Patent Application
20230400455
The present disclosure relates to an electrochemical sensor for use in diagnosis tests. There is also provided methods of testing a patient sample using the electrochemical sensor as disclosed.
Electrochemical biosensors are a promising route to realising rapid and sensitive detection of a large range of pathogens and clinically important biomarkers1. The most well-known example is the glucose biosensor (most commonly an amperometric sensor), which is in widespread use for the home testing of blood glucose levels and which serves diabetic patients so well in the routine monitoring of blood sugar levels2. Numerous other biosensors have been developed which operate by a range of principles3, including, cyclic voltammetry (CV)4, linear sweep voltammetry (LSV)5, electrochemical impedance spectroscopy (EIS)6 and differential pulse voltammetry (DPV)7. EIS involves a measurement setup where a small AC excitation potential is imposed at the working electrode (often under open circuit potential) and the resulting current response of the electrochemical cell is measured. Various parameters associated with the cell and its response can be extracted from the EIS response, and these include the solution resistance (RS), the double layer capacitance (CDL), the charge transfer resistance (RCT) and the Warburg impedance (W)8. The double layer capacitance and the charge transfer resistance have been shown to be particularly effective for the label-free monitoring of binding at biologically functionalised electrode surfaces. These techniques enable the sensitive and specific measurement of DNA and protein biomarkers which has been shown repeatedly in the literature9.
For many electrochemical biosensors, like those described above, surface functionalisation and attachment chemistry play a major role in sensor design and performance10,11. For gold sensors, the attachment of biological molecules often takes place through use of gold-thiol attachment and more specifically through the formation of self-assembled monolayers (SAMs)12. SAMs serve the dual purpose of blocking the electrode surface from non-specific binding of proteins, cells and other components in a sample medium and can ensure correct orientation of the bio-recognition element (e.g. DNA sequence, antibody or enzyme)13. Self-assembled layers are often formed by incubation of gold surfaces with solutions of thiolated biomolecules and can contain single molecule (monolayer) or multi component forms where additional complexity is introduced in order to ensure adequate orientation of the receptor and good resistance to surface fouling.
The present disclosure is based on work conducted by the investigators into the development of sensors for diagnostic use, which employ fluoro organothiol or fluoro organosilane molecules in the formation of SAMs. In one embodiment SAMs are formed using 1H,1H,2H,2H-Perfluorodecanethiol (PFDT) on sensor surfaces which is then bio-functionalised with a biological agent. PFDT spontaneously forms a dense hydrophobic SAM on gold surfaces, and reduces surface biofouling14. PFDT has been previously used to enhance performance of organic transistors15 and to reversibly organise DNA onto a micro-patterned substrate16.
In a first teaching, the present disclosure provides an electrochemical biosensor for use in detecting a target analyte, the sensor comprising:
A hydrofluorocarbon is an organic compound, which contains fluorine and hydrogen atoms. A fluorocarbon is a compound in which all the C—H bonds have been replaced by C—F bonds. The hydrofluorocarbon or fluorocarbon molecule may take the form of a linear, branched or cyclic alkane, alkene or alkyne molecule having a single or multiple reactive sulphur or silicon groups.
In one embodiment, the reactive sulphur group(s) may be a thiol, as such the molecule is a fluoro organothiol. In one embodiment, the reactive silicon group(s) may be a silane, as such the molecule is a fluoro organosilane.
In one embodiment, the fluorocarbon molecule is a linear fluoro alkanethiol, or fluoro alkanesilane.
Exemplary compounds suitable for use in the present disclosure include:
In one embodiment the compound for use in forming a SAM is 1H,1H,2H,2H-Perfluorodecanethiol
The electrodes of the present disclosure may be made of any suitable electrically conductive material, which may be coated by the hydrofluorocarbon or fluorocarbon molecules as described herein. Suitable materials include glassy carbon; metal oxide; conducting polymer; and noble metals including gold, ruthenium, rhodium, palladium, platinum and silver. In one particular embodiment, the electrode is gold.
A self-assembled monolayer (SAM) is a self-organized layer of typically amphiphilic molecules in which one end of the molecule shows a specific affinity for a substrate material. SAM molecules can include a head group that anchors the molecule to the substrate, as well as a tail portion or functional group at the terminal end. SAM layers can be formed by the chemisorption of head groups, such as reactive sulphur (e.g. thiol) and silicon (e.g. silane) groups onto a substrate material from the vapor or liquid phase. The tail portion group of the SAMs according to the present disclosure make use of the known, fluorous effect17 and are hydrophobic in nature, allowing capture or physisorption of a variety of biological agents as will be described in more detail herein.
Advantageously, the hydrophobic nature of the SAM not only allows a relevant biological agent to be captured by the SAM, but also provides a barrier layer which serves to reduce fouling and/or interference from other materials, such as proteins, cells and other molecules, which may be present with a sample. A sample may be obtained from a subject and may blood, saliva or any other suitable biological fluid, such as urine, semen, or tissue sample. Alternatively, the sample may be an environmental sample, for example a water sample, soil sample, or even a plant sample.
In a further teaching, the present disclosure further provides an electrochemical biosensor as described in the first teaching and embodiments, further comprising a biological agent captured, such as by physisorption, by the SAM layer coated on the surface of the electrode.
In a further teaching, the present disclosure provides a method of making an electrochemical biosensor as described in the first or further teachings and embodiments described above, the method comprising: forming a SAM on a surface of at least one detection electrode, by contacting the surface of the at least one detection electrode with a solution comprising an organic solvent and a hydrofluorocarbon or fluorocarbon molecule as described hereinabove,
In some embodiments of the present disclosure, the electrochemical biosensors further comprise at least one reference and/or counter electrode, which may be electrically coupled to said at least one detection electrode, e.g. via a measurement system or connections for connecting to a measurement system. The electrodes are typically on a substrate and may, for example, be provided in the form of screen printed electrodes; microelectrodes; on a printed circuit board; FETS/OFETS and the like. The electrochemical biosensors may comprise, or be provided with connections for connecting to, a measurement system. The measurement system may be configured to apply an electrical signal between the at least one detection electrode and the at least one reference and/or counter electrode. The measurement system may be configured to measure an electrical response resulting from the detection of an analyte by the biological agent. The measurement system may be configured to perform impedance, voltammetric, or amperometric measurements. The measurement system may be configured to perform electrochemical impedance spectroscopy (EIS).
An electrochemical biosensor in accordance with the present disclosure may be used detect a target analyte which is capable of binding, typically specifically binding, to the biological agent captured by the sensor. The target analyte may be a chemical (such as a hormone, narcotic or pollutant) or biological molecule, such as a peptide, protein, glycoprotein, enzyme, glycolipid, cell surface receptor, cytokine, antibody, nucleic acid or the like. The target analyte may be free within the sample being analyzed, or may still be part of a cell, cell membrane, virus coat etc in which the target analyte, such as biological molecule, is normally found in situ. Binding between the biological agent and the target analyte involves the attractive binding (i.e. non-repelling) of two or more species held together by attractive forces. Such binding comprises an interaction between the biological agent and the target analyte that pulls (or draws) the biological agent and target analyte in a sample together. Binding includes, but is not limited to covalent interactions; electrostatic interactions; ion-ion interactions, for example between attractive or opposite charges; dipole interactions; ion-dipole interactions; hydrogen bonding interactions; van der Waals interactions; pi-stacking interactions; the sharing of electron density or combinations thereof.
The electrochemical sensors of the present disclosure provide a signal transduction method, whereby any binding of the target analyte to the biological agent may be converted into a signal for processing and/or display. The biological agents may be similar to the target analytes and may include proteins, enzymes, antibodies, nucleic acids, etc. The biosensors can be configured to use specific chemical interaction properties (such as an enzyme and its substrate) or molecular recognition mechanisms (such as a protein binding to a receptor or antibody binding to an antigen) to identify target analytes. Biosensors can use the electrode to transform an electrical signal resulting from the detection of an analyte by the biological agent into a different signal that can be addressed by optical, electronic or other means.
In one embodiment, impedance, voltammetric, or amperometric measurements may be performed in order to detect the signal, or a change in signal, by the electrochemical sensor. In one embodiment, the electrochemical detection of a target analyte is carried out by electrochemical impedance spectroscopy (EIS). In this manner, the electrochemical biosensor of the present disclosure transduces changes in the interfacial properties between the electrode and an electrolyte induced by the biological macromolecule binding to the target analyte to an electrical signal. Typically, a redox pair, such as K3[Fe(CN)6]/K4[Fe(CN)6] may be used as a redox indicator for the electrode kinetics at the interface. Sensors based on the use of EIS detection are label-free and, thus, possess advantages of low cost, simplicity and ease of miniaturization. EIS is particularly useful because it is sensitive to surface interactions and quantifies the interfacial charge transfer resistance (RCT) that is associated with charged redox probes. RCT is strongly affected by changes in charge distributions near the electrode-solution interface in a sample solution, which results in surface sensitivity.
The biosensors of the present disclosure may find application in many different diagnostic applications, including detection of infectious agents, such as bacteria, viruses (including as influenza, SARS-COV2) and the like; molecules such as IL-16 and procalcitonin associated with sepsis; cardiac biomarkers such as troponin; and liver biomarkers, for example. Advantageously, the principles of detection using the sensors of the present invention have broad applicability due to the relative ease of manufacture and speed of detection.
In one non-limiting embodiment, the present disclosure is directed to the detection of COVID-19. In this example, the biological macromolecule, which is captured by the electrochemical biosensor is ACE-2, which has been identified as an entry receptor for SARS-CoV-2, the virus responsible for COVID-19. Thus, in an embodiment, an electrochemical biosensor in accordance with the present disclosure, wherein the biological macromolecule is ACE-2, may used in the detection of SARS-CoV-2, such as COVID 19. Alternatively, a COVID 19 specific antibody18, such as an antibody, which is specifically able to bind the spike protein of COVID 19 may be used as the biological macromolecule. In this manner, the target analyte to be detected is the virus or a coat protein of SARS-CoV-2 or more specifically COVID 19.
The chemical terminologies as used herein have their standard meanings known in the art, in accordance with the IUPAC Goldbook, unless explicitly stated. Unless the context clearly requires otherwise, throughout the description and the claims, the words “comprise”, “comprising” and the like, are to be construed in an inclusive sense as opposed to an exclusive or exhaustive sense, that is to say, in the sense of “including, but not limited to”.
The present disclosure will now be further described by way of example and with reference to the following figures, which show:
Abbreviations 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-assembled monolayer; Rd, Charge transfer resistance; OCP, open circuit potential; IQR, inter quartile range.
Chemicals. 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). Deionised water (5.00 μS/cm @ 25° C.) was purchased from Scientific Laboratory Supplies Limited (Nottingham, UK). Inactivated SARS-CoV-2 and negative control obtained from Randox laboratories Ltd (Crumlin, UK). ACE2 was purchased from Abcam (Cambridge, UK), HRP conjugated spike protein was purchased from The Native Antigen Company (Oxford, UK) and HRP conjugated streptavidin was purchased as part of an IL-6 diagnostics kit from Bio-techne (Abingdon, UK).
Preconditioning. SEPI BIOTIP multichannel electrode PCB platform (biotip ltd, Bath, UK) were cleaned according to the supplied protocol. This consisted of a 15-minute submersion in a solution of 50 mM KOH in H2O2 30% (v/v) at room temperature. The PCB was then rinsed with DI water and dried using compressed air. The PCB was then electrochemically cleaned by submerging in 50 mM KOH (DI water as solvent) with an external platinum counter electrode (Metrohm, Runcorn, UK) and 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 was −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. All electrochemical measurements were performed using a PalmSens4 potentiostat and the accompanying PSTrace software, both supplied by Palmsens BV (Houten, Netherlands).
Fluorous SAM and ACE2 immobilisation. The SAM solution was prepared by magnetically stirring toluene and adding PFDT until a 1 mM solution was formed. Stirring aids in dispersing the PFDT throughout the solution. Fluorocarbons can have low miscibility in organic solvents and have a propensity for self-interaction forming separate phases via the fluorous effect17. The PCBs were orientated horizontally in a small glass petri dish and the PFDT solution added to cover the PCB with excess solution. Toluene evaporates quickly, therefore having excess solution and a film covering reduced evaporative losses. The PCBs were incubated overnight at room temperature, then rinsed with DI water (10 second water bottle flow per electrode) and dried with compressed air. All work with toluene was performed in a suitable fume hood with proper halogenated solvent waste disposal routes.
ACE2 was diluted from stock in 1×PBS to 1 μg/ml and 10 μl aliquots were applied to each working electrode on the PCB and left to incubate for 1 hour at room temperature. Following incubation, the PCBs were rinsed with 1×PBS (10 second water bottle flow per electrode) and dried with compressed air.
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 the positive control HRP conjugated SARS-CoV-2 spike protein and negative controls of similar sized proteins (HRP conjugated streptavidin and IL-6) were incubated at room temperature for 30 minutes on the PCB sensor arrays with rinsing with 1×PBS (10 seconds water bottle flow per electrode) and EIS measurements between each concentration incubation. HRP conjugated SARS-CoV-2 spike protein and IL-6 concentrations used were 1, 10, 50 and 100 ng/ml (all dilutions in 1×PBS). HRP conjugated streptavidin was obtained as part of an ELISA kit and the concentration was not disclosed. The accompanying instructions recommended a 1:40 dilution for ELISA assays. The series of dilutions used (1:100, 1:75, 1:50, 1:25 and 1:5) were distributed about the 1:40 recommended dilution.
Inactivated virus detection. For detection of inactivated virus a clinical molecular standards kit for SARS-CoV-2 (Qnostics) was purchased. The kit contained positive and negative samples of the virus present in a complex “transport medium” representative of a clinical sample. A series of dilutions of the positive control (inactivated virus+transport medium and human cells) was incubated for 30 mins at room temperature on the PCBs. The concentrations used were 102, 103, 104, 105 and 106 dC/ml (digital copies per ml). Due to small volume of solutions provided, the negative control (transport medium+human cells) was incubated twice for 30 minutes at room temperature. Room temperature incubations were chosen to replicate the operational environmental conditions likely required for a diagnostic device. The PCBs were rinsed with 1×PBS (10 seconds wash bottle flow per electrode) and EIS measurements preformed between each incubation.
EIS parameters. All EIS measurements used the following parameters. Eac=0.01 V rms, Edc=0 V, frequency range=100 kHz to 1 Hz with 50 frequencies at 9.8/decade and measurements were made versus the open circuit potential (OCP). All measurements were obtained using 5 mM K3[Fe(CN)6]/K4[Fe(CN)6] in 1×PBS.
Results and Discussion
Fluorocarbon SAM functionalisation. Commonly electrochemical biosensors will have their probe molecule directly attached to the sensor surface (via covalent bonding, physisorption and chemisorption) and surrounded by a hydrocarbon-based SAM. Less commonly the hydrocarbon SAM is immobilized first, and the biomolecule adsorbed into it via hydrophobic physisorption interactions, which can effect better orientation of the probe biomolecule, increasing the likelihood of receptor-target binding. Such an approach does however have the disadvantage in also being a weaker immobilization method than covalent attachment and therefore a higher probability of removal during incubations and wash steps. The investigators sought to consider the use of fluorocarbons as they can offer greatly increased amphiphobicity (hydrophobic and lipophobic character) over hydrocarbons offering stronger physisorption and anti-biofouling properties17. The ability of fluorocarbons to form a SAM on the PCB electrode surfaces was investigated. An overnight incubation of 1 mM PFDT affected an increase in the measured impedance of the electrodes, which is evident as a larger R d semi-circle (SAM) compared to the clean impedance in the Nyquist plot (
Percentage change was calculated using Equation 1, where A % is percentage change, Rct-Before is the Rct of the initial stage, and Rct-After is the Rct of the incubation stage. Significant differences were gauged from the box plots. If the median of one group lies outside the inter quartile range (IQR) of another it is likely there is a significant difference between the groups. T-test analysis was not performed as these experiments were not designed with hypothesis testing in mind and as such may report false results. The box plot (
ACE2 hydrophobic immobilisation. A further benefit of the strongly hydrophobic fluorous SAM is that it provides an ideal environment to facilitate hydrophobic physisorption of ACE2 biomolecules. To test this ACE2 protein was incubated in the presences of the electrode SAM. After 1 hour of incubation with 1 μg/ml of ACE2 solution on the SAM functionalized electrodes, a small impedance increase was apparent (
HRP conjugated SARS-CoV-2 spike protein (positive) and HRP conjugated streptavidin protein (negative). Having confirmed successful immobilisation of ACE2, HRP conjugated SARS-CoV-2 spike protein (HRP conjugated version was used to enable visual determination of binding) was incubated with the functionalised sensor surface for 30 minutes. The measured impedance for 1, 10, 50 and 100 ng/ml consistently increased compared to the preceding concentration demonstrating dose dependant behaviour (
It was also observed (
P
LOD
=Y
i+3SDi (Eq. 2)
where YLOD was the limit of detection of the Y-axis parameter (normalized Rct), Yi was the y-intercept value obtained from linear regression of the data and SDi was the accompanying standard deviation of the y-intercept. The value obtained obtained for the normalised Rct YLOD for SARS-CoV-2 HRP conjugated spike protein was 2.1 (
HRP conjugated SARS-CoV-2 spike protein (positive) versus IL-6 (negative). A second negative control was investigated using the protein IL-6 (26 kDa) with equal concentrations as used for the positive control (1, 5, 10, 50, 100 ng/ml). IL-6 is a myokine and cytokine common in the human body under normal circumstances, especially after exercise. It has inflammatory and immune effects in a multitude of diseases including bacterial and viral infection. IL-6 has been shown to be present at elevated levels in the ‘cytokine storm’ which is observed in many advanced cases of COVID-19. This would therefore represent a potential source of artefact that could affect specific virus detection and was thus chosen as a negative control. This time, each control group used a single PCB array instead of portioning a single board into positive and negative sections. This increased the amount of collected data for both groups from n=4 to n=8. The HRP conjugated SARS-CoV-2 spike protein response for increasing concentration was once again seen to sequentially increase (
Inactivated SARS-CoV-2 detection. Having shown that the spike protein ligand was able to specifically bind to the ACE2 receptor in the presence of negative control proteins, the focus changed to virus detection. A dilution of series of inactivated whole virus (102, 102, 104, 105 and 106 dC/ml) was tested against undiluted negative control samples from the same molecular standards kit and containing lysed cells and proteins in the “transport medium” in order to mimic a complex clinical sample. Incubations resulted in a consistently increasing Rct (
The preparation and testing of a simple and easily produced electrochemical biosensor for SARS-CoV-2 has been demonstrated. The sensor consists of a base SAM composed entirely of PFDT with ACE2 hydrophobically absorbed into the layer. It was possible, using solutions of HRP-conjugated spike protein (positive) and HRP conjugated streptavidin and IL-6 (negatives) to detect the viral spike protein in a sensitive, specific and dose dependant manner. Detection and discrimination of inactivated SARS-CoV-2 virus present in a complex medium (cell culture lysate) was demonstrated to confirm the sensitivity, specificity and resistance to biological fouling necessary for a useful biosensor for SARS-CoV-2. The ease with which the sensor can be prepared and the compatibility of the preparation steps with mass manufacturing techniques mean the assay is potentially adoptable on existing commercial biosensor formats. This would allow for wide distribution of point of care assays for rapid testing of the population with diagnostics being at the centre of test, track and tracing of contacts, central to efforts to control the COVID-19 pandemic.
The presented sensor uses EIS to detect binding from solutions of recombinant SARS-CoV-2 spike protein and positive and negative samples of inactivated SARS-CoV-2 from a fully validated molecular standards kit. Advantages of the sensor design are that the result can be produced in a label free manner (i.e. there is no need to add a fluorescent or electrochemical label during the assay steps), the test is designed to measure viral particles in saliva so there is no chance of detecting residual viral RNA post infection and crucially the sensor has been designed for ease of upscaling and manufacture with two simple production steps: (1) facile SAM formation and (2) ACE2 functionalisation. In the work, the assay is demonstrated on a low cost eight working electrode PCB sensor system, however, the assay can be transferred onto even more mass manufacturable platforms such as screen-printed devices or glucose format test strips. Importantly, this would unlock integration with a well-established high volume production environment and lead to a diagnostic with the potential for widespread, rapid, point of need use.
| Number | Date | Country | Kind |
|---|---|---|---|
| 2017047.8 | Oct 2020 | GB | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/GB2021/052781 | 10/26/2021 | WO |
