SAM-functionalized electrode, a slight increase in impedance was evident (Figure 2A). Absolutely, this was a further 2 kΩ Rct increase compared to SAM alone (Figure 2C). This was an indicator of hydrophobic physicoadsorption of the enzyme into the supporting SAM layer. The ACE2 electrode differed significantly from the clean group but not significantly from the SAM group. This finding was not entirely unexpected, as fluorescein SAMs cover a previously clean surface with a densely packed layer that produces a large impedance change. ACE2 was adsorbed into the fluorescein SAM and further added to this layer by blocking the electrode surface; however, because there were relatively fewer ACE2 molecules compared to the fluorescein molecules present on the surface, it caused a small relative change in impedance. Since the successful assembly of the PFDT SAM was demonstrated and evidence of ACE2 incorporation into the SAM structure was observed, a series of ligand binding experiments were then performed.

[0038] HRP-conjugated SARS-CoV-2 spike protein (positive) and HRP-conjugated streptavidin protein (negative). Since successful fixation of ACE2 was confirmed, HRP-conjugated SARS-CoV-2 spike protein (the HRP-conjugated form was used to allow for visual determination of binding) was incubated with the functionalized sensor surface for 30 minutes. Measurement impedances at 1, 10, 50, and 100 ng / ml consistently increased compared to the previous concentration, showing dose-dependent behavior (Figure 3A). ctThe mean percentage change (n=4) ranged from 96% of the lowest concentration to 156% of the highest concentration (Figure 3B, red). This indicated that the HRP-conjugated SARS-CoV-2 spike protein bound to the PFDT-ACE2 modification sensor. The addition of a series of dilutions of HRP-conjugated streptavidin (negative controls 1:100, 1:75, 1:50, 1:25, and 1:5) allowed for confirmation of the specific binding of the HRP-conjugated SARS-CoV-2 spike protein. ctThe mean percentage change (n=4) ranged from 6.2% at the lowest concentration to 52.8% at the highest concentration (Figure 3B, blue). Negative responses appeared to plateau in two consecutive percentage change measurements at 1:25 (53.1%) and 1:5 (52.8%), with two similar data spreads and values ​​(Figure 2C, blue). All normalized data in these experiments used the ACE2 signal as the normalization factor. There appeared to be a significant difference between ACE2 and all positive control concentrations, further demonstrating strong SARS-CoV-2 spike protein binding (Figure 2C, red). Negative control experiments did not differ significantly across a series of dilutions, indicating weak binding to the PFDT-ACE2 modified sensor. All positive groups are likely to be significantly different from all negative groups, as indicated by the median of negative data being outside the IQR of the positive groups. Considering this evidence, it was concluded that the positive HRP-conjugated SARS-CoV-2 spike protein successfully and specifically binds to the ACE2 receptor, while HRP-conjugated streptavidin does not bind specifically. The signal produced by the negative control is most likely due to small amounts of absorption into the fluorescein SAM. The signal is lower and appears to saturate at low levels compared to the positive control, suggesting that the fluorescein SAM layer provides anti-biofouling properties, allowing the positive signal to prevail. It should also be noted that the starting concentration of the HRP-labeled streptavidin solution is in the 1 mg / mL range, meaning that a series of dilutions of the negative control protein solution are significantly more concentrated (at least an order of magnitude) than the HRP-conjugated spike protein solution. The fact that there is strong evidence of specific binding for the positive and relatively weaker binding for the negative also confirms that ACE2 is physically adsorbed into the fluorescein SAM in a significantly sufficient amount and orientation to bind to the positive ligand. If ACE2 binds in an undesirable orientation, ligand access to the receptor binding site will be hindered, and the positive signal will be significantly reduced.

[0039] It was also confirmed that binding efficiency significantly decreased when shorter 8-carbon octanethiols and longer 11-chain undecanethiols were used (Figure 6). This further indicates that the strong hydrophobic properties of the PFDT layer give rise to an adsorption mechanism that contributes to the sensor's behavior. Since HRP labeling was used on the positive protein sample, it was prudent to also use an HRP-labeled negative control to consider the potential for HRP to contribute to the binding signal. The HRP-conjugated SARS-CoV-2 spike protein is approximately 154 kDa, and the HRP-conjugated streptavidin is approximately 104 kDa. The two proteins are therefore relatively similar in size, both contain HRP labeling, and an excellent comparison between the two was possible. Equation 2:

number

[0040] HRP-conjugated SARS-CoV-2 spike protein (positive) vs. IL-6 (negative). A second negative control was examined using IL-6 protein (26kDa) at the same concentrations (1, 5, 10, 50, 100 ng / ml) as used for the positive control. IL-6 is a myokine and cytokine commonly found in the human body under normal conditions, especially after exercise. IL-6 has inflammatory and immune effects in numerous diseases, including bacterial and viral infections. IL-6 has been shown to be present at elevated levels in the "cytokine storm" observed in many advanced cases of COVID-19. Therefore, it represents a potential source of artifacts that could affect specific viral detection and was thus selected as a negative control. In this study, each control group used a single PCB array instead of dividing a single board into positive and negative sections. This increased the amount of data collected from n=4 to n=8 for both groups. The HRP-conjugated SARS-CoV-2 spike protein reaction with increasing concentration was again shown to increase sequentially (Figure 4A). This also applies to the R range from the lowest concentration of 24.4% to the highest concentration of 300%. ct This was also evident from the percentage changes (Figure 4B, red). This again indicated that the HRP-conjugated SARS-CoV-2 spike protein was binding to the PFDT-ACE2 complex. Negative control IL-6 showed similar smaller percentage increases of 57% and 59% at 10 and 50 ng / ml (Figure 4B, blue). The mean ranged from 14% at the lowest concentration to 77% at the highest concentration. Similar differences were observed between positive and negative controls at 1, 50, and 100 ng / ml concentrations (Figure 4C). Positive data (Figures 4B and 4C) showed the previously observed increasing dose-dependent behavior. Negative data increased slowly, consistent with the small mean percentage change (Figure 4B). These results suggest that the HRP-conjugated SARS-CoV-2 spike protein was successfully and specifically detected, and the negative IL-6 signal was suppressed, again indicating the anti-biofouling properties arising from fluorescein SAM. The IL-6 concentration used is 10% higher than the IL-6 levels detected in COVID-19 patients.3 ~10 5 It is important to note that the level is twice as high. Patients who progressed to acute respiratory distress syndrome (ARDS) had a median level of 7.39 pg / mL. 22 The deceased patients had a median level of 11.4 pg / mL. 23 This experiment was able to demonstrate differentiation at contamination levels far exceeding those seen in clinical COVID-19 samples. Normalized R ct Y LOD It was found to be 1.21 (Figure 4D), which is the same as reported in the previous section (Y LOD This is a slight improvement over (=2.1). Since only the 1 ng / ml concentration data point intersects this limit, it is suggested that 1 ng / ml may not be a reliable value for clear detection. However, concentration X LOD It was found to be 1.68 ng / ml (R 2 (=0.99).

[0041] Detection of inactivated SARS-CoV-2. Since it was shown that the spike protein ligand can specifically bind to the ACE2 receptor in the presence of a negative control protein, the focus was shifted to virus detection. A series of dilutions of inactivated whole virus (10 2 , 10 3 , 10 4 , 10 5 and 10 6 The dC / ml was tested against an undiluted negative control sample containing lysed cells and proteins in a "transport medium" to mimic a complex clinical sample derived from the same molecular standard kit. Incubation showed a consistently increasing R ctThis occurred (Figure 5A). The mean percentage change (n=7) for the positive control ranged from a minimum of 106% to the second highest concentration of 211% (Figure 5A, red). The highest concentration showed a decrease from 211% to 168%. This is likely due to the removal of the virus or the virus + ACE2 complex, or the subsequent reordering or absorption of SAM resulting from the presence of a large amount of virus through successive experiments. Similar effects have been observed in other studies within our group (confidential data). The negative control contained human cells as the positive control, in addition to the same background transport medium, but lacked the total virus. Actual concentrations and compositions were not provided by the supplier, but it was mentioned that the samples were substitutes for clinical human specimens, and quantified data were supplied in the form of digital copies per mL (dC / mL). The first negative sample was applied simultaneously with the first positive control and underwent the same processing. The second negative application was performed at the same time as the second positive. Due to the insufficient volume supplied relative to the required sample volume, only two negative tests were possible. Both negative reactions showed nearly identical mean percentage changes of 114.4% and 113.9% (Figure 5B, blue). This indicated that the background solution produced a high signal but quickly saturated. In contrast, the signal from the SARS-CoV-2-containing sample continued to increase with increasing virus concentration. Normalized data were obtained from 10 2 and 10 3 The dC / ml concentration is not significantly different from the negative result, but 10 4 , 10 5 and 10 6 The results showed that dC / ml appeared to be significantly different (Figure 5C). This data suggests that the virus specifically binds to the ACE2 receptor, and 10 4 It was shown that a value of dC / ml or higher can be distinguished from a negative result. The clinical level is 10. 4 ~10 11 It is in the range of RNA copies / ml. 25,26 This is within the distinguishable region presented. The performance of the sensor itself is normalized to R1.83. ct Y LOD This was shown (Figure 5D). There are no data points that intersect this limit, and the lowest 102 The concentration was successfully detected. LOD The value was 37.8 dC / ml (R 2 (=0.96064). Therefore, the sensor has the capability to detect across the entire range tested and has the potential to distinguish lower concentrations if the positive-to-negative signal ratio is improved. The results of the tests with inactivated viruses were very compelling: firstly, the viruses were heated at 65°C for 30 minutes and then gamma-irradiated, so their three-dimensional structure was significantly destroyed, and the positive and negative virus samples were present in complex media used to culture cells that produce viruses, and thus similar to other biological media such as saliva and serum. The 30-minute incubation time provided a compelling signal increase, which means the measurement is relatively rapid, especially compared to nucleic acid amplification detection, which is the criterion for judgment. Finally, there is considerable room for optimization of the assay protocol, such as shortening the virus incubation step and optimizing the washing method to maximize discriminative power.

[0042] conclusion The preparation and testing of samples for SARS-CoV-2, as well as an easily produced electrochemical biosensor, are demonstrated. The sensor consists of a base SAM composed entirely of PFDT containing ACE2 hydrophobically absorbed within the layer. Using solutions of HRP-conjugated spike protein (positive) and HRP-conjugated streptavidin and IL-6 (negative), it was possible to detect the viral spike protein with high sensitivity, specificity, and in a dose-dependent manner. Detection and differentiation of inactivated SARS-CoV-2 present in complex media (cell culture lysates) demonstrated the assurance of the sensitivity, specificity, and resistance to biological fouling required for a useful SARS-CoV-2 biosensor. The ease of sensor preparation and the compatibility of the preparation process with mass production technologies mean that the assay is potentially employable in commercially available biosensor formats. This will enable the widespread distribution of point-of-care assays for rapid population testing with laboratory diagnostics, contact tracing, and tracing, which are central to efforts to control the COVID-19 pandemic.

[0043] The presented sensor uses EIS to detect the binding of recombinant SARS-CoV-2 spike protein, as well as from solutions of positive and negative samples of inactivated SARS-CoV-2 derived from fully validated molecular standard kits. The advantages of the sensor design are that results can be obtained in a label-free manner (i.e., no fluorescent or electrochemical labeling is required during the assay process), the test is designed to measure viral particles in saliva and therefore has no possibility of detecting residual viral RNA after infection, and importantly, the sensor is designed for ease of upscaling and manufacturing with two sample production steps: (1) easy SAM formation and (2) ACE2 functionalization. In practice, the assay is validated with respect to a low-cost 8-working electrode PCB sensor system, but the assay may also be introduced to more mass-production platforms, such as screen-printed devices or glucose-format test strips. Importantly, this unlocks integration with well-established high-volume production environments, leading to diagnostics with the potential for widespread, rapid, on-site use.

[0044] <Note> Item 1 An electrochemical biosensor for use in detecting a target analyte, At least one detection including a surface coated with a self-assembled monolayer (SAM) The SAM includes electrodes and is located on a hydrofluorocarbon or fluorocarbon. The hydrofluoric acid bonded to the electrode surface via reactive sulfur or silicon groups Contains, or is essentially derived from, fluorocarbon molecules. These consist of, Electrochemical biosensor. Section 2 The hydrofluorocarbon or fluorocarbon molecule reacts to one or more reactions. Linear, branched, or cyclic alkanes, alkenes, or apheresis having sulfur or silicon groups. A lukin molecule, an electrochemical biosensor as described in item 1. Section 3 The electrochemical biosensor according to claim 1 or 2, wherein the reactive sulfur group is a thiol or a silane group. Section 4 The fluorocarbon molecule is a linear fluoroalkanethiol or fluoroalkane An electrochemical biosensor that is a silane, as described in any of items 1 to 3. Section 5 The aforementioned linear fluoroalkanethiol or fluoroalkanesilane: 1H,1H,2H,2H-perfluorodecanethiol, 3,3,4,4,5,5,6,6,7,7,8,8,8-Tridecafluoro-1-Octane thiol, 3,3,4,4,5,5,6,6,7,7,8,8,9,9,10,10,10-hepta Decafluoro-1-decanethiol, 3,3,4,4,5,5,6,6,6-nonafluoro-1-hexanethiol, 2,2,2-trifluoroethanethiol, 1H,1H,2H,2H-perfluorooctyltriethoxysilane, 1H,1H,2H,2H-perfluorodecyltriethoxysilane, and 1H,1H,2H,2H-Perfluorododecyltrichlorosilane An electrochemical biosensor as described in item 4, selected from the group consisting of the following: Section 6 The aforementioned linear fluoroalkanethiol is 1H,1H,2H,2H-perfluorodecanethiol The electrochemical biosensor described in item 4. Section 7 The electrode surface may be made of glassy carbon; metal oxide; conductive polymer; or gold, ruthenium. An electrochemical biosensor according to any one of items 1 to 6, formed from a precious metal including rhodium, palladium, platinum, and silver. Section 8 The electrochemical biosensor according to item 7, wherein the electrode surface is made of gold. Section 9 The biological agent captured by the SAM layer coated on the electrode surface Furthermore, the electrochemical biosensors described in any of items 1 to 8. Request 10 The electrochemical biosensor according to item 9, wherein the biological agent is captured in the SAM layer by physical adsorption. Section 11 A method for producing an electrochemical biosensor as described in any of items 1 to 8: Organic solvents and hydrofluorocarbons or fluorocarbons as described herein By bringing a solution containing the molecule into contact with at least one detection electrode surface, A step of forming a SAM on the surface of at least one detection electrode, The solvent is evaporated, and the SAM is formed on the surface of at least one detection electrode. The process that makes this possible Methods that include... Section 12 The electrode coated with the SAM is brought into contact with the solution containing the biological agent. The SAM layer coated on the electrode captures the biological agent. The method according to item 11, further comprising a step that enables the following. Section 13 At least one of the above detection electrodes is electrically coupled to at least one An electrochemical biosensor according to any one of items 1 to 10, further comprising a reference electrode and / or a counter electrode. Section 14 The electrochemical biosensor according to any one of items 1 to 10, wherein the electrode is provided on a substrate. Section 15 Screen-printed electrodes; microelectrodes; on printed circuit boards; or on FETs / OFETs An electrochemical biosensor as described in item 14, provided in the form of... Section 16 The biological agent captured by the sensor binds, and is typically capable of specific binding. Use of an electrochemical biosensor as described in any of sections 1-10 and 14-15 for the detection of a target analyte. Section 17 The target analyte is a chemical substance (e.g., hormone, narcotic, or contaminant) or a biological Physical molecules (e.g., peptides, proteins, glycoproteins, enzymes, glycolipids, cell surface receptors) Use as described in item 16 (body, cytokines, antibodies, nucleic acids). Section 18 The target analyte is either free within the sample to be analyzed or contained within it. The aforementioned target analytes are cells, cell membranes, and viral coats that are typically found in situ. Uses described in paragraphs 16 or 17, which remain in part, etc. Section 19 The use according to item 18, wherein the target analyte is a viral coat protein. Section 20 The biological macromolecule captured by the electrochemical biosensor is ACE-2 Yes, and the target analyte is SARS-CoV-2 or its coat protein. , use as described in item 16. Section 21 The use described in paragraph 20, wherein the SARS-CoV-2 is COVID-19. < References >

[0045] Table 1-1 Table 1-2 Table 1-3

Claims

1. An electrochemical biosensor for use in detecting a target analyte, At least one detection including a surface coated with a self-assembled monolayer (SAM) The SAM includes electrodes, and the SAM is located on a hydrofluorocarbon or fluorocarbon. The hydrofluoric acid bonded to the electrode surface via reactive sulfur or silicon groups Contains, or is essentially derived from, fluorocarbon molecules. It consists of these, The electrochemical biosensor is provided by the SAM coated on the electrode surface. This includes further biological drugs that have been captured, Electrochemical biosensor.

2. The hydrofluorocarbon or fluorocarbon molecule reacts to one or more reactions. Linear, branched, or cyclic alkanes, alkenes, or apheresis having sulfur or silicon groups. The electrochemical biosensor according to claim 1, which is a lukin molecule.

3. The electrochemical biosensor according to claim 1 or 2, wherein the reactive sulfur or silicon group is a thiol or silane group.

4. The fluorocarbon molecule is a linear fluoroalkanethiol or a linear fluoro An electrochemical biosensor according to any one of claims 1 to 3, wherein the sensor is a quilsilane.

5. The linear fluoroalkanethiol or linear fluoroalkylsilane is: 1H,1H,2H,2H-perfluorodecanethiol, 3,3,4,4,5,5,6,6,7,7,8,8,8-tridecafluoro-1-octa thiol, 3, 3, 4, 4, 5, 5, 6, 6, 7, 7, 8, 8, 9, 9, 10, 10, 10 - hepta Decafluoro-1-decanethiol, 3,3,4,4,5,5,6,6,6-nonafluoro-1-hexanethiol, 2,2,2-trifluoroethanethiol, 1H,1H,2H,2H-perfluorooctyltriethoxysilane, 1H,1H,2H,2H-perfluorodecyltriethoxysilane, and 1H,1H,2H,2H-perfluorododecyltrichlorosilane An electrochemical biosensor according to claim 4, selected from the group consisting of the following.

6. The aforementioned linear fluoroalkanethiol is 1H,1H,2H,2H-perfluorodecanethiol The electrochemical biosensor according to claim 4, which is an all-purpose device.

7. The electrode surface is formed from glassy carbon; metal oxide; conductive polymer; or noble metal. The electrochemical biosensor according to any one of claims 1 to 6.

8. Claim that the precious metal is gold, ruthenium, rhodium, palladium, platinum, or silver. The electrochemical biosensor described in 7.

9. The electrochemical biosensor according to claim 8, wherein the electrode surface is made of gold.

10. The biological agent is captured by the SAM by physical adsorption, any one of claims 1 to 9. An electrochemical biosensor as described in item 1.

11. A method for producing an electrochemical biosensor according to any one of claims 1 to 10. hand: A solution containing an organic solvent and hydrofluorocarbon or fluorocarbon molecules, By bringing the surface of at least one detection electrode into contact with the solvent and evaporating the solvent, A step of forming a SAM on the surface of at least one detection electrode, The electrode coated with the SAM is brought into contact with the solution containing the biological agent. A step of capturing the biological agent with the SAM coated on the electrode. A method that includes and

12. The hydrofluorocarbon or fluorocarbon molecule is any of claims 2 to 6. The method according to claim 11, as defined in paragraph 1.

13. At least one of the above detection electrodes is electrically coupled to at least one Any one of claims 1 to 10 further comprises a reference electrode and / or a counter electrode. The electrochemical biosensor described above.

14. The electrode is provided on a substrate, the electrochemical battery according to any one of claims 1 to 10. Io sensor.

15. They are provided in the form of screen-printed electrodes or microelectrodes, or on printed circuit boards. Claim 1, provided on a field-effect transistor or an organic field-effect transistor. The electrochemical biosensor described in 4.

16. The target analyte that can bind to the biological agent captured by the SAM In detection, the electrochemical bio in any one of claims 1 to 10 and 13 to 15 Use of sensors.

17. The target analyte specifically binds to the biological agent captured by the SAM. The use described in claim 16 is possible.

18. Claim 16 or 17, wherein the target analyte is a chemical substance or a biological molecule. Use as described above.

19. The chemical substance is a hormone, narcotic, or pollutant, and the biological molecule is a peptide Proteins, glycoproteins, enzymes, glycolipids, cell surface receptors, cytokines, antibodies, The use according to claim 18, or a nucleic acid.

20. The target analyte is either free within the sample to be analyzed, or the target Claims 16-19, wherein the GET analyte is a cell, cell membrane, or part of a viral coat. Use as described in any one of the items.

21. The use according to claim 20, wherein the target analyte is a viral coat protein.

22. The biological agent captured by the SAM is ACE-2, and the target The claim 16 states that the analyte is SARS-CoV-2 or its coat protein. Use.

23. The use according to claim 22, wherein the SARS-CoV-2 is COVID-19.

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