SENSOR FOR DETECTING BIOANALYTES AND METHOD THEREOF - Patent application

Abstract

The present invention provides a sensor for detecting a bioanalyte, the sensor comprising: a substrate; a pair of terminal electrodes disposed on the substrate in a spaced apart opposing relationship; and a non-insulating sensing element applied to a surface of the substrate between and in electrical contact with the pair of terminal electrodes, the sensing element providing a conductive path between the terminal electrodes, the sensing element comprising an oxygen-deficient metal oxide layer and a bioanalyte binding site, the sensing element generating an electrical signal proportional to a change in conductance of the sensing element corresponding to binding of the bioanalyte to the bioanalyte binding site when a voltage is applied across the sensor.

Description

[Technical field]

[0001] The present invention relates to sensors, and more particularly to non-invasive sensors and methods for detecting bioanalytes in body fluids.

[0002] The present invention has been developed primarily for use in the detection of various bioanalytes in body fluids and will hereinafter be described with reference to this application.

[0003] The following discussion of the background of the invention is intended to facilitate an understanding of the present invention, but it should be understood that this discussion is not an admission that any of the material mentioned was published, known or part of the general knowledge in Australia or any other country at the priority date of any of the claims herein. [Background technology]

[0004] In tissues and body fluids Targeted Biomarkers (Hereinafter, bioanalytes (bioanalytes) ) The level of There are two approaches to monitoring and measuring. The first approach is to use invasive sensors, where the sensor components come into direct contact with tissues and body fluids, potentially causing infection, tissue damage, and discomfort. On the other hand, non-invasive sensors measure the concentration of bioanalytes in body fluids using a variety of techniques, including optical absorption, electrochemistry, transduction, and conductometry. These sensors can be used to measure the concentration of bioanalytes in body fluids.

[0005] In terms of non-invasive sensors, those based on optical absorption are not particularly precise due to the close overlap of weak absorption bands of the various bioanalytes that may be present in bodily fluids and the temperature sensitivity of their measurements.

[0006] Electrochemical sensors, on the other hand, are more precise and therefore currently dominate the field of biosensing. Electrochemical sensors work by measuring an electrical signal generated by a sensing element attached to the sensor in reaction with the bioanalyte of interest, where the generated electrical signal is proportional to the concentration of the bioanalyte. The electrochemical reaction of the sensor is measured by a measurable current (amperometric), a measurable charge accumulation or potential (potentiometric), the conductive properties of the medium (conductometric), and an impedance measurement that is a combination of resistance and reactance.

[0007] Sensors that utilize electrochemical transduction typically require a working electrode, a counter (or auxiliary) electrode, and a reference electrode. The reference electrode is maintained at a distance from the site of interaction of the biological recognition element and the analyte to establish a known and stable potential. The working electrode acts as the transduction component when the interaction occurs, and the counter electrode measures the current and facilitates the supply of an electrolyte that allows current transfer to the working electrode.

[0008] Conductivity sensors also use electrodes to measure the ability to pass an electric current between media. However, they do not require a reference electrode. They also operate with low amplitude AC voltages, which avoids the Faraday effect of the electrodes, and can be easily miniaturized and integrated using thin film technology.

[0009] Although electrical conductivity sensors have certain advantages, the use of polymers as the sensing element often impairs sensitivity and reduces the durability and long-term stability of the sensors.

[0010] As an alternative to sensors that directly conduct electricity, sensors using field effect transistors have also been developed. A field effect transistor is a device with three terminals: source, gate, and drain. The device operates when a change in the gate generates a field effect, changing the conductivity between the source and drain.

[0011] For example, US Patent Application Publication No. 2010 / 2016256 describes a biosensor including a substrate, a source electrode on the substrate, a drain electrode on the substrate, and at least one functionalized nanobelt on the surface of the substrate between the source and drain electrodes, the functionalized nanobelt having a chemically functionalized surface with one or more detection molecules attached thereto for binding a bioanalyte to be detected such that a field gate effect occurs upon binding of the analyte to the one or more detection molecules attached to the nanobelt surface. The device functions such that binding of the molecules changes the field effect of the nanobelt (gate), changing the conductivity of the path between the source and drain, and the change in conductivity can be monitored. In general, this type of device suffers from two drawbacks:

[0012] First, field-effect transistors are generally devices that repeatedly switch on and off, and have a nonlinear response. Because the resistance value does not change linearly, there is little linear region and then a plateau, making it difficult to use under a wide range of conditions.

[0013] Second, as will be appreciated by those skilled in the art, for such devices to operate as described, it is necessary that an insulating (dielectric) layer be present between the source-drain conductive path and the source of the gate bias (in this case the nanobelt). Therefore, this type of device has the disadvantage that it is relatively complex to fabricate due to the various structural elements, and is more difficult to produce on an industrial scale than structurally simpler sensors. Summary of the Invention [Problem to be solved by the invention]

[0014] The present invention provides a sensor for use in the detection of a bioanalyte and a method for the detection thereof that overcomes or substantially ameliorates or at least provides an alternative to at least some of the shortcomings of the prior art. [Means for solving the problem]

[0015] According to a first aspect of the present invention, there is provided a sensor for detecting a bioanalyte comprising a substrate, a pair of terminal electrodes disposed on the substrate in spaced apart opposing relationship, and a non-insulating sensing element applied to a surface of the substrate between and in electrical contact with the pair of terminal electrodes, the sensing element providing a conductive path between the terminal electrodes, the sensing element having an oxygen deficient metal oxide layer and a bioanalyte binding site, wherein application of a voltage across the sensor produces an electrical signal proportional to a change in conductance of the sensing element corresponding to binding of the bioanalyte to the bioanalyte binding site.

[0016] Preferably, the oxygen-deficient metal oxide layer is formed from a metal oxide selected from the group consisting of zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium dioxide.

[0017] In one embodiment, the oxygen deficient metal oxide layer has a thickness in the range of about 50 nm to about 200 μm.

[0018] Preferably, the oxygen-deficient metal oxide layer is applied to the substrate surface by a technique selected from the group consisting of reactive sputtering, physical vapor deposition (PVD), chemical vapor deposition (CVD), metalorganic chemical vapor deposition (MOCVD), pulsed laser deposition (PLD) and molecular beam epitaxy (MBE).

[0019] Preferably, the bioanalyte binding sites are immobilized on the oxygen-deficient metal oxide layer via an intermediate layer that is physically or chemically adsorbed to the oxygen-deficient metal oxide layer.

[0020] In one embodiment, the intermediate layer is prepared by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functionality selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxyl groups, and hydroxyl groups.

[0021] In one embodiment, the silanizing agent is selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyltrimethoxysilane (AEAPTS).

[0022] In one embodiment, the oxygen-deficient metal oxide layer has a conductivity of about 0.08 Siemens / m 2 From about 0.6 siemens / m 2 The conductance falls in the range

[0023] Preferably, the bioanalyte binding site is a biomolecule.

[0024] Suitably, the biomolecule is a protein, a peptide, a lipopeptide, a protein-binding carbohydrate or a protein-binding ligand.

[0025] In one embodiment, the biomolecule is a capture protein.

[0026] Suitably the capture protein is a protein binding scaffold, a T cell receptor, a binding fragment of a TCR, a variable lymphocyte receptor, an antibody and / or a binding fragment of an antibody.

[0027] Preferably, the protein binding scaffold is selected from the group consisting of adnectins, affilins, affibodies, affima molecules, affitins, alpha bodies, aptamers, anticalins, armadillo repeat protein based scaffolds, atrimers, avimers, designed ankyrin repeat proteins (DARPins), finomers, inhibitor cystine knot (ICK) scaffolds, Kunitz domain peptides, monobodies and / or nanophytins.

[0028] Preferably, said binding fragment of an antibody comprises a Fab, (Fab')2, Fab', single chain variable fragment (scFv), di- and tri-scFvs, single domain antibody (sdAb), diabody or a fusion protein comprising the binding domain of an antibody.

[0029] In one embodiment, the bioanalyte binding site binds interleukin-6 (IL-6).

[0030] In one embodiment, the bioanalyte binding site binds C-reactive protein (CRP).

[0031] Preferably, the substrate is made from a material selected from the group consisting of a silicon wafer, a polymer, a glass and a ceramic.

[0032] Preferably, the polymer is selected from the group consisting of polydimethylsiloxane (PDMS), polyimide (PI) and polyethylene naphthalate (PEN).

[0033] Preferably, the ceramic is selected from the group consisting of aluminum oxide (Al2O3), sapphire, and silicon nitride (Si3N4).

[0034] According to a second aspect of the invention there is provided a method of detecting a bioanalyte comprising the steps of contacting a sensing element of a sensor according to the first aspect with a sample solution containing the bioanalyte, applying a voltage across the sensor and detecting an electrical signal generated proportional to a change in conductance corresponding to detection of the bioanalyte upon binding of the bioanalyte to the bioanalyte binding site.

[0035] Preferably, the bioanalyte binding site is a biomolecule.

[0036] In one embodiment, the bioanalyte binding site binds interleukin-6 (IL-6).

[0037] Preferably, the change in conductance detected in a sample solution containing 4 femtomolar IL-6 is about 9.2%.

[0038] In one embodiment, the bioanalyte binding site binds C-reactive protein (CRP).

[0039] Preferably, the change in conductance detected in a sample solution containing a CRP concentration of 13 femtomolar is about 12.5%.

[0040] According to a third aspect of the present invention, there is provided a method of manufacturing a sensor for detecting a bioanalyte, the method comprising the steps of providing a substrate, depositing a pair of terminal electrodes on the substrate in spaced apart opposing relationship to each other, Biological Analytes applying a non-insulating sensing element in the form of an oxygen deficient metal oxide layer coated with binding sites between and in electrical contact with a pair of terminal electrodes, where the sensing element provides a conductive path between the terminal electrodes and the bioanalyte binding sites are selective for detection of the bioanalyte upon binding to the bioanalyte binding sites.

[0041] Preferably, the oxygen-deficient metal oxide layer is formed from a metal oxide selected from the group consisting of zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium dioxide.

[0042] In one embodiment, the oxygen deficient metal oxide layer has a thickness in the range of about 50 nm to about 200 μm.

[0043] Preferably, the oxygen-deficient metal oxide layer is applied to the substrate surface by a technique selected from the group consisting of reactive sputtering, physical vapor deposition (PVD), chemical vapor deposition (CVD), metalorganic chemical vapor deposition (MOCVD), pulsed laser deposition (PLD) and molecular beam epitaxy (MBE).

[0044] Preferably, the method further comprises the step of physically or chemically adsorbing an intermediate layer to the oxygen-deficient metal oxide layer in order to immobilize the bioanalyte binding sites on the oxygen-deficient metal oxide layer.

[0045] Preferably, the intermediate layer is prepared by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functional group selected from the group consisting of an epoxy group, a thiol group, an amino group, a carboxyl group and a hydroxyl group.

[0046] In one embodiment, the silanizing agent is selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-2(2-aminoethyl)-3-aminopropyltrimethoxysilane (AEAPTS).

[0047] Other aspects of the invention are also disclosed.

[0048] Notwithstanding other forms which may fall within the scope of the invention, preferred embodiments of the invention will now be described, by way of example only, with reference to the accompanying drawings in which: [Brief description of the drawings]

[0049] [Figure 1] FIG. 1 is a schematic diagram for the fabrication of a non-invasive conductometric sensor for detecting a bioanalyte in accordance with a preferred embodiment of the present invention, the sensor having a sensing element including an oxygen-deficient metal oxide thin film layer having a plurality of bioanalyte binding sites bound thereto. [Figure 2A] 2A is a plot reflecting the change in resistance (%) as a function of concentration (M) on an IL-6 antibody-immobilized conductometric sensor. The dashed line in the plot indicates the antigen concentration (M) of a healthy human body fluid (i.e., IL-6 in sweat). [Figure 2B]Figure 2B is a plot reflecting the change in resistance (%) as a function of concentration (M) on a CRP antibody-immobilized conductometric sensor. The dashed line in the plot indicates the antigen concentration (M) of a healthy human body fluid (i.e., CRP in saliva). [Figure 3A] 3A is a plot reflecting the resistance change (%) of CRP on an anti-IL-6 antibody-immobilized conductivity sensor for a cross-selection test. The nominal concentration (M) of IL-6 is 4 pM. [Figure 3B] 3B is a plot reflecting the resistance change (%) of IL-6 on an anti-CRP antibody-immobilized conductivity sensor for a cross-selection test. The nominal concentration (M) of CRP is 13 pM. [Figure 4A] FIG. 4A shows (a,c) Zn2p core level XPS spectra collected from a sputtered ZnOx thin film formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 4B] FIG. 4B is an XPS spectrum of the (b, d) O1s core level collected from the sputtered ZnOx thin film formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 4C] FIG. 4C shows (a,c) XPS spectra of the Zn2p core level collected from the sputtered ZnOy thin film formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 4D] Figure 4D shows (b, d) O1s core-level XPS spectra collected from the sputtered ZnOy thin film formed on the substrate surface of the electrical conductivity sensor in Figure 1. Ov represents an oxygen vacancy. [Figure 5A] FIG. 5A shows (a, d) resolved core-level XPS spectra of Sr element in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 5B] FIG. 5B is a (b, e) resolved core-level XPS spectrum of Ti element in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 5C]FIG. 5C shows the (c,f)O element resolved core level XPS spectra in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 5D] FIG. 5D shows (a, d) resolved core-level XPS spectra of Sr element in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 5E] FIG. 5E shows (b,e) resolved core-level XPS spectra of Ti element in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 5F] FIG. 5F shows the (c,f)O element resolved core level XPS spectra in sputtered STO thin films having different chemical compositions formed on the substrate surface of the electrical conductivity sensor of FIG. [Figure 6A] FIG. 6A is a plot reflecting the change in resistance (%) of selectivity for IL-6 antigen on an IL-6 antibody immobilized device in the presence of other test antigens. [Figure 6B] FIG. 6B is a plot reflecting the change in resistance (%) of selectivity for the CRP antigen on the CRP antibody immobilized device in the presence of other test antigens. [Figure 7A] FIG. 7A is a plot reflecting (a) the change in resistance of new (day 0) and old (day 450) devices as a function of IL-6 concentration. [Figure 7B] FIG. 7B is a plot reflecting the change in resistance of new (0 days old) and old (450 days old) devices as a function of CRP concentration. [Figure 8A] FIG. 8A is (a) a plot reflecting the change in resistance as a function of IL-6 concentration in PBS and artificial saliva. [Figure 8B] FIG. 8B is a plot reflecting a selectivity study of IL-6 on an IL-6 antibody immobilized device in the presence of other antigens used in this study. [Figure 8C]FIG. 8C is a plot reflecting the change in resistance as a function of CRP concentration in PBS and artificial saliva. [Figure 8D] FIG. 8D is a plot reflecting a selectivity study for CRP on the CRP antibody immobilized device in the presence of other antigens used in this study. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0050] In the following description, it should be noted that like or identical reference numbers in different embodiments indicate the same or similar features.

[0051] The present invention is based on the discovery of an inexpensive, non-invasive sensor for detecting the levels of various bioanalytes in bodily fluids such as human saliva and sweat to aid in the prediction and diagnosis of disease conditions. As described in detail below, the sensor has a simple and relatively easy to fabricate device structure, providing a cost-effective alternative to conventional non-invasive sensors that require specialized substrates or employ sensing technologies that provide limited accuracy in the results.

[0052] The inventors believe that the electrical conductance sensor detailed below, being compatible with CMOS circuitry, can be easily integrated with flexible / wearable electronics to provide a portable, personal and reusable sensor that can be used to continuously monitor the levels of target bioanalytes via bodily fluids without the need for invasive procedures. These bioanalytes can act as biomarkers indicative of an individual's condition or health.

[0053] The following provides a detailed description of a non-invasive conductivity sensor and its application for detecting a range of levels of bioanalytes (eg, biomarkers) in body fluids.

[0054] <Sensor> Next, a sensor for detecting a biological analyte according to a preferred embodiment of the present invention will be described.

[0055] In its simplest form, as shown in the schematic diagram of FIG. 1, the sensor comprises a substrate, a pair of terminal electrodes disposed on the substrate in a spaced apart opposing relationship, and a non-insulating sensing element applied to the surface of the substrate between and in electrical contact with the pair of terminal electrodes, the sensing element providing a conductive path between the terminal electrodes, the sensing element being comprised of an oxygen-deficient metal oxide layer surface-modified with suitable surface modifiers and synthetic conjugates or biomolecules to form sites to which a target biomarker or bioanalyte can selectively bind for detection purposes.

[0056] Each element constituting the non-invasive electrical conductivity sensor will be described below.

[0057] <Substrate> The substrate may be made from a material selected from the group consisting of a silicon wafer, a polymer, a glass or a ceramic.

[0058] For example, a polymer suitable for use as a substrate may be selected from the group consisting of polydimethylsiloxane (PDMS), polyimide (PI), and polyethylene naphthalate (PEN), while a suitable ceramic may be selected from the group consisting of aluminum oxide (Al2O3), sapphire, and silicon nitride (Si3N4).

[0059] Here, to explain the fabrication procedure of a non-invasive electrical conductivity sensor, the substrate is a hard silicon wafer with a SiO2 surface, as shown in step (1) of FIG. 1.

[0060] However, those skilled in the art will appreciate that if the desired objective is to provide a non-invasive electrical conductivity sensor that can be used as a device in applications where portability and flexibility are desired, the substrate used would ideally be a flexible polymer such as polyimide foil rather than the rigid SiO2 / Si wafer described above. The fabrication procedure for a flexible non-invasive electrical conductivity sensor using polyimide foil is similar to that described above (see FIG. 1).

[0061] <Detection element> In its simplest form, the sensing element consists of an oxygen-deficient metal oxide layer and one or more bioanalyte binding sites bound to the surface of the oxygen-deficient metal oxide layer by chemical or physical adsorption.

[0062] The metal oxide layer may be formed using any suitable metal oxide selected from the group consisting of zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium oxide.

[0063] In a preferred embodiment, the metal oxide layer is an oxygen-deficient metal oxide layer formed using zinc oxide (ZnO) or strontium titanium oxide (STO). As described below, the inventors have found that good results can be obtained when the metal oxide layer is a thin oxygen-deficient zinc oxide (ZnO) layer.

[0064] The oxygen-deficient metal oxide layer can be applied to the substrate surface by a technique selected from the group consisting of reactive sputtering, physical vapor deposition (PVD), chemical vapor deposition (CVD), metalorganic chemical vapor deposition (MOCVD), pulsed laser deposition (PLD), and molecular beam epitaxy (MBE).

[0065] In a preferred embodiment, as shown in step (2) of FIG. 1, an oxygen-deficient metal oxide layer is applied to the surface of a rigid (SiO2 / Si) wafer or flexible polyimide foil by reactive sputtering to result in a thin metal oxide film having a thickness in the range of about 50 nm to about 200 μm.

[0066] For example, as shown in step (2) of FIG. 1, zinc oxide is sputtered onto the surface of a hard (SiO2 / Si) wafer to form an oxygen-deficient zinc oxide layer (ZnO 1-x The oxygen-deficient ZnO layer thus deposited may be of any suitable thickness to suit a desired application.

[0067] Good results were obtained when the oxygen-deficient ZnO layer was in the range of about 10 nm to about 1 μm.

[0068] <Metal oxide thin film> ZnO x and ZnO y Binary metal oxide thin films such as those based on ZnO and composite metal oxide thin films such as SrTiO3 have been applied as non-invasive electrical conductivity sensors to detect bioanalytes such as IL-6 and CRP.

[0069] Below, we explain the synthesis process and chemical composition of these binary and composite metal oxide thin films.

[0070] <Zinc oxide (ZnO)> Two types of ZnO thin films with different oxygen contents were prepared by magnetron sputtering, which results in sputtered thin films with different stoichiometries. The sputtering parameters and associated electrical conductivity are shown in Table 1.

[0071] [Table 1]

[0072] The sputtering parameters are selected to design a thin film with electrical conductivity in the range of 0.08-0.6 S / m, which is where the sensitivity of the sensor is maximized.

[0073] The stoichiometry of the sputtered ZnO thin films is evaluated by X-ray photoelectron spectroscopy (XPS). x Thin films and ZnO y Core-level Zn2p and O1s spectra collected from a thin film.

[0074] As shown in Figure 4, the O1s spectrum is fitted with three distinct peaks associated with the Zn-O bond (represented by peak (1)), the oxygen vacancy (Ov, represented by (2)), and the -OH bond (represented by (3)) (see refs. 1, 2).

[0075] The fitting parameters are shown in Table 2. Comparing the peaks (2) of the two thin films, ZnO x ZnO y It has been suggested that the oxygen deficiency is relatively greater than that in the control group.

[0076] [Table 2]

[0077] <Strontium Titanium Oxide (STO)> Two types of strontium titanium oxide (SrTiO3:STO) thin films with different oxygen contents were prepared by magnetron sputtering. The sputtering parameters are shown in Table 3.

[0078] [Table 3]

[0079] Figure 5 shows the results of sputtering STO in a reducing environment (0% oxygen) and an oxidizing environment (5% oxygen). x and STOs y The XPS core-level binding energy spectra of all three elements in both types of STO thin films are shown. For both STO oxides, the Sr3d core-level spectra (Figures 5A and 5D) were fitted with a single component, and no significant shift in chemical state was observed. In addition, for both oxides, the Sr3d 5 / 2 Binding energy of -132.9eV (±0.1eV) and Sr3d 3 / 2 The binding energy of -134.7 eV (±0.1 eV) is that of Sr in STO. 2+ species (see references 3 and 4 ).

[0080] The deconvolution spectrum analysis of Ti2p core-level binding energies is shown in Figures 5B and 5E. x and STOs y The Ti2p spectra of both 4+and Ti (4-x)+ The fitting is done by Ti 4+ The presence of the element alone corresponds to a perfectly stoichiometric STO oxide, whereas the presence of Ti (4-x)+ Seed (Ti 3+ Or Ti 2+ ) indicates the presence of oxygen vacancies in the oxide system. 3+ and Ti 2+ To avoid ambiguity, only one component is fitted at low binding energies, and Ti (4-x)+ It is written as STO. x and STOs y In oxides, Ti2p 3 / 2 The peak of the bond energy of Ti is 458.4 eV. 4+ The peaks at 456.2 eV and 456.5 eV are assigned to the Ti oxidation state. (4-x)+ The individual Ti species were assigned to the nuclei (see refs. 5 to 8) by integrating the fitted peaks. 4+ and Ti (4-x)+ The relative proportions of species are calculated. x Ti in 4+ and Ti (4-x)+ The relative ratios were calculated to be 72.9% and 27.1%, respectively.

[0081] On the other hand, STOs y Ti in 4+ and Ti (4-x)+ The relative ratios of STO and SGD were calculated to be 75.2% and 24.8%, respectively. x Ti in (4-x)+ The concentration of the species is STO y Thus, the STO x Thin film is STO y The oxygen deficiency is greater than in thin films.

[0082] Furthermore, the O1s spectrum (Fig. 5C, F) was fitted with two peaks at 529.5 eV and 531.3 eV, which correspond to the O in the STO oxide, respectively. 2-ions (see ref. 3) and C-O bonds formed by adventitious carbon adsorption on the surface (see refs. 3, 4, 9).

[0083] <Electrode> As shown in step (3) of FIG. 1, a pair of gold terminal electrodes are formed on the surface of the deposited oxygen-deficient metal oxide thin film, facing each other at a distance from each other, so as to be in electrical contact with the oxygen-deficient metal oxide thin film that functions as the detection element of the sensor.

[0084] Briefly, the terminal electrodes are formed by depositing a thin gold film (250 nm with a 100 nm chromium adhesion layer) on top of an oxygen-deficient metal oxide layer using electron beam lithography. The deposited gold film is then patterned using standard photolithography and wet etching techniques to form a pair of terminal electrodes.

[0085] <Middle Class> In a preferred embodiment, the bioanalyte binding sites are immobilized on the oxygen-deficient metal oxide layer via an intermediate layer formed using a plurality of long-chain molecules physically or chemically adsorbed to the oxygen-deficient metal oxide layer as surface modifiers.

[0086] In one embodiment, the intermediate layer is produced by silanizing the hydroxyl groups of an oxygen-deficient zinc oxide layer with a silanizing agent having a terminal functionality selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxyl groups, and hydroxyl groups.

[0087] For example, as shown in step (4) of Figure 1, the silanizing agent is (3-glycidyloxypropyl)trimethoxysilane (GPS), an epoxy-terminated silanizing agent. The protocol for adsorbing GPS onto the oxygen-deficient zinc oxide layer is described in the Materials and Methods section below.

[0088] In other embodiments, the surface modifier may be selected from the group consisting of (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS).

[0089] <Analyte binding site> As shown in step (5) of FIG. 1, the now silanized surface of the oxygen-deficient zinc oxide layer is further modified by immobilizing appropriate conjugates or biomolecules onto each end of the immobilized silanizing agent to act as binding sites for selectively binding desired bioanalytes from a biological sample.

[0090] There are a variety of biomolecules available as binding moieties for selectively binding a desired bioanalyte from a biological sample.

[0091] For example, such biomolecules may include proteins, peptides, lipopeptides, protein-binding carbohydrates, or protein-binding ligands.

[0092] In one embodiment, the biomolecule is a capture protein.

[0093] Suitably the capture protein is a protein binding scaffold, a T cell receptor, a binding fragment of a TCR, a variable lymphocyte receptor, an antibody and / or a binding fragment of an antibody.

[0094] <Protein-binding scaffold> Protein-binding scaffolds have emerged as effective molecules for binding a diverse range of bioanalytes, including proteins. Typically, protein-binding scaffolds have a stable protein structure (scaffold) that allows for modification of amino acids within a designated binding region without altering the relative arrangement of the binding domains. Such protein-binding scaffolds include, but are not limited to, adnectins, affilins (nanophytins), affibodies, affima molecules, affitins, α-bodies, aptamers, anticalins, armadillo repeat protein-based scaffolds, avimers, designed ankyrin repeat proteins (DARPins), finomers, inhibitor cystine knot (ICK) scaffolds, Kunitz domain peptides, monobodies (AdNectins™) and nanophytins.

[0095] Affilin is an artificially engineered protein of approximately 20 kDa. It contains a scaffold structurally related to human ubiquitin and vertebrate gamma B crystallin and has eight surface-exposed, engineerable amino acids. Affilins can be engineered to specifically bind target bioanalytes and can be specifically adapted to bind a wide variety of molecules using techniques such as site-directed mutagenesis and phage display libraries (see reference 10).

[0096] Affibodies are approximately 6 kDa proteins that have a protein scaffold of the Z domain of an IgG isotype antibody with modifications to one or more of 13 amino acid residues located in their two α-helical binding domains (see reference 11).

[0097] Affimer molecules are proteins of approximately 12-14 kDa that utilize a protein scaffold derived from the cystatin family of cysteine ​​protease inhibitors. They contain two peptide loop regions in addition to an N-terminal sequence that can accommodate target-specific binding. Affimer molecules with up to 1010 possible amino acid combinations in the binding site can be generated using phage display libraries and appropriate techniques (see reference 12).

[0098] Affitin is a protein of 66 amino acid residues (approximately 7 kDa) that uses a protein scaffold derived from the DNA-binding protein Sac7d found in Sulfolobus acidocaldarius (see reference 13). It is easily produced in vitro from prokaryotic cell cultures, contains 14 binding amino acid residues, and binds 3 × 10 12 These can be mutated to produce the above structural variations (see reference 13). Screening techniques such as surface plasmon resonance can be used to identify specific binding of these molecules.

[0099] Alpha-bodies are molecules of approximately 10 kDa that, unlike most macromolecules, can penetrate cell membranes (unless immobilized) and thus bind to molecules inside and outside the cell. The alpha-body scaffold is based on a computer-designed coiled-coil structure with three alpha-helices (A, B, and C) that are not similar to the native structure. Amino acids on the A and C alpha-helices can be modified to target specific antigens (see reference 14).

[0100] Aptamers for binding to proteins include a variety of nucleic acids (DNA, RNA, and XNA) and peptides that can be screened for binding to specific target molecules. Nucleic acid aptamer databases (see Reference 15) allow the selection of DNA aptamers identified in vitro. Peptide aptamers consist of short amino acid sequences that are typically embedded in a loop structure ("loop on a frame") within a stable protein scaffold frame. Typically, peptide loops of 5 to 20 residues provide a source of variability for selective binding to target molecules. Techniques such as combinatorial libraries and yeast-2 hybrid screening can be used to generate and screen peptide aptamers. Other techniques for generating and screening protein aptamers are also described in the literature (see Reference 16).

[0101] Anticalin proteins are protein-binding molecules derived from lipocalins. Typically, anticalins bind to molecules smaller than antibodies. Methods for screening and developing anticalins have been described in the literature (see references 17 and 18).

[0102] Armadillo repeat protein-based scaffolds are characterized by armadillo domains consisting of tandem armadillo repeats of approximately 42 amino acids organized into superhelices of repeating units, each consisting of three α-helices. (See Reference 19.) Modification of residues within the conserved binding domains allows for the preparation of diverse combinatorial libraries that can be used to select target-specific binders. (See Reference 19.)

[0103] Avimars (also known as avidity multimers, maxibodies, and low density lipoprotein receptor (LDLR) domain A) are composed of at least two 30-35 amino acid long peptides based on the A domain of a cysteine-rich cell surface receptor protein. Varying the A domain allows binding to different epitopes on the same target or multiple targets, and the number of peptides attached determines the number of targets per avimer. Avima phage display libraries are known in the art, including commercial libraries such as those from Creative Biolabs.

[0104] Designed ankyrin repeat proteins (DARPins) are binding proteins engineered from ankyrin proteins. Methods for screening and identifying DARPins have been described in the literature (see references 20, 21).

[0105] Inhibitor cystine knot (ICK) scaffolds are a family of miniproteins (30-50 amino acid residues long) that form stable three-dimensional structures consisting of three disulfide bridges connecting a series of loops with high sequence diversity. There are three families of inhibitor cystine knots: knottins, cyclotides, and growth factor cystine knots. Databases are known in the art, such as the KNOTTIN database (www.dsimb.inserm.fr / KNOTTIN / ), which disclose specific properties such as sequences, structures, and functions of known knottins and cyclotides. Furthermore, methods for producing ICKs and screening for binding have been described in the literature (see reference 22).

[0106] Monobodies (also known commercially as Adnectins) utilize an FN3 (fibronectin type III domain) scaffold with diverse and engineerable variable groups. Adnectins share antibody variable domains and β-stalks with antibodies. The binding affinity of monobodies can be diversified and customized by in vitro evolution methods such as mRNA display, phage display, and yeast display. Methods for screening and generating monobodies have been described in the literature (see references 22, 24).

[0107] <Antibodies and antibody fragments> In some embodiments, the bioanalyte binding moiety is an antibody, or a binding fragment thereof. Antibodies are exemplary protein binding molecules with potentially as many as 1011-1012 unique molecules in an individual, with genetic variation between individuals allowing for further diversity. Antibody diversity in vivo is generated by random recombination of a set of genes at the V(D)J junction.

[0108] The binding of an antibody is determined primarily by three hypervariable regions in the heavy and light chains, called complementarity determining regions (CDRs) 1, 2, and 3. Thus, each mature antibody has six CDRs (variable heavy chain (VH) CDR1, CDR2, and CDR3, variable light chain (VL) CDR1, CR2, and CDR3). These hypervariable regions form a three-dimensional antigen-binding pocket, and the binding specificity of the antibody is determined by the specific amino acid sequences in the CDRs, primarily CDR3.

[0109] Antibodies to specific bioanalytes may be commercially obtained or may be produced by methods known in the art, for example, antibodies to specific bioanalytes may be prepared using methods generally disclosed in the literature (see reference 25).

[0110] The specificity, activity and affinity of antibodies generated in a subject can be modified by in vitro processes such as affinity maturation (see reference 26). In this way, in vivo derived antibodies can be further modified to generate different but phylogenetically related antibodies. Thus, the term "antibody" encompasses in vivo derived antibodies as well as exogenously derived molecules that have undergone a process of mutation to modify the CDR binding sites so that they have unique sequences compared to antibodies generated in vivo.

[0111] The term antibody also includes non-conventional antibodies generated from species such as camelids, sharks, jawfish, etc. Thus, the term antibody includes camelid antibodies, heavy chain antibodies including IgNAR and variable lymphocyte receptors (VLR). Furthermore, these can be fragmented into their binding moieties (such as a single binding moiety of VNAR-IgNAR) or recombinantly integrated into fusion proteins. Methods for generating and adapting such non-conventional antibodies are described in the literature (see references 27, 28).

[0112] <Antibody-binding fragment> In some embodiments, the bioanalyte binding moiety is an antibody binding fragment. The antibody binding fragment may be derived from an antibody or may be a recombinant product having sequences identical to the CDRs of an antibody or antibody fragment. In fact, these CDRs may be from affinity matured antibodies and therefore may not be identical to the antibody of in vivo origin.

[0113] Antibodies consist of four chains (two heavy and two light) and can be separated into Fc (fraction crystallizable) and Fab (fraction antibody) domains. The Fc portion of an antibody interacts with Fc receptors and the complement system. As a result, the Fc portion is important for the immune function of the antibody. However, the Fab portion contains the binding region of the antibody and is important for the specificity of the antibody for the desired epitope.

[0114] Thus, in some embodiments, the bioanalyte binding site is a Fab fragment of an antibody. The Fab fragment may be an individual Fab fragment (i.e., the antibody fragment is produced without linked disulfide bridges) or a F(ab')2 fragment consisting of two Fab fragments of an antibody linked via disulfide bridges. These fragments are usually produced by fragmenting the antibody with a digestive enzyme such as pepsin. Methods are described in the literature (see reference 29).

[0115] Each Fab fragment of an antibody has six CDRs, three each for the VH and VL chains (within a framework of four framework regions). The constant regions of the Fab fragment can be removed, leaving only the VH and VL regions. Individual VH and VL chains (each consisting of only three CDRs) have been shown to bind specifically with high affinity. Typically, individual binding regions are known as single antibody domains (sdAbs). VH and VL chains can also be joined by a linker to form a fusion protein called single chain variable fragment (scFv - also called diabody). Unlike Fab, scFv is not fragmented from an antibody, but is usually formed recombinantly based on the CDRs and framework regions of an antibody. Additionally, sdAbs may also be recombinantly produced and contain moieties that may form the binding component of a larger fusion protein, act to stabilize the binding region, improve or facilitate immobilization to a sensing element or intermediate layer, improve binding by providing flexibility of the binding region, or optimize the length of the bioanalyte binding site, for example, by allowing access to an antigenic region of the bioanalyte. Thus, in some embodiments, the bioanalyte binding site is or comprises an scFv or sdAb. The scFv may comprise multiple VH and VL chains linked together to form a multivalent scFv, such as a di-scFv or tri-scFv.

[0116] Antibodies and antibody fragments, or fusion proteins containing antibody-derived sequences, directed against specific bioanalytes may be obtained commercially or produced by methods known in the art, such as those described above.

[0117] <Proteins and receptors> Protein receptors or ligands that interact with and bind to proteins can be used as bioanalyte binding sites. Such receptors and ligands include whole receptors or ligands, or specific fragments thereof (e.g., fragments that constitute the binding domain of the receptor or ligand). Particularly contemplated receptors include receptors for cytokines, such as interleukins, or chemokines that can provide information about the state of the immune system. In some embodiments, the receptors or ligands (or fragments thereof) may be combined to form a fusion protein.

[0118] For example, interleukin-6 (IL-6) is a proinflammatory cytokine and an important biomarker that can be used to monitor immune responses during cancer treatment, as well as psychological stress and insulin activity.

[0119] For example, the present inventors have been able to obtain good results when using anti-interleukin-6 (IL-6) antibodies to selectively recognize and bind IL-6.

[0120] For example, the present inventors have been able to obtain good results using an anti-C-reactive protein (CRP) antibody to selectively recognize and bind to CRP.

[0121] <Summary> In summary, the non-invasive conductance sensor described above is a passive electronic device constructed in a simple in-plane two-terminal electrode geometry, where the sensing element of the sensor takes the form of an oxygen-deficient metal oxide thin film applied to the surface of a sensor substrate and subsequently functionalized with specific bioanalyte binding sites that are selective for one or more bioanalytes to be detected in bodily fluids such as human saliva and / or sweat. When a voltage is applied across the sensor, an electrical signal is generated that is proportional to the change in conductance of the sensing element that is generated as a result of charge transfer between the complex formed between the bioanalyte and the bioanalyte binding sites and the oxygen-deficient metal oxide thin film layer. This electrical signal can be equated to the level of the target biomarker or bioanalyte present in the bodily fluid.

[0122] Next, a method for detecting a target bioanalyte in a body fluid using the above-mentioned non-invasive electrical conductivity sensor will be described.

[0123] <Detection method> In accordance with another preferred embodiment of the present invention, a method for detecting a bioanalyte is provided.

[0124] Briefly, a method for detecting a bioanalyte using an oxygen-deficient metal oxide-based sensor comprises the steps of: (i) contacting the oxygen-deficient metal oxide-based sensing element with a sample solution of a body fluid containing the bioanalyte; (ii) applying a voltage across the sensor; and (iii) using a current source meter to detect an electrical signal generated between a pair of terminal electrodes that is proportional to a change in conductance corresponding to detection of the bioanalyte when the bioanalyte binds to the bioanalyte binding sites on the surface of the oxygen-deficient metal oxide sensing element. EXAMPLES

[0125] <Antigen concentration dependency test> Both antigens, IL-6 and CRP, show a concentration-dependent change in resistance with respect to the baseline resistance of the device. The baseline resistance was measured for the antibody-immobilized-GPS silanized sensor before the addition of antigen. A linear correlation is observed for the change in resistance as a function of antigen concentration for both IL-6 and CRP antigens (see Figure 2). The responsivities (i.e., slopes of the curves) for IL-6 and CRP are 5.1 and 4.1% / M, respectively. These values ​​suggest that the anti-IL-6 antibody-immobilized ZnO sensor exhibits higher sensitivity for detecting IL-6 antigen compared to the anti-CRP antibody-immobilized ZnO conductometric sensor for detecting CRP antigen. To investigate the contribution from the matrix of the antigen solution, the resistance of both types of antibody-immobilized ZnO sensors to a PBS solution was measured. For both types of antibody-immobilized sensors, the change in resistance in the presence of a PBS solution was less than 1%. This indicates that the contribution from the matrix of the antigen solution to the resistance change is negligible.

[0126] For both antigens, the non-invasive zinc oxide conductivity sensor showed detectable responses at concentrations below those found in healthy human body fluids. The concentration of IL-6 in healthy human sweat is about 0.38 pM (10 ng / L) (see Reference 30), and the concentration of CRP in healthy human saliva is about 12 pM (285 ng / L) (see Reference 31). The ZnO conductivity sensor showed a resistance change of 9.2% for 4 fM, which is 100 times lower than the IL-6 concentration in healthy human sweat. Similarly, the resistance change for the lowest CRP concentration detected by the sensor (13 fM) was ~12.5%, which is almost 1000 times lower than the CRP concentration in healthy human saliva. This high sensitivity response to concentrations much lower than those found in healthy human body fluids clearly demonstrates the importance of ZnO-based conductivity sensors in detecting bioanalytes in human body fluids.

[0127] <Cross-selectivity test 1> A cross-selectivity test was conducted to investigate the feasibility of each antibody-immobilized device in the presence of other antigens. The anti-IL-6 antibody-immobilized device showed a 3% change in resistance in the presence of 13 pM CRP, and the CRP antibody-immobilized device showed a 3.5% change in resistance in the presence of 4 pM IL-6 (see Figure 3). The antigen concentrations used in this experiment were selected to be as close as possible to the antigen concentrations in healthy human body fluids. When both antigens were mixed and added to the device, the anti-IL-6 antibody-immobilized device showed a 17.6% change in resistance, which was about 4% lower than when 4 pM IL-6 was added. On the other hand, the resistance change of the anti-CRP antibody-immobilized device was 27.4%, which was about 3% higher than that of 13 pM CRP on the same device. The results of this cross-selectivity test revealed that the anti-IL-6 antibody-immobilized ZnO device showed selectivity for IL-6 in the presence of CRP antigen, and the anti-CRP antibody-immobilized ZnO device showed selectivity for CRP antigen in the presence of IL-6 antigen. The reason for the 4% decrease in conversion rate of the IL-6 antibody device and the 3% increase in conversion rate of the CRP antibody device in the presence of the antigen mixture is currently unknown.

[0128] <Cross-selectivity test 2> A cross-selectivity study was performed to determine the feasibility of each antibody immobilized device in the presence of other antigens. Cathelicidin, B natriuretic peptide (BNP), cardiac troponin I (cTnI) were used in this study along with IL-6 and CRP antigens. The nominal concentrations of the original and mixed antigens were maintained at 4 pM. When the antigen mixture was reacted with the corresponding antibody immobilized device, a significantly higher resistance change was observed for the target antigen (see Figure 6). When treated with the IL-6 antibody immobilized device, a 31% resistance change was observed for IL-6 in the antigen mixture. Only a 3% resistance change was observed for CRP. Also, the CRP antibody immobilized device showed a 30% resistance change for CRP in the antigen mixture, with IL-6 contributing 5% of the resistance change. It is important to note that cathelicidin, BNP, and cTnI did not contribute to the resistance change in the IL-6 and CRP antibody immobilized devices. The small contribution of IL-6 and CRP antigens to the resistance change may be due to the physical adsorption of these molecules to the antibody, and the negative contribution of cathelicidin, BNP, and cTnI to the resistance change suggests that these three antigens were not bound to the tested antibody or that the charge transfer effect of these antigens employs a different mechanism than that of CRP and IL-6 antigens on the antibody.

[0129] [material and method] <Sensor manufacturing> The sensors were fabricated by depositing 100 nm thick thin films of metal oxides such as zinc oxide (ZnO), which act as the sensing element in the sensor, on rigid (SiO2 / Si) and flexible plastic (polyimide foil) substrates. The composition of the sensing element ranges from 0.08 to 2 Siemens / m 2 More preferably, 0.08 to 0.6 Siemens / m 2 The reactive sputtering technique is designed to produce oxygen-deficient metal oxide films with conductances in the range of 16×10. For conductance measurements, two terminal in-plane electrodes are used. -6 m 2It is patterned with the detection area. The change in conductance corresponding to the distributed antibody and antigen is measured using a commercially available current source meter (B2901A Precision Source / Measurement Unit manufactured by Keysight Technologies).

[0130] <Preparation of Antibody and Antigen Solutions> Interleukin-6 (IL-6), anti-IL-6, C-reactive protein (CRP), and anti-CRP were purchased from a commercial vendor (Sigma-Aldrich) and used as received. Cathelicidin, B-type natriuretic peptide (BNP), and cardiac troponin I (cTnI) were supplied from multiple commercial vendors (Abcam, MyBioSource, and ProSpec Bio) and used as received.

[0131] The anti-IL-6 stock solution was diluted 1:106 with phosphate buffer saline (PBS, pH 7.4) and immobilized on the surface of the oxygen-deficient ZnO sensor. The anti-CRP solution was diluted 1:50 with PBS (pH 7.4) before immobilization. The received IL-6 powder was completely dissolved in a known amount of autoclaved Milli-Q water and diluted with a PBS solution of pH 7.4 to prepare a standard series of IL-6 solutions. The prepared IL-6 concentrations were 4 nM, 100 pM, 4 pM, 100 fM, and 4 fM. Also, the received CRP solution was diluted with a predetermined amount of PBS of pH 7.4 to prepare a standard series of CRP solutions. The prepared CRP concentrations were 13 nM, 100 pM, 13 pM, 100 fM, and 13 fM. Solutions of other antigens were prepared in the same manner.

[0132] <Preparation of GPS-Silanized Sensors> Silanization of oxygen-deficient metal oxide sensor surfaces with 3-glycidyloxypropyltrimethoxysilane (GPS) has been reported previously for invasive sensors (see reference 32). Here, silanization of ZnO was performed with minor modifications to the reported silanization procedure. Briefly, freshly prepared ZnO devices were exposed to O2 plasma (Plasma Cleaner PDC-002, Harrick Plasma) for 10 min to activate the hydroxyl groups on the ZnO surface. Next, 20 μL of freshly prepared GPS solution was dropped onto an Al foil, which was placed in a vacuum desiccator to allow GPS vapor to accumulate in the desiccator. The O2 plasma-cleaned ZnO sensor was then exposed to this GPS vapor for 30–45 min. The exposure of the plasma-cleaned ZnO sensor to GPS vapor was performed in an LC200 glove box system. After the exposure to GPS vapor was completed, the ZnO sensor was thoroughly washed with Milli-Q water for 2 min to remove unbound silane groups from the surface. The washed ZnO sensor was then heated at 150 °C for 10 min to enhance the binding of silane groups to the ZnO surface. This GPS-silanized sensor was then used for antibody immobilization.

[0133] It will be appreciated that silanization of an oxygen-deficient strontium titanium oxide (STO) sensor surface with 3-glycidyloxypropyltrimethoxysilane (GPS) or an alternative silanization agent can be carried out in a similar manner as described above.

[0134] <Immobilization of antibodies> Antibody (IgG) immobilization onto GPS-silanized ZnO sensors has been reported previously for invasive sensors (see reference 32). 25 μL of freshly prepared 1:106 diluted anti-IL-6 solution was dropped onto the surface of each freshly GPS-silanized ZnO sensor and incubated for 2 h to allow the immobilization of IL-6 antibody onto the ZnO sensor surface. The sensor was then rinsed with a PBS solution at pH 7.4 to remove unbound antibody. The PBS-washed ZnO sensors were then dried in a stream of N2 gas. These anti-IL-6 antibody immobilized sensors were used for IL-6 antigen concentration and cross-selectivity measurements. Using a similar procedure, CRP-immobilized GPS-silanized ZnO sensors were prepared using 25 μL of freshly prepared 1:50 diluted anti-CRP solution.

[0135] <Measurement of electrical conductivity of antigen> The baseline conductance of the antibody-immobilized ZnO sensor was measured before adding the antigen. 15 μL of antigen solution was dropped onto the surface of the antibody-immobilized ZnO sensor and incubated for 10 min. The antigen solution remaining on the sensor was then removed, and the sensor surface was dried under a stream of N2 gas. Conductance measurements were then performed for each concentration. For the cross-selectivity measurement, 4 pM IL-6 solution (15 μL) was dropped onto the anti-CRP antibody-immobilized ZnO device, and 13 pM CRP solution (15 μL) was dropped onto the anti-IL-6 antibody-immobilized ZnO device. The nominal concentrations of IL-6 antigen and CRP antigen in the premix solution for the cross-selectivity measurement were 4 pM and 13 pM, respectively.

[0136] <Device Shelf Life Test> GPS-silanized devices are stable for at least 15 months. This conclusion was drawn based on the resistance changes of IL-6 and CRP antigens obtained for 15-month (450-day) aged devices (see Figure 7). The resistance change for both antigens showed a linear change with their concentration. A similar linear trend was observed for both types of antigens in freshly prepared devices (0 days old). The almost overlapping standard errors of the resistance changes for a given concentration of selected antigens strongly suggests that there is no statistical difference between the mean resistance changes of fresh and old devices for a given concentration of selected antigen.

[0137] <Device performance in artificial saliva> The antibody-immobilized devices for IL-6 and CRP successfully detected the corresponding antigens in artificial saliva (see Figure 8). The resistance change for both types of antigens varied linearly with their concentrations. The reactivity (i.e. slope) for IL-6 and CRP in artificial saliva was 1.6 (% / M) and 1.2 (% / M), respectively, suggesting that the devices were more sensitive to IL-6 than to CRP in artificial saliva. The resistance change for a given concentration of these antigens was significantly lower in artificial saliva with respect to the resistance change in PBS. The large background contribution from artificial saliva (83% for IL-6 and 85% for CRP) was responsible for this significantly lower resistance change. This could be due to the higher ionic composition of artificial saliva compared to PBS, which resulted in a higher charge transfer effect. Also, when the devices were immobilized with the corresponding antibodies, similar to the effect in PBS, these devices showed significant selectivity for the target antigens in artificial saliva. For example, the resistance change for IL-6 in the antigen mixture prepared in artificial saliva was significantly higher in the presence of the IL-6 antibody immobilized device than for other antigens. Similarly, the CRP antibody immobilized device showed a significantly higher resistance change rate for the CRP antigen in the antigen mixture prepared in artificial saliva.

[0138] [advantage] From the foregoing discussion, it will be apparent to those skilled in the art that the non-invasive oxygen-deficient metal oxide-based conductometric sensors of the various embodiments of the present invention offer numerous advantages over existing counterparts.

[0139] Indeed, oxygen-deficient metal oxide-based conductometric sensors are capable of measuring concentrations of target bioanalytes lower than their respective levels in human body fluids. Therefore, the inventors believe that such sensors offer great potential for the development of cost-effective, biocompatible, and functional sensors with broad applicability as personalized, reusable healthcare monitoring devices. Indeed, the inventors highly anticipate that these sensors may have a significant impact on: 1. Alerting for cardiovascular disease: Regular testing with these sensors could alert people to elevated levels of inflammatory biomarkers, allowing for intervention before a heart attack or stroke occurs. 2. Cancer treatment: It is expected to be useful in monitoring biomarkers for cancer and its treatment. 3. Treatment of abdominal diseases: By detecting CRP antigens, it is expected that abdominal diseases, such as diseases of the body, can be detected.

[0140] The oxygen-deficient metal oxide-based electrical conductivity sensor can be combined with conventional portable integrated circuits and wearable electronic devices to form portable devices that can be worn as needed.

[0141] In fact, oxygen-deficient metal oxide-based conductometric sensors can be fabricated on any type of insulating or plastic substrate while maintaining selectivity for the target biomarkers or bioanalytes, and they are also reusable, which offers significant cost benefits.

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[0143] [Definition] Whenever a range is given herein, e.g., a temperature range, a time range, or a concentration range, all intermediate and subranges, as well as all individual values ​​contained within the given range, are intended to be included in the disclosure. It will be understood that any subrange or individual value within a range or subrange contained in the description herein can be excluded from the claims herein.

[0144] All definitions defined and used herein should be understood to control over dictionary definitions, definitions in documents incorporated by reference, and / or ordinary meanings of the defined terms.

[0145] As used herein, the indefinite articles "a" and "an" should be understood to mean "at least one," unless clearly indicated to the contrary.

[0146] The phrase "and / or" as used herein should be understood to mean "either or both" of the elements so conjoined, i.e., elements that are conjunctive in some cases and disjunctive in other cases. Multiple elements listed with "and / or" should be interpreted in the same manner, i.e., to mean "one or more" of the elements so conjoined. Other elements can optionally be present other than the elements specifically identified by the "and / or" clause, whether related to the elements specifically identified or not. Thus, as a non-limiting example, a reference to "A and / or B," when used in conjunction with open-ended language such as "comprising," can refer in one embodiment to only A (optionally including elements other than B), in another embodiment to only B (optionally including elements other than A), in yet another embodiment to both A and B (optionally including other elements), etc.

[0147] While the present invention has been described in connection with a limited number of embodiments, it will be appreciated by those skilled in the art that many alternatives, modifications, and variations are possible in light of the foregoing description. Accordingly, the present invention is intended to embrace all such alternatives, modifications, and variations that may fall within the spirit and scope of the present invention as disclosed.

[0148] When the terms "comprise", "comprises", "comprised" or "comprising" are used in this specification (including the claims), they are to be interpreted as specifying the presence of stated features, integers, steps or components, but not excluding the presence of one or more other features, integers, steps or components, or groups thereof. [Prior art documents] [Patent documents]

[0149] Patent Document 1: U.S. Patent Application Publication No. 2010 / 2016256 Patent Document 2: U.S. Patent Application Publication No. 2010 / 0305304 Patent Document 3 Korean Patent No. 101774891 Patent Document 4 China Patent Application Publication No. 102706939 Patent Document 5: U.S. Patent No. 8,133,698

Non-licensed documents

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Claims

1. 1. A sensor for detecting a bioanalyte, comprising: A substrate; a pair of terminal electrodes arranged on the substrate in an opposing relationship with a gap between them; A sensor comprising: a non-insulating sensing element applied to a surface of the substrate between and in electrical contact with a pair of terminal electrodes, the sensing element providing a conductive path between the terminal electrodes, the sensing element having an oxygen-deficient metal oxide layer and a bioanalyte binding site, wherein when a voltage is applied across the sensor an electrical signal is generated that is proportional to a change in conductance of the sensing element corresponding to binding of a bioanalyte to the bioanalyte binding site.

2. the oxygen-deficient metal oxide layer is formed from a metal oxide selected from the group consisting of zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium dioxide; Further, the oxygen-deficient metal oxide layer has a thickness ranging from about 50 nm to about 200 μm.

2. The sensor according to claim 1 .

3. 3. The sensor of claim 1 or 2, wherein the bioanalyte binding sites are immobilized on the oxygen-deficient metal oxide layer via an intermediate layer that is physically or chemically adsorbed on the oxygen-deficient metal oxide layer.

4. 4. The sensor of claim 3, wherein the intermediate layer is fabricated by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functionality selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxyl groups, and hydroxyl groups.

5. 5. The sensor of claim 4, wherein the silanizing agent is selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyltrimethoxysilane (AEAPTS).

6. The oxygen-deficient metal oxide layer has a resistivity of about 0.08 Siemens / m ２ to about 0.6 siemens / m ２ 6. The sensor according to claim 1, having a conductance falling within the range

7. the bioanalyte binding site is a biomolecule; Further, the biomolecule is a protein, a capture protein, a peptide, a lipopeptide, a protein-binding carbohydrate, or a protein-binding ligand.

7. A sensor according to any one of claims 1 to 6.

8. the capture protein is a protein binding scaffold, a T cell receptor, a binding fragment of a TCR, a variable lymphocyte receptor, an antibody and / or a binding fragment of an antibody; the protein-binding scaffold is selected from the group consisting of adnectins, affilins, affibodies, affimer molecules, affitins, alpha bodies, aptamers, anticalins, armadillo repeat protein-based scaffolds, atrimers, avimers, designed ankyrin repeat proteins (DARPins), finomers, inhibitor cystine knot (ICK) scaffolds, Kunitz domain peptides, monobodies and / or nanophytins; The binding fragment of an antibody comprises a Fab, (Fab') 2 , Fab', single chain variable fragment (scFv), di- and tri-scFvs, single domain antibodies (sdAb), diabodies or fusion proteins comprising the binding domain of an antibody; 8. The sensor according to claim 7.

9. The sensor of any one of claims 1 to 8, wherein the bioanalyte binding site binds interleukin-6 (IL-6) or C-reactive protein (CRP).

10. 10. The sensor of claim 1, wherein the substrate is made of a material selected from the group consisting of a silicon wafer, a polymer, a glass and a ceramic.

11. 11. The sensor of claim 10, wherein the polymer is selected from the group consisting of polydimethylsiloxane (PDMS), polyimide (PI) and polyethylene naphthalate (PEN).

12. The ceramic is made of aluminum oxide (Al ２ O ３ ), sapphire, and silicon nitride (Si ３ N ４ 11. The sensor of claim 10, wherein the sensor is selected from the group consisting of:

13. 1. A method for detecting a biological analyte, comprising: a) contacting a sensing element of a sensor according to any one of claims 1 to 12 with a sample solution containing a bioanalyte; b) applying a voltage across the sensor; c) detecting an electrical signal generated proportional to a change in conductance corresponding to detection of the bioanalyte upon binding of the bioanalyte to the bioanalyte binding site; A method for providing the above.

14. 14. The method of claim 13, wherein the bioanalyte binding site binds interleukin-6 (IL-6).

15. 14. The method of claim 13, wherein the bioanalyte binding site binds C-reactive protein (CRP).

16. 1. A method of manufacturing a sensor for detecting a bioanalyte, comprising: Providing a substrate; depositing a pair of terminal electrodes on the substrate in a spaced-apart, opposing relationship; applying between and in electrical contact with said pair of terminal electrodes a non-insulating sensing element in the form of an oxygen deficient metal oxide layer coated with bioanalyte binding sites, said sensing element providing an electrical conductive path between said terminal electrodes, said bioanalyte binding sites being selective for detection of said bioanalyte upon binding to said bioanalyte binding sites; A method for providing the above.

17. the oxygen-deficient metal oxide layer is formed from a metal oxide selected from the group consisting of zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium dioxide; the oxygen-deficient metal oxide layer has a thickness in the range of about 50 nm to about 200 μm; 17. The method of claim 16.

18. 17. The method of claim 16, wherein the oxygen-deficient metal oxide layer is applied to the substrate surface by a technique selected from the group consisting of reactive sputtering, physical vapor deposition (PVD), chemical vapor deposition (CVD), metal-organic chemical vapor deposition (MOCVD), pulsed laser deposition (PLD) and molecular beam epitaxy (MBE).

19. 17. The method of claim 16, further comprising physically or chemically adsorbing an intermediate layer to the oxygen deficient metal oxide layer to immobilize the bioanalyte binding sites thereon.

20. 20. The method of claim 19, wherein the intermediate layer is produced by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functionality selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxyl groups, and hydroxyl groups.

21. 21. The method of claim 20, wherein the silanizing agent is selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-2(2-aminoethyl)-3-aminopropyltrimethoxysilane (AEAPTS).

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