Sensors and detection methods for detecting bioanalytes

Abstract

The present invention provides a sensor for detecting a bioanalyte, comprising: - a substrate; - a pair of end electrodes disposed on the substrate in a mutually spaced and opposite relationship; and - a non-insulating sensing element applied to the surface of the substrate, in electrical contact between and with the end electrodes, wherein the sensing element provides a conduction path between the end electrodes, wherein the sensing element includes an oxygen-deficient metal oxide layer and a bioanalyte binding site, and wherein when a voltage is applied through the sensor, an electrical signal proportional to the change in conductivity of the sensing element is generated, the electrical signal corresponding to the binding of the bioanalyte to the bioanalyte binding site.

Description

Technical Field

[0001] This invention relates to sensors, and more particularly to non-invasive sensors for detecting bioanalytes in bodily fluids and methods thereof.

[0002] This invention is primarily developed for the detection of a range of bioanalytes in bodily fluids, and will be described below with reference to this application.

[0003] The following discussion of the background of the invention is intended to facilitate understanding of the invention. However, it should be understood that this discussion is not an affirmation or acknowledgment that any material mentioned herein has been published, is known, or is part of the general public knowledge in Australia or any other country by the priority date of any claim in this specification. Background Technology

[0004] There are two approaches for monitoring / measuring the levels of target biomarkers (hereinafter referred to as bioanalytes) in tissues and / or biofluids. The first approach relies on the use of invasive sensors, where the components of the sensor come into direct contact with the tissue or body fluids, potentially leading to infection, tissue damage, and discomfort. The second approach relies on the use of non-invasive sensors, employing various techniques to determine the levels of bioanalytes in body fluids, including optical absorption, electrochemical, transduction, and conductivity assays.

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

[0006] On the other hand, electrochemical sensors are more precise and therefore currently dominate the field of biosensing. These sensors operate by measuring the electrical signal generated by the reaction of a bioanalyte of interest with a sensing element associated with the sensor, where the generated electrical signal is proportional to the concentration of the bioanalyte. This electrochemical reaction in the sensor functions by: initiating a measurable current (amperometric measurement), a measurable charge accumulation or potential (potential measurement), altering the conductivity of the medium (conductivity measurement), or measuring the combination of resistance and reactance to form an impedance (impedance measurement).

[0007] Sensors using electrochemical transduction typically require a working electrode, a reverse (or auxiliary) electrode, and a reference electrode. The reference electrode is positioned at a distance from the site where the biorecognition element and analyte interact to establish a known and stable potential. When an interaction occurs, the working electrode acts as the transduction component, while the reverse electrode measures the current and facilitates the transport of the electrolyte solution to allow current to flow to the working electrode.

[0008] Conductivity sensors also rely on the use of electrodes to measure the current conducted through a medium. However, conductivity sensors do not require the use of reference electrodes. These sensors also operate at low-amplitude AC voltages, thus preventing Faraday processes on the electrodes, and can be easily miniaturized and integrated using thin-film technology.

[0009] While conductivity sensors offer certain advantages, their sensitivity is often hampered by the use of polymers as sensing elements, which frequently results in poor durability and long-term stability.

[0010] As an alternative to direct conductivity sensors, field-effect transistor (FET)-based sensors have also been developed. FETs are devices with three terminals: a source, a gate, and a drain. These devices operate on the principle that a change in the gate causes a field effect, which alters the conductivity between the source and drain.

[0011] For example, US 2010 / 2016256 describes a biosensor comprising: a substrate, a source electrode on the substrate, a drain electrode on the substrate, and at least one functionalized nanoribbon on a substrate surface between the source and drain electrodes, wherein the functionalized nanoribbon has a chemically functionalized surface connected to one or more detector molecules for binding to a bioanalyte to be detected, thereby generating an electric field gating effect through the binding of the analyte to the one or more detector molecules connected to the nanoribbon surface. The device operates by the molecular binding altering the field effect of the nanoribbon (gate), thereby changing the conductivity of the pathway between the source and drain electrodes, and the change in conductivity can be monitored. Typically, this type of device has two drawbacks.

[0012] First, field-effect transistors are typically devices that switch on and off and have a non-linear response. In these devices, the resistance does not change linearly because they typically have a small linear response region before leveling off, meaning that such devices are difficult to use under a wide range of conditions.

[0013] Secondly, as those skilled in the art will understand, in order for such a device to operate as described, an insulating (dielectric) layer is necessary between the conduction path (located between the source and drain) and the source of the gate bias (in this case, a nanoribbon). Therefore, a disadvantage of this type of device is that, due to the number of different structural elements, they are relatively complex to manufacture and thus more difficult to produce on an industrial scale than sensors with simpler structures.

[0014] The present invention seeks to provide a sensor and a detection method for detecting bioanalytes, which will overcome or substantially improve at least some of the defects of the prior art, or at least provide an alternative. Summary of the Invention

[0015] According to a first aspect of the invention, a sensor for detecting a bioanalyte is provided, comprising: a substrate; a pair of end electrodes disposed on the substrate in a mutually spaced and opposite relationship; and a non-insulating sensing element applied to a surface of the substrate, between and in electrical contact with the end electrodes, wherein the sensing element provides a conduction path between the end electrodes, wherein the sensing element includes an oxygen-deficient metal oxide layer and an analyte binding site, and wherein when a voltage is applied across the sensor, an electrical signal proportional to a change in the conductivity of the sensing element is generated, the electrical signal corresponding to the binding of the analyte to the analyte binding site.

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

[0017] In one embodiment, the thickness of the oxygen-deficient metal oxide layer falls 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 using a technique selected from the following: 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).

[0019] Preferably, the binding sites of the bioanalyte are anchored to the oxygen-deficient metal oxide layer by physical or chemical adsorption to the intermediate layer of the oxygen-deficient metal oxide layer.

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

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

[0022] In one embodiment, the conductivity of the oxygen-deficient metal oxide layer falls at approximately 0.08 Siemens / m. 2 Approximately 0.6 Siemens / m 2 Within the range.

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

[0024] Appropriately, biomolecules are proteins, peptides, lipopeptides, protein-bound carbohydrates, or protein-bound ligands.

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

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

[0027] Preferably, the protein-binding scaffold is selected from: Adnectin, Affilin, Affibody, Affimer molecule, Affitin, Alphabody, Aptamer, Anticalins, armadillo repeat protein-based scaffold, Atrimer, Avimer, Designed Ankyrin Repeat Protein (DARPin), Fynomer, inhibitory cystine knot (ICK) scaffold, Kunitz domain peptide, Monobody and / or Nanofitin.

[0028] Preferably, the antibody binding fragment includes Fab, (Fab')2, Fab', single-chain variable fragment (scFv), di-scFv and tri-scFv, single-domain antibody (sdAb), biantibody, or fusion protein containing an antibody binding domain.

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

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

[0031] Preferably, the substrate is made of a material selected from silicon wafers, polymers, glass, and ceramics.

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

[0033] Suitable ceramics are selected from alumina (Al2O3), sapphire, and silicon nitride (Si3N4).

[0034] According to a second aspect of the present invention, a method for detecting a bioanalyte is provided, the method comprising the steps of: contacting a sensing element of a sensor according to a first aspect with a sample solution containing the bioanalyte; applying a voltage through the sensor; and detecting a generated electrical signal, the electrical signal being proportional to a change in conductivity corresponding to the detection of the bioanalyte when it binds to a bioanalyte binding site.

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

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

[0037] Suitablely, the change in conductivity detected in a sample solution with an IL-6 concentration of 4 femtomoles was approximately 9.2%.

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

[0039] Suitablely, the conductivity change detected in a sample solution with a CRP concentration of 13 femtomoles was approximately 12.5%.

[0040] According to a third aspect of the invention, a method for manufacturing a sensor for detecting a bioanalyte is provided, the method comprising the steps of: providing a substrate; depositing a pair of end electrodes on the substrate in a spaced-apart and opposite relationship; applying a non-insulating sensing element between the end electrodes and in electrical contact with the end electrodes, the sensing element being in the form of an oxygen-deficient metal oxide layer coated with bioanalyte binding sites, wherein the sensing element provides a conduction path between the end electrodes, wherein the bioanalyte binding sites are selective for detecting the bioanalyte when the bioanalyte binds to them.

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

[0042] In one embodiment, the thickness of the oxygen-deficient metal oxide layer falls 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 using a technique selected from the following: 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).

[0044] Suitablely, the method further includes the step of: physically or chemically adsorbing an intermediate layer onto an oxygen-deficient metal oxide layer for anchoring a first biomolecule to the oxygen-deficient metal oxide layer.

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

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

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

[0048] Although any other forms may fall within the scope of this invention, preferred embodiments of the invention will now be described by way of example only and with reference to the accompanying drawings, wherein:

[0049] Figure 1 A schematic diagram of the manufacture of a non-invasive conductivity sensor for detecting bioanalytes according to a preferred embodiment of the present invention is shown, wherein the sensor has a sensing element comprising an oxygen-deficient metal oxide thin film layer, and a plurality of bioanalyte binding sites are coupled to the oxygen-deficient metal oxide thin film layer.

[0050] Figure 2 The graphs show the change in resistance (%) as a function of the concentration (M) of the following substances: (A) IL-6 on a conductivity sensor immobilized with anti-IL-6 antibody, and (B) CRP on a conductivity sensor immobilized with anti-CRP antibody. The dashed lines in each graph represent the antigen concentration (M) in healthy human body fluids (i.e., IL-6 in sweat and CRP in saliva);

[0051] Figure 3 The graphs showing the following resistance changes (%) are presented: (A) CRP on a conductivity sensor immobilized with anti-IL-6 antibody, and (B) IL-6 on a conductivity sensor immobilized with anti-CRP antibody, performed for cross-selectivity studies. The nominal concentrations (M) of IL-6 and CRP are 4 pM and 13 pM, respectively.

[0052] Figure 4 The images show ZnO sputtered from ZnO. x and ZnO y Nuclear-level XPS spectra of (a,c)Zn 2p and (b,d)O 1s collected from thin films, the thin films being formed on... Figure 1 On the substrate surface of the conductivity sensor. Ov represents oxygen vacancy;

[0053] Figure 5 The resolved nuclear-level XPS spectra of all three elements (a,d)Sr, (b,e)Ti, and (c,f)O in the sputtered STO thin film are shown. Figure 1 The substrate surface of the conductivity sensor has different chemical compositions;

[0054] Figure 6 The graphs show the following resistance changes (%): (a) selectivity for IL-6 antigen on the device immobilized with IL-6 antibody in the presence of other study antigens; (b) selectivity for CRP antigen on the device immobilized with CRP antibody in the presence of other study antigens.

[0055] Figure 7 The following graphs are shown: (a) Resistance changes of the new (day 0) and old (day 450) units as a function of IL-6 concentration; (b) Resistance changes of the new (day 0) and old (day 450) units as a function of CRP concentration; and

[0056] Figure 8 The following graphs are shown: (a) Electrical resistance change as a function of IL-6 concentration in PBS and artificial saliva. (b) Selectivity of the device with immobilized IL-6 antibody for IL-6 in the presence of other antigens used in the work. (c) Electrical resistance change as a function of CRP concentration in PBS and artificial saliva. (d) Selectivity of the device with immobilized CRP antibody for CRP in the presence of other antigens used in the work. Detailed Implementation

[0057] It should be noted that in the following description, the same or similar reference numerals in different embodiments denote the same or similar features.

[0058] This invention is based on the discovery of an inexpensive, non-invasive sensor employing conductivity sensing technology to detect the levels of a range of bioanalytes in bodily fluids (such as human saliva and / or sweat) for the prognosis / diagnosis of medical conditions. As will be described in more detail below, the conductivity sensor has a simple and relatively easy-to-manufacture device structure, providing a cost-effective alternative to traditional non-invasive sensors that require specialized substrates or sensing technologies that limit the accuracy of results.

[0059] The inventors believe that the conductivity sensor, described in more detail below, is compatible with CMOS circuits and can therefore be easily integrated with flexible / wearable electronic devices to provide portable, personalized, and reusable sensors that can be used to continuously monitor the levels of target bioanalytes via bodily fluids without invasive procedures. These bioanalytes can serve as biomarkers indicating an individual's condition and health.

[0060] The following is a detailed description of a non-invasive conductivity sensor and its application in detecting levels of a range of bioanalytes (such as biomarkers) in bodily fluids.

[0061] sensor

[0062] A sensor for detecting bioanalytes according to a preferred embodiment of the present invention will now be described.

[0063] In its simplest form, and as Figure 1 As shown in the schematic diagram, the sensor includes a substrate, a pair of end electrodes disposed on the substrate in a spaced-apart and opposite relationship, and a non-insulated sensing element applied to the surface of the substrate. The sensing element is in electrical contact with and between the end electrodes, wherein the sensing element provides a conduction path between the end electrodes, and wherein the sensing element includes an oxygen-deficient metal oxide layer. The oxygen-deficient metal oxide layer can be surface-modified using suitable surface modifiers and synthetic binding entities or biomolecules to form sites capable of selectively binding target biomarkers or bioanalytes for detection purposes.

[0064] The following is a description of each component of the non-invasive conductivity sensor.

[0065] substrate

[0066] The substrate can be made of a material selected from silicon wafers, polymers, glass, or ceramics.

[0067] For example, suitable polymers for use as substrates can be selected from polydimethylsiloxane (PDMS), polyimide (PI), and polyethylene naphthalate (PEN). Suitable ceramics can be selected from alumina (Al2O3), sapphire, and silicon nitride (Si3N4).

[0068] Here, for the purpose of describing the steps involved in manufacturing a non-invasive conductivity sensor, and as follows Figure 1 As shown in step (1), the substrate is a rigid silicon wafer with a SiO2 surface.

[0069] However, those skilled in the art will understand that if the desired objective is to provide a non-invasive conductivity sensor that can be used as a device in applications requiring portability and flexibility, the substrate used is ideally a flexible polymer (such as polyimide foil) rather than the rigid SiO2 / Si wafer described above. The steps for fabricating a flexible non-invasive conductivity sensor using polyimide foil are the same as described above (see...). Figure 1 ).

[0070] Sensing element

[0071] In its simplest form, the sensing element comprises an oxygen-deficient metal oxide layer and one or more bioanalyte binding sites, which are attached to the surface of the oxygen-deficient metal oxide layer by chemical or physical adsorption.

[0072] Metal oxide layers can be formed using any suitable metal oxide selected from the following: zinc oxide (ZnO), strontium titanium oxide (STO), tin oxide, and titanium dioxide.

[0073] 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 will be described below, the inventors have found that good results can be obtained when the metal oxide layer is a thin film oxygen-deficient zinc oxide (ZnO) layer.

[0074] Oxygen-deficient metal oxide layers can be applied to the substrate surface using techniques selected from: 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).

[0075] In a preferred embodiment, such as Figure 1 As shown in step (2), an oxygen-deficient metal oxide layer is applied to the surface of a rigid (SiO2 / Si) wafer or a flexible polyimide foil by reactive sputtering to provide a metal oxide film with a thickness ranging from about 50 nm to about 200 μm.

[0076] For example, such as Figure 1 As shown in step (2), zinc oxide has been sputtered onto the surface of a rigid (SiO2 / Si) wafer to provide an oxygen-deficient zinc oxide layer (ZnO). 1-x The oxygen-deficient zinc oxide layer has multiple hydroxyl (OH) groups on its surface. The deposited oxygen-deficient ZnO layer can be of any suitable thickness to suit the desired application.

[0077] Good results were obtained when the thickness of the oxygen-deficient ZnO layer fell within the range of about 10 nm to about 1 μm.

[0078] Metal oxide thin films

[0079] Different binary (ZnO) x and ZnO y ) and composite metal oxide (such as SrTiO3) thin films have been engineered for non-invasive conductivity sensor applications to sense various bioanalytes, such as IL-6 and CRP.

[0080] The following sections discuss the synthesis process and chemical composition of these binary and composite metal oxide films.

[0081] Zinc oxide (ZnO)

[0082] Two different types of ZnO thin films with different oxygen content ratios were prepared by magnetron sputtering. This resulted in different stoichiometry of the sputtered films. The sputtering parameters and associated conductivity are listed in Table 1.

[0083] Table 1. Parameters used for sputtering ZnO thin films with different stoichiometry and associated conductivity.

[0084]

[0085] Sputtering parameters were selected to engineer thin films with conductivity in the range of 0.08–0.6 S / m. This conductivity range allows the sensor to achieve maximum sensitivity.

[0086] The stoichiometry of sputtered ZnO thin films was evaluated using X-ray photoelectron spectroscopy (XPS). Figure 4 Showing from ZnO x and ZnO y Nuclear-level Zn 2p and O1s spectra collected by thin film.

[0087] like Figure 4 As shown, the O 1s spectrum is fitted with three different peaks associated with Zn-O bonding (represented by peak (1)), oxygen vacancies (Ov, represented by (2)) and -OH bonding (labeled as (3)). [1,2]

[0088] The fitting parameters are listed in Table 2. A relative comparison of peak (2) in the two types of films indicates that ZnO... x Compared to ZnO y It is relatively more oxygen-deficient.

[0089] Table 2. From ZnO x and ZnO y Fitting parameters for nuclear-level XPS spectra collected from thin films.

[0090]

[0091] Strontium titanium oxide (STO)

[0092] Two different types of strontium titanium oxide (SrTiO3:STO) films with different oxygen content ratios were prepared by magnetron sputtering. The sputtering parameters are summarized in Table 3.

[0093] Table 3. Parameters used for sputtering STO thin films with different stoichiometry and associated conductivity.

[0094]

[0095] Figure 5XPS nuclear-level binding spectroscopy spectra of all three elements in two types of STO thin films sputtered in reducing (0% oxygen) and oxidizing (5% oxygen) environments, respectively, are shown. x and STO y In the two STO oxides, Sr 3d( Figure 5 a, Figure 5 The nuclear-level spectra of Sr3d were fitted with single-component data, and no significant shift in chemical state was observed. For both oxides, Sr3d... 5 / 2 At 132.9 eV (±0.1 eV) and Sr 3d 3 / 2 The binding energy of 134.7 eV (±0.1 eV) is attributed to Sr in STO. 2+ substance. [3,4]

[0096] Analysis of the deconvolution spectrum of the nuclear-level binding energy of Ti 2p is shown in... Figure 5 b, Figure 5 In the e. STO x and STO y The two Ti 2p spectra in the image are both obtained using two different components (i.e., Ti). 4+ and Ti (4-x)+ Fitted spectrum. Only Ti exists. 4+ The composition corresponds to a fully stoichiometric STO oxide, while Ti (4-x)+ Substances (such as Ti) 3+ and Ti 2+ The ) indicates the presence of oxygen vacancies in the oxide system. However, at lower binding energies, Ti... 3+ and Ti 2+ The composition is fitted using only one component and denoted as Ti. (4-x)+ To avoid any ambiguity. In STO x and STO y In oxides, Ti 2p 3 / 2 The peak with a binding energy of 458.4 eV is attributed to Ti. 4+ The oxidation state, while the peaks at 456.2 eV and 456.5 eV are attributed to Ti. (4-x)+ substance [5-8] Ti is calculated by integrating the fitted peak. 4+ and Ti (4-x)+ The relative ratio of substances. Calculate STO. x Ti 4+ and Ti (4-x)+ The relative percentages were 72.9% and 27.1%, respectively.

[0097] On the other hand, calculating STO y Ti 4+ and Ti (4-x)+The relative proportions of the substances were 75.2% and 24.8%, respectively. This indicates that STO x Ti (4-x)+ The concentration of the substance is relatively higher than that of STO. y Therefore, STO x Thin film compared to STO y The thin film is more prone to oxygen defects.

[0098] In addition, the O 1s spectrum ( Figure 5 c Figure 5 f) was fitted with two components with peak positions at 529.5 eV and 531.3 eV, corresponding to O in STO oxide. 2- ion [3] and CO bond [3,4,9] These are formed due to the accidental adsorption of carbon onto the surface.

[0099] electrode

[0100] like Figure 1 As shown in step (3), a pair of gold-terminal electrodes are formed on the surface of the deposited oxygen-deficient metal oxide film in a mutually spaced and opposite relationship, and are electrically contacted with the oxygen-deficient metal oxide film to serve as the sensing element of the sensor.

[0101] Briefly, the terminal electrode is formed by evaporating a 250 nm gold film with a 100 nm chromium adhesion layer on top of an oxygen-defective metal oxide layer using electron beam lithography. The deposited gold film is then patterned using standard photolithography and wet etching techniques to define the terminal electrode.

[0102] Intermediate layer

[0103] In a preferred embodiment, the bioanalyte binding site is anchored to the oxygen-deficient metal oxide layer by using an intermediate layer formed by multiple long-chain molecules that have been physically or chemically adsorbed onto the oxygen-deficient metal oxide layer as surface modifiers.

[0104] In one embodiment, an intermediate layer is produced by silanizing the hydroxyl groups of an oxygen-deficient zinc oxide layer with a silanizing agent having terminal functional groups selected from epoxy, thiol, amino, carboxyl, and hydroxyl groups.

[0105] For example, such as Figure 1 As shown in step (4), the silanizing agent is an epoxy-terminated silanizing agent (3-glycidoxypropyl)trimethoxysilane (GPS). The method for adsorbing GPS onto an oxygen-deficient zinc oxide layer is described in the Materials and Methods section below.

[0106] In other embodiments, the surface modifier may be selected from (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS).

[0107] Analyte binding sites

[0108] like Figure 1 As shown in step (5), the now silanized surface of the oxygen-deficient zinc oxide layer is further modified by immobilizing suitable binding entities or biomolecules to the end of each anchored silanizing agent as binding sites for selectively binding desired bioanalytes from biological samples.

[0109] A range of biomolecules can be used as binding sites for the selective binding of desired bioanalytes from biological samples.

[0110] For example, such biomolecules may include proteins, peptides, lipopeptides, protein-bound carbohydrates, or protein-bound ligands.

[0111] In one implementation, the biomolecule is a capture protein.

[0112] Suitablely, the capture protein is a protein-binding scaffold, a T-cell receptor, a TCR-binding fragment, a variable lymphocyte receptor, an antibody, and / or an antibody-binding fragment.

[0113] Protein-bound scaffold

[0114] Protein-binding scaffolds have emerged as viable molecules for binding to a variety of bioanalytes, including proteins. These scaffolds typically contain a stable protein structure (scaffold) that can withstand amino acid modifications within a designated binding region without altering the relative arrangement of the binding domains. These protein-binding scaffolds include (but are not limited to): Adnectin, Affilin (Nanofitin), affinity molecules, Affimer molecules, Affitin, Alphabody, Aptamer, Anticalins, armadillo repeat-sequence-based scaffolds, Avimer, designed ankylosing repeat-sequence protein (DARPin), Fynomer, inhibitory cystine knot (ICK) scaffolds, Kunitz domain peptides, and Monobody (AdNectins). TM ) and Nanofitin.

[0115] Affilins are artificially produced proteins of approximately 20 kDa. They consist of a scaffold structurally associated with human ubiquitin and vertebrate γ-β lens proteins, with eight operable amino acids exposed on their surface. Affilins can be engineered to specifically bind to target bioanalytes and can be used using techniques such as site-directed mutagenesis and phage display libraries.

[10] This makes it specifically adapted to bind with a variety of molecules.

[0116] The affinity protein is approximately 6 kDa and contains a protein scaffold of the IgG isotype antibody Z domain, modified with one or more of the 13 amino acid residues located in its two α-helical binding domains.

[11]

[0117] Affimer molecules are proteins of approximately 12 to 14 kDa that utilize a protein scaffold derived from the cysteine ​​protease inhibitor family of cystatin. Affimer molecules contain two peptide loop regions and an N-terminal sequence adapted for target-specific binding. Affimer molecules, which have 10 kDa at their binding sites, can be generated using phage display libraries and appropriate techniques. 10 A combination of amino acids.

[12]

[0118] Affitins are 66 amino acid residues (approximately 7 kDa) proteins that use a protein scaffold derived from the DNA-binding protein Sac7d, found in *Sulfolobus acidocaldarius*. They are readily produced in vitro from prokaryotic cell cultures and contain 14 binding amino acid residues, which can be mutated to produce more than 3 × 10⁻⁶ affitins. 12 A structural variant.

[13] Screening techniques (such as surface plasmon resonance) can be used to identify the specific binding of these molecules.

[0119] Alphabody molecules are approximately 10 kDa. Unlike most macromolecules, they can penetrate cell membranes (when not immobilized), allowing them to bind to both intracellular and extracellular molecules. The scaffold of an alphabody is a computationally designed coiled-coil structure with three α-helices (A, B, and C), unlike natural structures. Amino acids on the A and C α-helices can be modified to target specific antigens.

[14]

[0120] Aptamers used for protein binding include a range of nucleic acids (DNA, RNA, and XNA) and peptides, which can be screened for binding to specific target molecules. Nucleic Acid Aptamer Database

[15] This allows for the selection of DNA aptamers for in vitro identification. Peptide aptamers consist of short amino acid sequences, typically embedded in circular structures (“loops on the framework”) within a stable protein scaffold framework. Typically, peptide loops of 5 to 20 residues are a variable source for selectively binding to target molecules. Combinatorial libraries and techniques (such as yeast two-hybrid screening) can be used to generate and screen peptide aptamers. Other techniques for generating and screening protein aptamers are described in the literature.

[16]

[0121] Anicalin is a protein-binding molecule derived from lipid transport proteins. Typically, anticalin binds to molecules smaller than antibodies. Methods for screening and developing anticalins are described in the literature. [17,18]

[0122] The armadillo repetitive sequence-based scaffold is characterized by an armadillo domain consisting of approximately 42 tandem armadillo repeat sequences forming supercoils of repeat units, each consisting of three α-helices. Modification of residues within the conserved binding domain allows for the preparation of a series of combinatorial libraries that can be used to select target-specific binders.

[19]

[0123] Avimers (also known as affinity multimers, macrobodies, or low-density lipoprotein receptor (LDLR) domain A) comprise at least two linked peptides of 30 to 35 amino acids in length, based on the A domain of a series of cysteine-rich cell surface receptor proteins. Modification of the A domain allows for targeted binding to a series of epitopes on or across the same target; the number of linking peptides determines the number of potential targets for each avimer. A range of avimer phage display libraries are known in the art, including commercial libraries such as those from Creative Biolabs.

[0124] Designed ankyrin repeat protein (DARPin) is an engineered binding protein derived from ankyrin. Methods for screening and identifying DARPin are described in the literature. [20,21]

[0125] Inhibitor cystine knots (ICK) scaffolds are a family of miniproteins (30 to 50 amino acid residues long) that form stable three-dimensional structures containing three disulfide bridges connecting a series of highly sequence-variable rings. Inhibitor cystine knots include three family members: knottins, cyclotides, and growth factor cysteine ​​knots. Databases known in the art, such as the KNOTTIN database (www.dsimb.inserm.fr / KNOTTIN / ), disclose specific properties of known knottins and cyclotides, such as their sequences, structures, and functions. Furthermore, methods for generating ICKs and screening for binding are described in the literature.

[22]

[0126] Monobody (also known as the trade name AdNectins) utilizes an FN3 (fibronectin type III domain) scaffold with multiple and operable variable groups. AdNectins share the antibody's variable domain and β-sheet loop with the antibody. The binding affinity of monobody can be diversified and customized through in vitro evolutionary methods such as mRNA display, phage display, and yeast display. Methods for screening and generating monobody are described in the literature. [23,24]

[0127] Antibodies and antibody fragments

[0128] In some implementations, the bioanalyte binding site is an antibody or a binding fragment thereof. Antibodies are protein-binding molecules with exemplary diversity, potentially containing up to 10 in a single individual. 11 Up to 10 12 Each unique molecule allows for further diversity due to genetic variations between individuals. In vivo antibody diversity is driven by random recombination of a range of genes in the V(D)J linker.

[0129] Antibody binding is primarily determined by three hypervariable regions on the heavy and light chains, referred to as complementarity-determining regions (CDRs) 1, 2, and 3. Therefore, each mature antibody possesses six CDRs (variable heavy (VH) chain CDRs 1, 2, and 3, and variable light (VL) chain CDRs 1, 2, and 3). These hypervariable regions form a three-dimensional antigen-binding pocket, and the antibody's binding specificity is determined by a specific amino acid sequence within the CDRs (primarily CDR 3).

[0130] Antibodies against specific bioanalytes can be commercially obtained or generated using methods known in the art. For example, antibodies against specific bioanalytes can be prepared using methods generally disclosed in the literature.

[25]

[0131] The specificity, affinity, and affinity of antibodies generated in a subject's body can be altered through in vitro processes (such as affinity maturation).

[26] Therefore, antibodies derived from the body can be further modified to produce different but lineage-related antibodies. Thus, the term "antibody" encompasses both in vivo antibodies and in vitro molecules that have undergone a mutation process to modify the CDR binding site, giving them a unique sequence compared to in vivo-generated antibodies.

[0132] The term antibody also includes unconventional antibodies generated by species such as camels, sharks, and jawed fish. Therefore, the term antibody includes heavy chain antibodies, including camel antibodies, IgNARs, and variable lymphocyte receptors (VLRs). Furthermore, these antibodies can be fragmented into their binding moieties (such as VNARs—a single binding moiety of IgNARs) or recombinantly integrated into fusion proteins. Methods for generating and modulating such unconventional antibodies are described in the literature. [27,28]

[0133] Antibody-binding fragments

[0134] In some implementations, the bioanalyte binding site is an antibody-binding fragment. This antibody-binding fragment can be derived from an antibody or can be recombinantly generated, and it has the same sequence as the CDR (Combined Derivative Record) of the antibody or antibody fragment. In fact, these CDRs may originate from affinity-matured antibodies and therefore may differ from antibodies derived from within the body.

[0135] Antibodies consist of four chains (two heavy chains and two light chains), which can be divided into Fc (crystallizable moiety) and Fab (part of the antibody) domains. The Fc moiety of an antibody interacts with Fc receptors and the complement system. Therefore, the Fc moiety is important for the immune function of the antibody. However, the Fab moiety contains the antibody-binding region, which is crucial for the specificity of the antibody for the desired epitope.

[0136] Therefore, in some embodiments, the bioanalyte binding site is a Fab fragment of the antibody. The Fab fragment can be a single Fab fragment (i.e., an antibody fragment generated in the absence of a disulfide bridge) or an F(ab')2 fragment containing two Fab fragments of the antibody linked by a disulfide bridge. These fragments are typically generated by fragmenting the antibody using a digestive enzyme (such as pepsin). Methods are described in the literature.

[29]

[0137] Each Fab fragment of an antibody has a total of six CDRs, with the VH and VL chains each containing three CDRs (within a framework consisting of four frame regions). The constant regions of the Fab fragment can be removed to leave only the VH and VL regions of the antibody. Individual VH and VL chains (each containing only three CDRs) have been shown to bind specifically with high affinity. Typically, these individual binding regions are called single-antibody domains (sdAbs). Alternatively, the VH and VL chains can be linked via linkers to form fusion proteins called single-chain variable fragments (scFvs—also known as biantibodies). Unlike Fabs, scFvs are not fragmented from antibodies but are typically formed based on recombination of the antibody's CDRs and frame regions. Furthermore, sdAbs can be recombined to generate binding components that form larger fusion proteins, which may also include portions that can function to: stabilize the binding region, improve or facilitate anchoring to sensing elements or intermediate layers, or improve binding (e.g., by providing flexibility to the binding region or optimizing the length of the bioanalyte binding site, thereby allowing access to the antigenic region of the bioanalyte). Therefore, in some implementations, the bioanalyte binding site is or includes scFv or sdAb. scFv may include multiple VH and VL chains linked together to form a multivalent scFv (such as di-scFv or tri-scFv).

[0138] Antibodies and antibody fragments, or fusion proteins containing antibody-derived sequences, can be commercially obtained or generated by methods known in the art (such as those discussed above).

[0139] Proteins and receptors

[0140] Protein receptors or ligands that interact with and bind to proteins can serve as binding sites for bioanalytes. Such receptors and ligands include the entire receptor or ligand, or specific fragments thereof (e.g., fragments containing the binding domain of the receptor or ligand). Specifically envisioned receptors include cytokine receptors, where cytokines (such as interleukins or chemokines) can provide information about the state of the immune system. In some embodiments, receptors or ligands (or fragments thereof) can be integrated to form fusion proteins.

[0141] For example, interleukin-6 (IL-6) is an inflammatory pluripotent cytokine and an important biomarker for monitoring immune responses during cancer treatment. It can also be used to monitor psychological stress and insulin activity.

[0142] For example, the inventors have achieved good results when using anti-interleukin-6 (IL-6) antibodies to selectively recognize and bind to IL-6.

[0143] For example, the inventors have achieved good results when using anti-C-reactive protein (CRP) antibodies to selectively recognize and bind to CRP.

[0144] Summarize

[0145] In summary, the aforementioned non-invasive conductivity sensor is a passive electronic device configured with a simple in-plane geometry of two end electrodes. The sensing element of the sensor is in the form of an oxygen-deficient metal oxide thin film, which is applied to the surface of a sensor substrate and subsequently functionalized with specific bioanalyte binding sites. These binding sites are selective for one or more analytes in bodily fluids (such as human saliva and / or sweat). When a voltage is applied through the sensor, an electrical signal is generated that is proportional to the change in conductivity of the sensing element. This change is due to charge transfer between the bioanalyte and the complex formed by the bioanalyte binding sites and the oxygen-deficient metal oxide thin film layer. This electrical signal can be equivalent to the level of the target biomarker or bioanalyte present in the bodily fluid.

[0146] The method for detecting the level of target bioanalytes in body fluids using the aforementioned non-invasive conductivity sensor is now described below.

[0147] Detection methods

[0148] According to another preferred embodiment of the present invention, a method for detecting bioanalytes is provided.

[0149] Briefly, a method for detecting bioanalytes using an oxygen-deficient metal oxide-based sensor includes the following steps: (i) contacting an oxygen-deficient metal oxide-based sensing element with a bodily fluid sample solution containing the bioanalyte; (ii) applying a voltage through the sensor; and (iii) detecting an electrical signal generated between a pair of terminal electrodes using a current source meter, the electrical signal being proportional to a change in conductivity corresponding to the detection of the bioanalyte when the bioanalyte binds to a bioanalyte binding site on the surface of the oxygen-deficient metal oxide sensing element.

[0150] Example

[0151] Antigen concentration-dependent studies:

[0152] Both IL-6 and CRP antigens showed concentration-dependent resistance changes relative to the device's baseline resistance. The baseline resistance of the GPS silanized sensor with immobilized antibodies was measured before antigen addition. For both IL-6 and CRP antigens, a linear correlation was observed between the resistance change and the antigen concentration. Figure 2The response rates (i.e., the slopes of the curves) for IL-6 and CRP were 5.1% and 4.1% / M, respectively. These values ​​indicate that the ZnO sensor immobilized with anti-IL-6 antibody exhibits higher sensitivity for detecting IL-6 antigen compared to the ZnO conductivity sensor immobilized with anti-CRP antibody. To determine the contribution of the antigen solution matrix, the resistance of the PBS solution was measured on both types of ZnO sensors with immobilized antibodies. For both types of sensors with immobilized antibodies, the resistance change in the presence of PBS solution was less than 1%. Therefore, the contribution of the antigen solution matrix to the resistance change is negligible.

[0153] For both antigens, the non-invasive conductivity ZnO sensor exhibited a detectable response even at concentrations lower than those in healthy human body fluids. The reported concentration of IL-6 in the sweat of healthy individuals is approximately 0.38 pM (10 ng / L).

[30] The concentration of CRP in the saliva of healthy individuals is approximately 12 pM (285 ng / L).

[31] The ZnO conductivity sensor showed a 9.2% resistance change for an IL-6 concentration of 4 fM, which is more than 100 times lower than the IL-6 concentration in healthy human sweat. Similarly, the sensor detected a resistance change of approximately 12.5% ​​for the lowest CRP concentration (13 fM), which is almost 1000 times lower than the CRP concentration in healthy human saliva. This high sensitivity to concentrations far below those found in healthy human body fluids clearly demonstrates the importance of ZnO-based conductivity sensors in detecting bioanalytes in human body fluids.

[0154] Cross-selective study 1:

[0155] To determine the viability of each immobilized antibody device in the presence of other antigens, cross-selectivity studies were conducted. Devices immobilized with anti-IL-6 antibody showed a 3% change in electrical resistance in the presence of 13 pM CRP, while devices immobilized with CRP antibody showed a 3.5% change in electrical resistance in the presence of 4 pM IL-6. Figure 3The antigen concentrations used in this experiment were chosen to be as close as possible to those found in healthy human body fluids. When the two antigens were mixed before being added to the device, the device immobilized with anti-IL-6 antibody showed a 17.6% change in resistance, approximately 4% lower than that of 4 pM IL-6 on the same device. In contrast, the device immobilized with anti-CRP antibody showed a 27.4% change in resistance, approximately 3% higher than that of 13 pM CRP on the same device. The cross-selectivity test results clearly indicate that the ZnO device immobilized with anti-IL-6 antibody is selective for IL-6 in the presence of CRP antigen, while the ZnO device immobilized with anti-CRP antibody is selective for CRP antigen in the presence of IL-6 antigen. The reasons for the 4% decrease in change in the IL-6 antibody device and the 3% increase in change in the CRP antibody device in the presence of the antigen mixture are currently unclear.

[0156] Cross-selective study 2:

[0157] To determine the feasibility of each immobilized antibody device in the presence of other antigens, cross-selectivity studies were performed. Cathelicidin, β-diuretic natriuretic peptide (BNP), and cardiac troponin I (cTnI) were used in this study along with IL-6 and CRP antigens. The nominal concentrations of the antigens in the original antigens and mixtures were maintained at 4 pM. Significantly high electrical resistance changes were observed against the target antigens when the antigen mixtures reacted with the devices of the corresponding immobilized antibodies. Figure 6 When IL-6 in the antigen mixture was treated with a device immobilized with IL-6 antibody, a 31% change in resistance was observed. For CRP, only a 3% change in resistance was observed. Furthermore, the device immobilized with CRP antibody showed a 30% change in resistance for CRP in the antigen mixture, while IL-6 contributed 5%. Notably, the krait antimicrobial peptide, BNP, and cTnI did not contribute to the resistance changes of the device immobilized with IL-6 and CRP antibodies. The small contributions of IL-6 and CRP antigens to the resistance changes of their non-corresponding antibodies may be due to the physical adsorption of these molecules onto the antibodies. The negative impact of the krait antimicrobial peptide, BNP, and cTnI on the resistance changes suggests that either these three antigens did not bind to the antibodies under study, or the charge transfer effects of these antigens employed a different mechanism than those of the CRP and IL-6 antigens on the antibodies.

[0158] Materials and methods

[0159] Sensor manufacturing:

[0160] A sensor is fabricated by depositing a 100 nm thick metal oxide (e.g., zinc oxide (ZnO)) thin film on a rigid (SiO2 / Si) or flexible plastic (polyimide foil) substrate, which serves as the sensing element in the sensor. The composition of the sensing element is modified by reactive sputtering engineering to achieve a conductivity of 0.08–2 Siemens / m. 2 Within the range, more preferably 0.08-0.6 Siemens / m 2 Oxygen-deficient metal oxide films within the range. For conductivity measurements, two in-plane terminal electrodes are patterned, with a sensing area of ​​16 × 10⁻⁶. -6 m 2 The changes in conductivity corresponding to the dispensed antibodies and antigens were measured using a commercial current source meter (Keysight Technologies' B2901A precision source / measurement unit).

[0161] Preparation of antibody and antigen solutions:

[0162] Interleukin-6 (IL-6), anti-IL-6, C-reactive protein (CRP), and anti-CRP were purchased from a commercial supplier (Sigma-Aldrich) and used as received. Cobra antimicrobial peptide, β-diuretic natriuretic peptide (BNP), and cardiac troponin I (cTnI) were obtained from multiple commercial suppliers (Abcam, MyBioSource, and ProSpec Bio) and used as received.

[0163] In phosphate buffer (PBS, pH 7.4), 1:10 6 The anti-IL-6 stock solution was diluted to immobilize it on the surface of an oxygen-deficient ZnO sensor. Prior to immobilization, the received undiluted anti-CRP solution was diluted 1:50 in PBS (pH 7.4). The received undiluted IL-6 powder was completely dissolved in a known volume of autoclaved Milli-Q water and diluted in PBS solution at 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. A standard series of CRP solutions was also prepared by diluting the received undiluted CRP solution in PBS at pH 7.4 with predetermined volumes. The prepared CRP concentrations were 13 nM, 100 pM, 13 pM, 100 fM, and 13 fM. Solutions of other antigens were prepared in a similar manner.

[0164] Fabrication of GPS silanized sensors:

[0165] Previously, the use of (3-glycidoxypropyl)trimethoxysilane (GPS) to silanize the surface of oxygen-deficient metal oxide sensors has been reported, although it was for invasive sensors.

[32] In this paper, a slight modification was made to the reported silanization process for ZnO. Briefly, the newly prepared ZnO device was exposed to O2 plasma (Plasma Cleaner PDC-002, Harrick Plasma) for 10 minutes to activate the hydroxyl groups on the ZnO surface. Then, 20 μL of a freshly prepared GPS solution was drop-coated onto an Al foil placed in a vacuum dryer to allow GPS vapor to accumulate. The O2 plasma-cleaned ZnO sensor was then exposed to this GPS vapor for 30–45 minutes. The plasma-cleaned ZnO sensor was then exposed to GPS vapor within an LC200 glovebox system. After GPS vapor exposure, the ZnO sensor was thoroughly rinsed with Milli-Q water for 2 minutes to remove any unbonded silane groups from the ZnO sensor surface. The washed ZnO sensor was then heated at 150 °C for 10 minutes to enhance the bonding of the silane groups to the ZnO surface. These GPS silanized sensors are then used for antibody immobilization.

[0166] It will be understood that the surface of oxygen-deficient strontium titanium oxide (STO) sensors can be silanized using (3-glycidoxypropyl)trimethoxysilane (GPS) or an alternative silanizing agent in the same manner as described above.

[0167] Antibody immobilization:

[0168] Antibodies (IgG) have been immobilized on GPS silanized ZnO sensors, although for invasive sensors.

[32] 25 μL of freshly prepared 1:10 6 Diluted anti-IL-6 solution was drop-coated onto the surface of each new GPS silanized ZnO sensor and incubated for 2 hours to allow the IL-6 antibody to immobilize on the ZnO sensor surface. The sensors were then washed with pH 7.4 PBS solution to remove any unbound antibody. The PBS-washed ZnO sensors were then dried in an N2 stream. These sensors immobilized with anti-IL-6 antibody were used for IL-6 antigen concentration and cross-selectivity measurements. GPS silanized ZnO sensors immobilized with CRP were prepared using 25 μL of freshly prepared 1:50 diluted anti-CRP solution following the same procedure.

[0169] Antigen conductivity method

[0170] The baseline conductivity of the ZnO sensor with immobilized antibody was measured before the addition of antigen. A 15 μL volume of antigen solution was dropped onto the surface of the ZnO sensor with immobilized antibody and incubated for 10 minutes. After this time, any remaining antigen solution was removed from the sensor, and the surface was dried under a stream of N2. Conductivity measurements were then performed for each concentration. For cross-selectivity measurements, 15 μL of 4 pM IL-6 solution and 15 μL of 13 pM CRP solution were dropped onto the ZnO device with immobilized anti-IL-6 antibody. The nominal concentrations of IL-6 and CRP antigen in the premixed solutions used for cross-selectivity measurements were 4 pM and 13 pM, respectively.

[0171] Study on the shelf life of the device

[0172] The GPS silanization device is stable for at least 15 months. This conclusion is based on the changes in the electrical resistance of IL-6 and CRP antigen obtained from the aged device over 15 months (450 days). Figure 7 The resistance changes of both types of antigens were linear with their concentrations. Similar linear trends were observed for both types of antigens on the newly prepared device (day 0). The close overlap of the standard errors of the resistance changes for a given concentration of the selected antigen strongly suggests that there is no statistically significant difference between the mean resistance changes of the new and old devices for a given concentration of the selected antigen.

[0173] Performance of the device in artificial saliva

[0174] The device immobilizing IL-6 and CRP antibodies successfully detected the corresponding antigens in artificial saliva. Figure 8 The resistivity changes of both types of antigens were linear with their concentrations. The responsiveness (i.e., slope) of IL-6 and CRP in artificial saliva was 1.6 (% / M) and 1.2 (% / M), respectively, indicating that the device was more sensitive to IL-6 in artificial saliva than to CRP. The resistivity changes of these antigens at a given concentration were significantly lower in artificial saliva compared to PBS. This significantly lower resistivity change was attributed to the large background contribution of artificial saliva (83% for IL-6 and 85% for CRP). This is likely due to the higher charge transfer effect resulting from the higher ionic composition of artificial saliva compared to PBS. Similar to the effect in PBS, these devices exhibited significant selectivity for the target antigens in artificial saliva when the corresponding antibodies were immobilized. For example, in the presence of a device immobilized with IL-6 antibody, the resistivity change of IL-6 in antigen mixtures prepared in artificial saliva was significantly higher than that of other antigens. Similarly, devices immobilized with CRP antibody exhibited significantly higher resistivity changes for CRP antigens in antigen mixtures prepared in artificial saliva.

[0175] advantage

[0176] From the foregoing discussion, those skilled in the art will understand that the non-invasive conductivity sensors based on oxygen-deficient metal oxides of various embodiments of the present invention offer many advantages over their existing counterparts.

[0177] In fact, conductivity sensors based on oxygen-deficient metal oxides can measure the concentration of target bioanalytes at levels lower than their corresponding levels in human body fluids. Therefore, the inventors believe that such sensors offer significant potential for developing cost-effective, biocompatible, and functional sensors that can be widely used as personalized and reusable healthcare monitoring devices. Indeed, the inventors broadly anticipate that these sensors will have a significant impact on:

[0178] 1. Cardiovascular disease warning: Routine tests using these sensors are expected to alert to elevated levels of inflammatory biomarkers, allowing for intervention before a heart attack or stroke.

[0179] 2. Cancer Treatment: These sensors are expected to have invaluable significance in monitoring cancer biomarker levels and influencing treatment; and

[0180] 3. Treatment of abdominal disorders: These sensors are also expected to detect abdominal disorders (such as liver disorders) by sensing CRP antigens.

[0181] Conductivity sensors based on oxygen-deficient metal oxides can be integrated with conventional portable integrated electronic devices and wearable electronic devices / devices, making them portable devices that can be worn when necessary.

[0182] In fact, conductivity sensors based on oxygen-deficient metal oxides can be fabricated on any type of insulating and plastic substrate while still retaining selectivity for target biomarkers or bioanalytes. Furthermore, these sensors are reusable, further enhancing their cost-effectiveness.

[0183] References

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[0216] definition

[0217] Whenever a range is given in the specification, such as a temperature range, time range, or concentration range, all intermediate ranges and subranges, as well as all individual values ​​included within a given range, are intended to be included in this disclosure. It will be understood that any subrange or individual value included in the ranges or subranges described herein may be excluded from the claims herein.

[0218] All definitions defined and used herein should be understood to take precedence over dictionary definitions, definitions incorporated by reference in other documents, and / or the usual meaning of the defined terms.

[0219] Unless explicitly stated otherwise, the indefinite article “a” used in this specification should be understood to mean “at least one”.

[0220] As used herein, the phrase “and / or” should be understood to mean “one or both” of the elements so combined, i.e., elements that exist together in some cases and separately in others. The multiple elements listed with “and / or” should be interpreted in the same way, i.e., “one or more” of the elements so combined. In addition to the elements specifically indicated by the “and / or” clause, other elements may optionally exist, whether related to or unrelated to those specifically indicated elements. Thus, as a non-limiting example, when used in conjunction with open-ended language such as “including”, references to “A and / or B” may, in one embodiment, refer only to A (optionally including elements other than B); in another embodiment, refer only to B (optionally including elements other than A); in yet another embodiment, refer to both A and B (optionally including other elements); and so on.

[0221] Although the invention has been described in conjunction with a limited number of embodiments, those skilled in the art will understand that many alternatives, modifications, and variations are possible based on the foregoing description. Therefore, the invention is intended to include all such alternatives, modifications, and variations that may fall within the spirit and scope of the disclosed invention.

[0222] When the terms “comprising,” “including,” or “contains” are used in this specification (including the claims), they should be interpreted as specifying the presence of the stated feature, integer, step, or component, but not excluding the presence of one or more other features, integers, steps, or components, or groups thereof.

Claims

1. A sensor for detecting a bioanalyte, comprising: substrate; A pair of end electrodes are disposed on the substrate in a mutually spaced and opposite relationship; and A non-insulating sensing element is applied to the surface of a substrate, between and in electrical contact with the opposing electrodes, wherein the sensing element provides a conduction path between the electrodes, wherein the sensing element includes an oxygen-deficient metal oxide layer and a bioanalyte binding site, and wherein when a voltage is applied through a sensor, an electrical signal proportional to the change in the conductivity of the sensing element is generated, the electrical signal corresponding to the binding of the bioanalyte to the bioanalyte binding site.

2. The sensor according to claim 1, wherein the oxygen-deficient metal oxide layer is formed of a metal oxide selected from the group consisting of zinc oxide, strontium titanium oxide, tin oxide, and titanium dioxide.

3. The sensor according to claim 1 or 2, wherein the thickness of the oxygen-deficient metal oxide layer falls within the range of 50 nm to 200 μm.

4. The sensor according to claim 1, wherein the oxygen-deficient metal oxide layer is applied to the substrate surface by means of a technique selected from: reactive sputtering, physical vapor deposition, chemical vapor deposition, metal-organic chemical vapor deposition, pulsed laser deposition, and molecular beam epitaxy.

5. The sensor of claim 1, wherein the bioanalyte binding site is anchored to the oxygen-deficient metal oxide layer by physical or chemical adsorption to an intermediate layer of the oxygen-deficient metal oxide layer.

6. The sensor according to claim 5, wherein the intermediate layer is produced by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functional group selected from epoxy, thiol, amino, carboxyl, and hydroxyl groups.

7. The sensor according to claim 6, wherein the silanizing agent is selected from (3-glycidoxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane, (3-aminopropyl)triethoxysilane, and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane.

8. The sensor of claim 1, wherein the conductivity of the oxygen-deficient metal oxide layer falls within 0.08 Siemens / m. 2 Up to 0.6 Siemens / m 2 Within the range.

9. The sensor according to claim 1, wherein the bioanalyte binding site is a biomolecule.

10. The sensor according to claim 9, wherein the biomolecule is a protein, peptide, lipopeptide, or protein-bound carbohydrate.

11. The sensor of claim 9, wherein the biomolecule is a capture protein.

12. The sensor of claim 11, wherein the capture protein is a protein-binding scaffold, a T-cell receptor, a TCR-binding fragment, a variable lymphocyte receptor, an antibody, and / or an antibody-binding fragment.

13. The sensor of claim 12, wherein the protein binding scaffold is selected from: Adnectin, Affilin, affinity, Affimer molecule, Affitin, Alphabody, aptamer, Anticalins, armadillo repeat sequence-based scaffold, Atrimer, Avimer, designed ankylosing repeat sequence protein, Fynomer, inhibitory cystine knot scaffold, Kunitz domain peptide, Monobody, and Nanofitin.

14. The sensor of claim 12, wherein the antibody binding fragment comprises Fab, (Fab')2, Fab', single-chain variable fragments, di-scFv and tri-scFv, single-domain antibodies, biantibodies, or fusion proteins containing antibody-binding domains.

15. The sensor of claim 1, wherein the bioanalyte binding site binds interleukin-6.

16. The sensor of claim 1, wherein the bioanalyte binding site binds to C-reactive protein.

17. The sensor of claim 1, wherein the substrate is made of a material selected from silicon wafers, polymers, glass, and ceramics.

18. The sensor of claim 17, wherein the polymer is selected from polydimethylsiloxane, polyimide, and polyethylene naphthalate.

19. The sensor of claim 17, wherein the ceramic is selected from alumina, sapphire, and silicon nitride.

20. A method for detecting a bioanalyte, the method comprising the following steps: a) Contact the sensing element of the sensor according to any one of claims 1 to 19 with a sample solution containing a bioanalyte; b) Apply voltage via a sensor; and c) Detect the generated electrical signal that is proportional to the change in conductivity, the change corresponding to the detection of the bioanalyte when it binds to the bioanalyte binding site.

21. The method of claim 20, wherein the bioanalyte binding site is a biomolecule.

22. The method of claim 20 or 21, wherein the bioanalyte binding site binds interleukin-6.

23. The method of claim 22, wherein the change in conductivity detected in a sample solution with an IL-6 concentration of 4 femtomoles is 9.2%.

24. The method of claim 20 or 21, wherein the bioanalyte binding site binds to C-reactive protein.

25. The method of claim 24, wherein the change in conductivity detected in the sample solution with a CRP concentration of 13 femtomolar is 12.5%.

26. A method for manufacturing a sensor for detecting a bioanalyte, the method comprising the steps of: Provide substrate; A pair of end electrodes are deposited on the substrate in a spaced-apart and opposite relationship; and A non-insulating sensing element is applied between and in electrical contact with the opposing electrodes. The sensing element is in the form of an oxygen-deficient metal oxide layer coated with bioanalyte binding sites. The sensing element provides a conduction path between the electrodes, wherein the bioanalyte binding sites are selective for detecting the bioanalyte when the bioanalyte binds to them.

27. The method of claim 26, wherein the oxygen-deficient metal oxide layer is formed of a metal oxide selected from zinc oxide, strontium titanium oxide, tin oxide, and titanium dioxide.

28. The method of claim 26, wherein the thickness of the oxygen-deficient metal oxide layer falls within the range of 50 nm to 200 μm.

29. The method of claim 26, wherein the oxygen-deficient metal oxide layer is applied to the substrate surface by means of a technique selected from: reactive sputtering, physical vapor deposition, chemical vapor deposition, metal-organic chemical vapor deposition, pulsed laser deposition, and molecular beam epitaxy.

30. The method of claim 26, further comprising the following steps: The intermediate layer is physically or chemically adsorbed onto the oxygen-deficient metal oxide layer to anchor the binding sites of bioanalytes to the oxygen-deficient metal oxide layer.

31. The method of claim 30, wherein the intermediate layer is produced by silanizing the oxygen-deficient metal oxide layer with a silanizing agent having a terminal functional group selected from epoxy, thiol, amino, carboxyl, and hydroxyl groups.

32. The method according to claim 31, wherein the silanizing agent is selected from (3-glycidoxypropyl)trimethoxysilane, (3-mercaptopropyl)trimethoxysilane, (3-aminopropyl)triethoxysilane and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane.

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