Conductivity sensor for detecting bioanalytes and detection method thereof

Abstract

The present invention provides a sensor for detecting a bioanalyte, comprising a substrate, a pair of terminal electrodes arranged on the substrate in a spaced-apart opposing relationship, and a sensor element located between the pair of terminal electrodes and in electrical contact with the pair of terminal electrodes, the sensor element comprising: (i) a semiconductor portion of the substrate comprising a high resistivity non-oxide semiconductor, wherein a conductive path between the terminal electrodes passes through the semiconductor portion; and (ii) bioanalyte binding sites on a surface of the semiconductor portion, where binding of the bioanalyte to the bioanalyte binding sites causes a change in the electrical resistance of the sensor.

Description

[Technical field]

[0001] The present invention relates to sensors, and in particular to a conductivity sensor for detecting bioanalytes in fluids and a method for detecting bioanalytes using such a sensor. The present invention has been developed primarily for use in the detection of various bioanalytes in body fluids and will be described below with reference to this exemplary application.

[0002] The following discussion of the background of the invention is intended to facilitate an understanding of the present invention, but it is understood that the discussion is not an affirmation or admission that any of the subject matter referred to was published or publicly known or part of the common general knowledge in Australia or any other country at the priority date of any one of the claims herein. [Background technology]

[0003] Sensors are traditionally used to monitor / measure the levels of target biomarkers (hereafter bioanalytes) in tissues and / or biological fluids. One approach uses invasive sensors where the sensor components come into direct contact with the tissue or bodily fluids, potentially causing infection, tissue damage and discomfort. Another approach relies on the use of non-invasive sensors to measure the levels of bioanalytes in a sample solution that contains a sample of bodily fluid.

[0004] Sensors for such applications rely on a variety of detection techniques, such as optical absorption and electrochemical methods. Optical absorption-based sensors are not very accurate due to the closely overlapping weak absorption bands of the various bioanalytes that may be present in bodily fluids, as well as the temperature sensitivity of such assays.

[0005] On the other hand, electrochemical sensors currently dominate the biosensing field because they are more accurate. Such sensors operate by measuring an electrical signal generated after interaction of the bioanalyte of interest with a sensor element associated with the sensor, the generated electrical signal being proportional to the concentration of the bioanalyte. The interaction of the bioanalyte with the sensor produces a measurable change in the current (amperometric sensor), charge accumulation or potential (potentiometric sensor), conductive properties of the sensor element (conductivity sensor) or impedance of the sensor element (impedance sensor).

[0006] Amperometric and potentiometric sensors that use electrochemical transduction typically require a working electrode, a counter electrode (or auxiliary electrode), and a reference electrode. The reference electrode is maintained away 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 a transduction component when an interaction occurs, while the counter electrode measures the current and facilitates the delivery of electrolyte to enable current transfer to the working electrode.

[0007] In the case of a conductivity sensor, the analyte-sensitive resistance of the sensor element is measured by applying a voltage between two electrodes and measuring the current response through the analyte-sensitive sensor element between the electrodes. Advantageously, such devices therefore do not require a reference electrode. Furthermore, conductivity sensors can be operated with low amplitude AC voltages, thus preventing Faradaic reactions on the electrodes, and also, due to their simple operating principle, can be miniaturized and integrated into various electronic devices. Thus, although conductivity sensors offer certain advantages compared to amperometric and potentiometric sensors, the sensitivity of many conductivity sensors is hindered by the use of polymers as the sensor element, which can result in sensors with low durability and poor long-term stability.

[0008] Co-pending International Application PCT / AU2020 / 051396 discloses conductivity sensors comprising thin film metal oxide based sensor elements. Although these devices offer excellent sensitivity and selectivity in detecting various bioanalytes in body fluids, the device structure is complex and it is desirable to avoid the use of thin film fabrication techniques typically required to prepare metal oxide sensing layers.

[0009] As an alternative to direct conductivity sensors, sensors based on field effect transistors have also been developed. A field effect transistor is a device with three terminals: source, gate and drain. The interaction of the bioanalyte with the sensor element (gate) results in a field effect that changes the conductivity between the source and drain.

[0010] For example, US 2010 / 2016256 describes a biosensor that includes a substrate, a source electrode on the substrate, a drain electrode on the substrate, and at least one functionalized metal oxide nanobelt on the surface of the substrate between the source and drain electrodes, where the functionalized nanobelt has a chemically functionalized surface bound to one or more detection molecules for binding to a biological analyte to be detected, such that a field gating effect occurs upon binding of the analyte to the one or more detection molecules bound to the nanobelt surface. Binding of the analyte modifies the field effect of the nanobelt (gate), thereby modifying the conductivity of the path between the source and drain, allowing the change in conductivity to be monitored.

[0011] This type of device generally suffers from several drawbacks. First, field effect transistors are typically devices that turn on and off and have a nonlinear response. In these devices, there is typically a small region of linear response that then plateaus, so the resistance does not change linearly, meaning that the device is difficult to use over a wide range of conditions. Second, as the skilled artisan will appreciate, for such devices to work as described, it is necessary that an insulating (dielectric) layer is present between the conductive path between the source and drain and the gate bias (the nanobelt in US 2010 / 2016256). Thus, this type of device suffers from the drawback that it is relatively complex to manufacture due to the number of different structural elements, and thus difficult to manufacture on an industrial scale. Third, many field effect sensors use functionalized metal oxides in thin films or other microstructured configurations as sensor elements. Again, it would be desirable to avoid the complexity of the microfabrication techniques required to manufacture such device structures.

[0012] The present invention seeks to provide a sensor for use in the detection of a bioanalyte and a method for detecting a bioanalyte, which will overcome or substantially ameliorate at least some of the shortcomings of the prior art, or at least provide a useful alternative. [Prior art documents] [Patent documents]

[0013] [Patent Document 1] International Application No. PCT / AU2020 / 051396 [Patent Document 2] US Patent Application Publication No. 2010 / 2016256 Summary of the Invention [Means for solving the problem]

[0014] 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 arranged on the substrate in a spaced-apart opposing relationship, and a sensor element located between the pair of terminal electrodes and in electrical contact with the pair of terminal electrodes, wherein the sensor element comprises: (i) a semiconductor portion of the substrate, the semiconductor portion comprising a high resistivity non-oxide semiconductor, wherein a conductive path between the terminal electrodes passes through the semiconductor portion; and (ii) bioanalyte binding sites on a surface of the semiconductor portion, wherein binding of the bioanalyte to the bioanalyte binding sites causes a change in the electrical resistance of the sensor.

[0015] In some embodiments, the non-oxide semiconductor has a resistivity greater than 100 ohm-cm, hi some embodiments, the non-oxide semiconductor has a resistivity in the range of about 500 ohm-cm to about 50,000 ohm-cm, or in the range of about 1000 ohm-cm to about 10000 ohm-cm.

[0016] In some embodiments, the sensor has an electrical resistance in the range of about 10 kilohms to about 10,000 kilohms.

[0017] In some embodiments, the non-oxide semiconductor is selected from the group consisting of elemental semiconductors and compound semiconductors, hi some embodiments, the non-oxide semiconductor is an elemental semiconductor.

[0018] In some embodiments, the non-oxide semiconductor is a silicon semiconductor. The silicon semiconductor may be an intrinsic silicon semiconductor. The silicon semiconductor may be a float-zone type silicon semiconductor.

[0019] In some embodiments, the substrate includes a semiconductor portion as an integral part thereof. The substrate may be a wafer of a non-oxide semiconductor.

[0020] In some embodiments, the bioanalyte binding moiety is chemically bonded to the semiconductor moiety by an organic linker, which may be, for example, a residue of a silanizing agent. The bioanalyte binding moiety may be chemically bonded to the semiconductor layer by a process comprising: (i) silanizing a non-oxide semiconductor 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 carboxy group, and a hydroxy group; and (ii) reacting a precursor comprising the bioanalyte binding moiety with the terminal functional group. The silanizing agent is selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane (GPS), (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS).

[0021] In some embodiments, the bioanalyte binding site is present on a biomolecule or a molecularly imprinted polymer.

[0022] In some embodiments, the bioanalyte binding site is present on a biomolecule selected from the group consisting of a protein, a peptide, a lipopeptide, a protein-binding carbohydrate, and a protein-binding ligand.

[0023] In some embodiments, the biomolecule is a capture protein. The capture protein may be selected from 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.

[0024] Suitable protein binding scaffolds may be selected from the group consisting of adnectins, affilins, affibodies, affimer molecules, affitins, alphabodies, aptamers, anticalins, armadillo repeat protein based scaffolds, atrimers, avimers, designed ankyrin repeat proteins (DARPins), finomers, inhibitory cystine knot (ICK) scaffolds, Kunitz domain peptides, monobodies and / or nanophytins.

[0025] Binding fragments of antibodies may include Fab, (Fab')2, Fab', single chain variable fragments (scFv), di- and tri-scFv, single domain antibodies (sdAb), diabodies, or fusion proteins comprising the binding domain of an antibody.

[0026] In some embodiments, the bioanalyte binding site binds interleukin-6 (IL-6) or C-reactive protein (CRP).

[0027] In some embodiments, the bioanalyte binding site binds to a viral protein.

[0028] The sensor is preferably a conductivity sensor. The sensor may therefore comprise a device for applying a voltage between the terminal electrodes and measuring the current flow through the conductive path of the sensor. This device may preferably be a potentiostat. In an embodiment, the sensor is therefore not a field effect transistor.

[0029] According to a second aspect of the present invention there is provided a method for detecting a bioanalyte comprising the steps of: a) contacting a sensor element of a sensor according to any embodiment of the first aspect with a substance which may contain the bioanalyte; b) measuring an electrochemical parameter of the sensor corresponding to the resistance of the sensor; and c) detecting the presence or absence of the bioanalyte on the sensor element based on the electrochemical parameter measured in step b).

[0030] In some embodiments, measuring an electrochemical parameter of the sensor includes (i) applying a voltage across the sensor, and (ii) measuring a current flow through the sensor.

[0031] In some embodiments, detecting the presence or absence of the bioanalyte comprises comparing the electrochemical parameter measured in step b) to a reference value for that parameter of the sensor.

[0032] In some embodiments, the bioanalyte is interleukin-6 (IL-6) or C-reactive protein (CRP).

[0033] In some embodiments, the bioanalyte is a viral protein.

[0034] In some embodiments, the substance is a sample solution, and optionally the sample solution comprises a bodily fluid.

[0035] According to a third aspect of the present invention, there is provided a method for manufacturing a sensor for detecting a bioanalyte, the method comprising the steps of: preparing a substrate including a semiconductor portion, the semiconductor portion comprising a high resistivity non-oxide semiconductor; preparing a pair of terminal electrodes on the substrate that are spaced apart and facing each other, the semiconductor portion of the substrate being disposed between and in electrical contact with the terminal electrodes, and a conductive path between the terminal electrodes passing through the semiconductor portion; and immobilizing a bioanalyte binding site on a surface of the semiconductor portion, thereby preparing a sensor element comprising (i) the semiconductor portion and (ii) the bioanalyte binding site.

[0036] In some embodiments, the non-oxide semiconductor has a resistivity greater than 100 ohm-cm, hi some embodiments, the non-oxide semiconductor has a resistivity in the range of about 500 ohm-cm to about 50,000 ohm-cm, or in the range of about 1000 ohm-cm to about 10000 ohm-cm.

[0037] In some embodiments, the sensor has an electrical resistance in the range of about 10 kilohms to about 10,000 kilohms.

[0038] In some embodiments, the non-oxide semiconductor is selected from the group consisting of elemental semiconductors and compound semiconductors, hi some embodiments, the non-oxide semiconductor is an elemental semiconductor.

[0039] In some embodiments, the non-oxide semiconductor is a silicon semiconductor. The silicon semiconductor may be an intrinsic silicon semiconductor. The silicon semiconductor may be a float-zone type silicon semiconductor.

[0040] In some embodiments, the substrate includes the semiconductor layer as an integral part thereof. The substrate may be a wafer of a non-oxide semiconductor.

[0041] In some embodiments, immobilizing the bioanalyte binding moiety comprises chemically binding the bioanalyte binding moiety to the semiconductor moiety. Chemically binding the bioanalyte binding moiety to the semiconductor layer may comprise (i) silanizing the non-oxide semiconductor 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 carboxy group, and a hydroxy group, and (ii) reacting a precursor comprising the binding moiety with the terminal functional group. The silanizing agent may be selected from the group consisting of (3-glycidyloxypropyl)trimethoxysilane (GPS), (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS).

[0042] In some embodiments, the precursor comprising the binding site is a biomolecule or a molecularly imprinted polymer.

[0043] Other aspects of the invention are also disclosed.

[0044] 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]

[0045] [Figure 1]FIG. 1 shows a schematic diagram of a conductivity sensor for detecting a bioanalyte according to an embodiment of the invention, the sensor having a sensor element including a semiconductor portion of a sensor substrate comprising a high resistivity non-oxide semiconductor, with bioanalyte binding sites immobilized on the surface of the semiconductor portion. [Diagram 2] FIG. 2 shows a schematic diagram of a method for manufacturing the conductivity sensor shown in FIG. [Diagram 3] FIG. 3 shows a schematic diagram of a method for immobilizing bioanalyte binding sites on a surface of a semiconductor portion comprising a high resistivity non-oxide semiconductor according to an embodiment of the present invention. [Figure 4] FIG. 4 shows a plot reflecting the change in resistance (%) as a function of the concentration of IL-6 on a conductivity sensor functionalized with immobilized anti-IL-6 antibody, according to one embodiment of the present invention. [Diagram 5] FIG. 5 shows a plot reflecting the percent change in resistance as a function of the concentration of CRP on a conductivity sensor functionalized with immobilized anti-CRP antibody, according to another embodiment of the present invention. [Figure 6] FIG. 6 shows a plot reflecting the change in resistance (%) as a function of the concentration (mg / L) of SARS-COV-2 viral protein on a conductivity sensor functionalized with immobilized plastic antibody (MIP) according to another embodiment of the present invention. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0046] The present invention relates to a conductivity sensor for detecting a bioanalyte. The sensor comprises a substrate, a pair of terminal electrodes arranged on the substrate in a spaced apart facing relationship, and a sensor element arranged between and in electrical contact with the pair of terminal electrodes. The sensor element comprises (i) a semiconductor portion of the substrate comprising a high resistivity non-oxide semiconductor, and (ii) bioanalyte binding sites on a surface of the semiconductor portion. An electrically conductive path between the terminal electrodes passes through the semiconductor portion. In use, binding of a bioanalyte to the bioanalyte binding sites causes a change in the electrical resistance of the sensor. The increase in resistance can be ascertained by measuring the current response when a voltage is applied across the sensor, and thus the presence and / or concentration of a bioanalyte can be detected.

[0047] Thus, the sensor of the present invention uses conductivity sensing technology to detect various bioanalytes in fluids, such as bodily fluids such as human saliva, sweat, urine, tears, blood, plasma, interstitial fluid, or respiratory aerosols / droplets, for the prognosis / diagnosis of medical disease. As described in more detail below, the conductivity sensor has a simple and relatively easy to fabricate device structure that provides a cost-effective alternative to conventional non-invasive sensors that either require specialized substrates or employ detection technologies that limit their accuracy.

[0048] In particular, the sensor does not rely on a thin metal oxide layer for the conductivity layer, as disclosed in International Application PCT / AU2020 / 051396, but instead uses a high resistivity non-oxide semiconductor, such as an intrinsic silicon semiconductor, in the sensor element. Surprisingly, the inventors have found that various bioanalyte binding sites can be immobilized directly on such semiconductor materials, thereby resulting in sensor elements with excellent sensitivity and selectivity for various complementary bioanalytes. Advantageously, therefore, there is no need to deposit a metal oxide layer on the substrate by techniques such as reactive sputtering. Instead, the semiconductor portion of the sensor element may be an integral part of the substrate itself, resulting in a very simple yet highly effective device structure.

[0049] The inventors believe that the conductivity sensors described in more detail below are compatible with CMOS circuitry and therefore can be easily integrated into flexible / wearable electronic devices to provide portable, personalized, reusable sensors that can be used to continuously monitor levels of targeted bioanalytes via bodily fluids without the need for invasive procedures. These bioanalytes can act as biomarkers indicative of an individual's condition and health.

[0050] Below is a detailed description of a non-invasive conductivity sensor and its application method for detecting the levels of various bioanalytes (e.g., biomarkers) in body fluids. It should be noted that in the following description, similar or identical reference numbers in different embodiments indicate the same or similar features.

[0051] Sensors In its simplest form, as shown in the schematic diagram of Figure 1, the sensor 100 comprises a substrate 102, a pair of terminal electrodes 104, 106 disposed in spaced apart facing relationship on the substrate, and a sensor element 108 between and in electrical contact with the terminal electrodes 104, 106. The sensor element 108 includes a semiconductor portion 110 including a high resistivity non-oxide semiconductor 112 and bioanalyte binding sites 114 on a surface 116 of the semiconductor portion 110. A conductive path 120 between the terminal electrodes 104 and 106 passes through the semiconductor portion 110 and thus through the non-oxide semiconductor 112.

[0052] In the embodiment shown in Figure 1, the substrate 102 includes the semiconductor portion 110 as an integral part of the substrate, such that the remainder of the substrate is composed of the same high resistivity non-oxide semiconductor 112. The conductive path 120 between the terminals 104 and 106 is substantially confined to a surface layer of the substrate (the surface layer corresponding to the semiconductor portion 110) by the electric field lines that are established when a voltage is applied across the sensor in use. Advantageously, therefore, there is no need to fabricate a separate thin-film semiconductor layer on the sensor substrate. The substrate 102 may therefore be of any convenient thickness, for example as provided when using a wafer of the high resistivity non-oxide semiconductor 112.

[0053] Alternatively, the sensor element 108 may include a semiconductor portion 110 formed as a separate surface layer on the substrate 102, at least between the terminal electrodes 104 and 106, but optionally extending across the entire substrate surface. In such embodiments, the substrate 102 may be constructed from any suitable material capable of receiving and supporting the semiconductor layer 110.

[0054] In use, the sensor 100 is contacted with a substance, such as a sample solution 122, that contains (or may contain) a bioanalyte 124. If present, the bioanalyte binds to the bioanalyte binding sites 114, thereby causing a change in the electrical resistance of the sensor. The change in electrical resistance occurs due to a charge transfer by donating or accepting an electron to or from the semiconductor when the bioanalyte binds. When a voltage is applied across the sensor, i.e., between the terminal electrodes 104 and 106, the resulting current flowing between the terminal electrodes along the conductive path 120 can be measured, and thus the electrical resistance of the sensor can be determined. By comparing this resistance to a predetermined reference resistance of the sensor, the presence or absence of the bioanalyte in the sample solution 122 can be detected.

[0055] Those skilled in the art will appreciate that the sensor element 108 typically includes a plurality of bioanalyte binding sites 114, and the percentage of those bioanalyte binding sites that bind the bioanalyte 124 may depend on the concentration of the bioanalyte in the sample solution 122. Because the resistance of the conductive pathway 120 is proportional to the percentage of the binding sites 114 that are occupied, the concentration of the bioanalyte 124 in the fluid 116 can be determined, for example, by comparing the determined resistance with a calibration curve.

[0056] Each component of the conductivity sensor will now be described.

[0057] substrate In the broadest form of the invention, the substrate is generally not particularly limited and may be made of a material selected from the group consisting of semiconductors, polymers, glasses, or ceramics. In such embodiments, the semiconductor portion of the sensor element may be supported on a support layer of the substrate, optionally only in the substrate area covered by the sensor element. However, in some preferred embodiments, the substrate comprises or consists of a high resistivity non-oxide semiconductor content. Thus, as seen in FIG. 1, the semiconductor portion of the sensor element may be an integral part of the substrate, simplifying the device structure. In some embodiments, the substrate comprises a wafer of a high resistivity non-oxide semiconductor.

[0058] electrode The sensor includes a pair of terminal electrodes disposed on a substrate in spaced-apart, opposing relationship such that a sensor element of the sensor is disposed in a detection region between the spaced-apart terminal electrodes, and as will be appreciated by those skilled in the art, the terminal electrodes are conductive and configured for electrical connection to a device, such as a potentiostat, for applying a voltage across the sensor.

[0059] As shown in Figure 1, the terminal electrodes are formed as separate structures on the substrate surface and are in electrical contact with the underlying semiconductor portion, which includes a high resistivity non-oxide semiconductor, although other configurations are contemplated. For example, the terminal electrodes may be recessed into the substrate, and the semiconductor portion of the sensor element may lie horizontally between the terminal electrodes along the substrate surface.

[0060] The terminal electrodes may comprise a conductive metal or alloy, preferably a chemically inert metal or alloy. Gold is one example of a suitable metal.

[0061] In some embodiments, terminal electrodes are formed on the substrate by microfabrication techniques. Gold terminal electrodes can be formed by evaporating a thin gold film (250 nm with a 100 nm chromium adhesion layer) onto the semiconductor layer using electron beam lithography. The as-deposited gold film is then patterned using standard photolithography and wet etching techniques to define a pair of terminal electrodes.

[0062] The terminal electrodes may generally be sized and positioned relative to one another in any configuration suitable for the conductivity sensor. In some embodiments, the terminal electrodes are spaced apart by a distance ranging from 1 micrometer to 100 micrometers. In some embodiments, the terminal electrodes have a length (i.e., the length in the direction perpendicular to the inter-electrode gap distance) ranging from 200 to 4000 micrometers. The inventors have obtained good results using two parallel electrodes spaced 40 micrometers apart and 4000 micrometers long, achieving a 16×10 -8 m 2 A detection area with an area of

[0063] Sensor element The sensor comprises a sensor element that includes (i) a semiconductor portion of a substrate that includes or consists of a high resistivity non-oxide semiconductor, and (ii) bioanalyte binding sites on a surface of the semiconductor portion.

[0064] The sensor element is located between and in electrical contact with the terminal electrodes such that the device is configured such that an electrical conductive path between the terminal electrodes passes through the semiconductor portion and thus through the high resistivity non-oxide semiconductor of the semiconductor portion.

[0065] In some embodiments, as seen in FIG. 1, the semiconductor portion is an integral part of the substrate, specifically a region or surface portion of the substrate that extends across the sensing region between the terminal electrodes.

[0066] In other embodiments, the semiconductor portion is a separate surface layer of the substrate supported on a support layer of the underlying substrate. The semiconductor layer is located at least in the detection region between the terminal electrodes, but may optionally extend across the entire substrate surface. In such embodiments, the terminal electrodes may be formed, for example by metal deposition, on the surface of the separate semiconductor layer of the substrate. Alternatively, the terminal electrodes may be formed on the support layer, and the semiconductor portion of the substrate is subsequently formed on the support layer of the substrate at least in the detection region between the terminal electrodes.

[0067] The semiconductor portion of the substrate comprises, and typically consists of, a high resistivity non-oxide semiconductor. As used herein, non-oxide semiconductor includes both elemental and compound semiconductor materials, but does not include metal oxide semiconductors.

[0068] Common semiconductors used in electrochemical devices, including many non-oxide semiconductors such as doped silicon, are too conductive for use in conductivity sensor elements. Any effect on the electronic properties of such semiconductors caused by binding of bioanalytes to the surface is too small to provide sufficient sensitivity. For this reason, conventional conductivity sensors are typically constructed with a separate conductivity sensing layer of high resistivity polymer or metal oxide material.

[0069] However, it has now surprisingly been found that the use of a high resistivity non-oxide semiconductor in the conductivity sensor element can result in good conductivity sensor performance. By selecting a non-oxide semiconductor with high resistivity, the sensor has an overall resistance that is in a suitable range for the detection of bioanalytes when bound to the sensor element surface.

[0070] In some embodiments, high resistivity non-oxide semiconductors have a resistivity of greater than 100 ohm-cm, or greater than 200 ohm-cm, or greater than 500 ohm-cm, or greater than 1000 ohm-cm. In contrast, doped silicon semiconductors commonly used in electrochemical sensing devices generally have a resistivity of about 1 to 10 ohm-cm.

[0071] In some embodiments, the high resistivity non-oxide semiconductor has a resistivity in the range of 500 ohm-cm to about 50,000 ohm-cm, for example, in the range of about 1000 ohm-cm to about 10000 ohm-cm. The inventors have had good results with non-oxide semiconductors having resistivities of 1000-2000 ohm-cm and 5000-10000 ohm-cm.

[0072] The high resistivity non-oxide semiconductor may be selected so that the sensor has an appropriate electrical resistance when measured between the terminal electrodes (and along the conductive path). In some embodiments, the sensor has an electrical resistance in the range of about 10 kilohms to about 10,000 kilohms, for example, when no bioanalyte is bound to the bioanalyte binding sites. The inventors have found that the use of low resistance sensors results in very low sensitivity to the bioanalyte.

[0073] In some embodiments, the non-oxide semiconductor is selected from the group consisting of elemental semiconductors and compound semiconductors.

[0074] Suitable elemental semiconductors include silicon and germanium semiconductors, preferably silicon semiconductors. High purity intrinsic (undoped) silicon semiconductors have been found to be particularly suitable due to their resistive properties. The intrinsic silicon semiconductor may be float-zone silicon, which is high purity silicon prepared by a float-zone purification technique. In this technique, a molten zone passes slowly along a rod of silicon, and impurities preferentially remain in the molten zone rather than being reincorporated into the recrystallized silicon. In contrast, most silicon semiconductors are produced by the Czochralski process and therefore incorporate high levels of impurities, making the silicon too conductive for use in conductivity sensor elements. Suitable float-zone silicon is available in typical diameters (e.g., 3-inch and 4-inch diameters). <100> It is an oriented silicon wafer.

[0075] Intrinsic silicon semiconductors have been found to be particularly suitable, but this is not to be excluded as the non-oxide semiconductor may be a doped elemental semiconductor, provided that the doping level is sufficiently low to maintain the semiconductor highly resistive.

[0076] Suitable compound semiconductors include binary semiconductors such as gallium arsenide (GaAs), indium phosphide (InP) and indium antimony (InSb), and ternary semiconductors such as gallium aluminum arsenide (GaAlAs).

[0077] As previously mentioned, the semiconductor portion comprising the high resistivity non-oxide semiconductor may be an integral part of the substrate, i.e. the substrate may comprise or consist of the non-oxide semiconductor, for example the substrate may be a wafer of a non-oxide semiconductor, such as a wafer of a high resistivity intrinsic silicon semiconductor.

[0078] The surface of the semiconductor portion need not be nanostructured, and thus in some embodiments the high resistivity non-oxide semiconductor is not nanostructured, i.e., does not exist as discrete nanoparticles (dimensions less than 100 nm) or in the form of nanostructured surface features (dimensions less than 100 nm).

[0079] The sensor element comprises a bioanalyte binding site on a surface of the detection moiety, hi some embodiments, the sensor element comprises a plurality of such bioanalyte binding sites.

[0080] The bioanalyte binding site may be immobilized on the detection portion of the substrate by either physical absorption or chemical binding. In a preferred form, the bioanalyte binding site is chemically bound to the surface of the semiconductor portion. In some embodiments, the bioanalyte binding site is tethered to the semiconductor portion by an organic linker, which is covalently bound to the surface of the semiconductor portion. The covalent binding may occur by any suitable reaction, for example, a silanization reaction. The length of the organic linker may be selected to appropriately space the bioanalyte binding site from the surface of the semiconductor portion. A shorter linker is typically preferred to ensure that binding of the bioanalyte to the bioanalyte binding site produces a strong sensor response. However, in some embodiments, some spacing is preferred to allow the bioanalyte binding site to accept and bind the bioanalyte. Thus, the organic linker may include at least three or at least four, for example five or more atoms in the linking group between the bioanalyte binding site (or the biomolecule containing the bioanalyte binding site) and the terminal functional group of the organic linker that is covalently bound to the surface.

[0081] Non-oxide semiconductors, including silicon semiconductors, typically contain surface functional groups, such as hydroxy groups, that are susceptible to covalent bond-forming reactions with surface modifiers, such as silanizing agents (surface modifiers that contain silanizing groups, such as alkoxysilanes). Thus, bioanalyte binding sites can be chemically attached to the semiconductor moiety by a process that includes: (i) silanizing the non-oxide semiconductor with a silanizing agent having a terminal functional group selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxy groups, and hydroxy groups; and (ii) reacting a precursor that includes a bioanalyte binding site with the terminal functional group. As a result of this process, the binding site is immobilized on the surface of the semiconductor moiety by an organic linker that is a residue of the silanizing agent.

[0082] Suitable silanizing agents include (3-glycidyloxypropyl)trimethoxysilane (GPS), (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS). For example, when an epoxy-functionalized silanizing agent such as (3-glycidyloxypropyl)trimethoxysilane (GPS) is used, silanization of a non-oxide semiconductor functionalizes the surface with pendant epoxy groups. Thus, precursor molecules containing bioanalyte binding sites can be immobilized on this surface by conjugation reaction of epoxy-reactive functional groups such as amines present in the precursor molecule.

[0083] In another set of embodiments, the bioanalyte binding sites are initially present on a biomolecule or other entity that has been pre-functionalized with a surface-reactive functional group, such as a silanized group, and thus the bioanalyte binding site can be chemically attached to the semiconductor moiety by contacting the pre-functionalized biomolecule (or other entity) with the non-oxide semiconductor under conditions suitable to allow for covalent bond formation and thus surface immobilization.

[0084] The semiconductor portion of the sensor element may include an oxide surface layer on a non-oxide semiconductor, which may include surface functional groups susceptible to covalent bond formation reactions with a surface modifier. Such passivation layers are generally very thin, for example, about 1 nm for a native silicon oxide layer on a silicon semiconductor, so that binding of a bioanalyte to the bioanalyte bond changes the resistance of the underlying high resistivity non-oxide semiconductor during use, allowing tunneling of current through that layer. Thus, any oxide surface layer on a non-oxide semiconductor at the surface of the semiconductor portion may be less than 10 nm in thickness.

[0085] Bioanalyte Binding Sites The sensor element of the sensor comprises at least one bioanalyte binding site, typically a plurality of bioanalyte binding sites, on the surface of the semiconductor portion. The bioanalyte binding sites may suitably be located on biomolecules or non-biological entities immobilized on the semiconductor portion. As previously described herein, the binding sites may be immobilized on the semiconductor portion either by physical absorption or by chemical binding.

[0086] In some embodiments, the bioanalyte binding site is present on a natural or synthetic biomolecule immobilized on the semiconductor moiety. A wide variety of biomolecules can be utilized as binding sites for selective binding of the desired bioanalyte from a biological sample. For example, such biomolecules can include proteins, peptides, lipopeptides, protein-binding carbohydrates, or protein-binding ligands.

[0087] In some embodiments, the biomolecule is a capture protein. 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.

[0088] Protein-binding scaffolds have emerged as viable molecules for binding various bioanalytes, including proteins. Protein-binding scaffolds typically comprise stable protein structures (scaffolds) that can tolerate modification of amino acids within a designated binding region without altering the relative arrangement of the binding domains. These protein-binding scaffolds include (but are not limited to) adnectins, affilins (nanophytins), affibodies, affimer molecules, affitins, alphabodies, 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.

[0089] Affilins are artificially engineered proteins of approximately 20 kDa. They contain a scaffold structurally related to human ubiquitin and vertebrate gamma-B crystallin, with eight surface-exposed, manipulable 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.

[0090] Affibodies are approximately 6 kDa proteins that contain the protein scaffold of the Z domain of an IgG isotype antibody, with modifications to one or more of the 13 amino acid residues located in their two alpha-helical binding domains.

[0091] Affimer molecules are proteins of approximately 12 to 14 kDa that utilize a protein scaffold derived from the cystatin cysteine ​​protease inhibitor family. Affimer molecules contain two peptide loop regions in addition to an N-terminal sequence that can be adapted for target-specific binding. Affimer molecules with combinations of 1010 amino acids in the binding site can be generated using phage display libraries and appropriate techniques.

[0092] Affins are proteins consisting of 66 amino acid residues (approximately 7 kDa) and 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 can produce up to 3 × 10 12 It contains 14 binding amino acid residues that can be mutated to produce more than one structural variant. Screening techniques such as surface plasmon resonance can be used to identify the specific binding of these molecules.

[0093] Alphabodies are approximately 10 kDa molecules that, unlike most macromolecules, can penetrate cell membranes (if not immobilized) and therefore bind intracellular and extracellular molecules. The alphabody scaffold is based on a computer-designed coiled-coil structure with three alpha-helices (A, B, and C) that do not resemble the native structure. Amino acids on the A and C alpha helices can be modified to target specific antigens.

[0094] Aptamers that bind to proteins include a variety of nucleic acids (DNA, RNA, and XNA) and peptides that can be screened for binding to a specific target molecule. Databases of nucleic acid aptamers allow the selection of DNA aptamers identified in vitro. Peptide aptamers generally consist of short amino acid sequences embedded in a loop structure ("loop on frame") within a stable protein scaffold frame. Typically, a peptide loop of 5 to 20 residues is responsible for the variability of selective binding to the target molecule. 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 can be found in the literature.

[16] It is described in.

[0095] Anticalin proteins are protein binding molecules derived from lipocalins. Typically, anticalins bind to molecules smaller than antibodies. Methods for screening and developing anticalins are described in the literature.

[0096] Armadillo repeat protein-based scaffolds feature an armadillo domain composed of tandem armadillo repeats of approximately 42 amino acids organized into a superhelix of repeating units, each composed of three α-helices. Modification of residues within the conserved binding domain allows for the preparation of a variety of combinatorial libraries that can be used for the selection of target-specific binders.

[0097] Avimers (also known as avidity multimers, maxibodies or low density lipoprotein receptor (LDLR) domain A) contain at least two linked 30-35 amino acid long peptides based on the A domain of a range of cysteine-rich cell surface receptor proteins. Modification of the A domain allows for directed binding to different epitopes on the same target or across multiple targets, with the number of linked peptides determining the number of possible targets per avimer. Various avimer phage display libraries are known in the art, including commercially available libraries such as those from Creative Biolabs.

[0098] Designed ankyrin repeat proteins (DARPins) are engineered binding proteins derived from ankyrin proteins. Methods for screening and identifying DARPins are described in the literature.

[0099] The inhibitor cystine knot (ICK) scaffold is a family of mini-proteins (30 to 50 amino acid residues long) that form stable three-dimensional structures that contain three disulfide bridges connecting a series of loops with high sequence variability. The inhibitor cystine knot includes three family members: knottins, cyclotides and growth factor cysteine-knots. Databases are known in the art that disclose certain properties of known knottins and cyclotides, such as sequences, structures and functions, for example the KNOTTIN database (www.dsimb.inserm.fr / KNOTTIN / ). Furthermore, methods for making ICKs and screening for binding are described in the literature.

[0100] Monobodies (also known by the trade name AdNectins) utilize an FN3 (fibronectin type III domain) scaffold with diverse and engineerable variable groups. Adnectins share antibody variable domains and beta-sheet loops 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 producing monobodies have been described in the literature.

[0101] In some embodiments, the biomolecule comprising the bioanalyte binding site is an antibody or a binding fragment thereof. An antibody can potentially bind up to 10 11 From 10 12 Antibody diversity is exemplified by protein-binding molecules with as many as 10 unique molecules, with further diversity possible due to genetic differences between individuals. In vivo antibody diversity is induced by random recombination of a series of genes at the V(D)J junctions.

[0102] 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 (VH) chain CDR1, CDR2 and CDR3, and variable light (VL) chain 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 of the CDRs, primarily CDR3.

[0103] Antibodies to specific bioanalytes can be commercially obtained or can be made by methods known in the art. For example, antibodies to specific bioanalytes can be prepared using methods generally disclosed in the literature (e.g., Howard and Kaser, Making and Using Antibodies: A Practical Handbook, CRC Press, 2007).

[0104] The specificity, avidity and affinity of antibodies produced in a subject can be modified by in vitro processes such as affinity maturation. Thus, in vivo derived antibodies can be further modified to generate distinct, but systematically related, antibodies. As a result, the term "antibody" encompasses in vivo derived antibodies as well as in vitro derived molecules that have undergone a process of mutation to modify the CDR binding sites to have unique sequences compared to antibodies produced in vivo.

[0105] The term antibody also includes non-conventional antibodies produced from species such as camelids, sharks and jaw fish. Thus, the term antibody includes heavy chain antibodies such as camelid antibodies, IgNARs and variable lymphocyte receptors (VLRs). Furthermore, these may be fragmented into their binding moieties (e.g., the single binding moiety of VNAR-IgNAR) or recombinantly incorporated into fusion proteins. Methods for making and adapting such non-conventional antibodies are described in the literature.

[0106] In some embodiments, the biomolecule is an antibody-binding fragment.Antibody-binding fragments can be derived from antibodies or can be produced by recombinant techniques using sequences identical to the CDRs of antibodies or antibody fragments.Indeed, these CDRs can be derived from affinity-matured antibodies and therefore may not be identical to in vivo derived antibodies.

[0107] Antibodies are composed of four chains (two heavy and two light) and can be separated into Fc (crystallizable portion) and Fab (antibody portion) domains. The Fc portion of an antibody interacts with the Fc receptors and the complement system. It is therefore important for the immune function of the antibody, while the Fab portion contains the binding region of the antibody and is important for the specificity of the antibody for the desired epitope.

[0108] Thus, in some embodiments, the biomolecule containing the bioanalyte binding site is a Fab fragment of an antibody. The Fab fragment may be an individual Fab fragment (i.e., an antibody fragment is generated when the linking disulfide bridge is released) or a F(ab')2 fragment that contains two Fab fragments of an antibody linked via 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.

[0109] 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 of four framework regions). The constant regions of the Fab fragment can be removed, leaving 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. Generally, individual binding regions are known as single antibody domains (sdAbs). Alternatively, the VH and VL chains can be linked via a linker to form a fusion protein known as a single chain variable fragment (also known as scFv-diabody). Unlike Fabs, scFvs are not fragmented from antibodies, but rather are usually formed recombinantly based on the CDRs and framework regions of antibodies. Additionally, sdAbs can be recombinantly produced to form larger fusion protein binding components, which may also contain moieties that may act to stabilize the binding region, improve or facilitate immobilization to the sensor element or intermediate layer, improve binding, for example by providing flexibility of the binding region or optimizing the length of the bioanalyte binding site, thereby allowing access to antigenic regions of the bioanalyte. Thus, in some embodiments, the biomolecule comprising the bioanalyte binding site is or comprises an scFv or sdAb. The scFv may comprise multiple VH and VL chains that are linked together to form a multivalent scFv, such as a di-scFv or tri-scFv.

[0110] Antibodies and antibody fragments, or fusion proteins containing antibody-derived sequences, directed against specific bioanalytes can be obtained commercially or produced by methods known in the art, as described above.

[0111] In some embodiments, the biomolecule comprising the bioanalyte binding site is a protein receptor or ligand that interacts with and binds to a protein. Such receptors and ligands include any receptor or ligand, or specific fragments thereof (e.g., fragments comprising the binding domain of the receptor or ligand). Particularly contemplated receptors include receptors for cytokines such as interleukins or chemokines, which may provide information regarding the state of the immune system. In some embodiments, the receptor or ligand (or fragments thereof) may be incorporated to form a fusion protein.

[0112] For example, interleukin-6 (IL-6) is a proinflammatory pluripotent cytokine and is an important biomarker that can be used to monitor immune responses during cancer treatment. It can also be used to monitor psychological stress and insulin activity.

[0113] For example, the inventors have obtained good results when using anti-interleukin-6 (IL-6) antibodies for selective recognition and binding of IL-6. For example, the inventors have obtained good results when using anti-C-reactive protein (CRP) antibodies for selective recognition and binding of CRP.

[0114] In some embodiments, the bioanalyte binding site is present on a non-biological entity immobilized on the semiconductor moiety. In some embodiments, the non-biological entity is a molecularly imprinted polymer with binding sites that mimic the biological binding sites of the target biomolecular analyte. Such polymers may have an extended shelf life because they are non-biological and do not degrade or denature like biological antibodies, which have limited shelf life due to degradation / denaturation over time. The inventors have obtained good results using commercially available molecularly imprinted polymers custom designed for selective binding to SARS-COV-2 proteins.

[0115] Detection Method The present invention also relates to a method of detecting a bioanalyte, the method comprising the steps of contacting a sensor element of a sensor described herein with a substance that may contain the bioanalyte, measuring an electrochemical parameter of the sensor that corresponds to the resistance of the sensor, and detecting the presence or absence of the bioanalyte on the sensor element based on the measured electrochemical parameter.

[0116] In a typical operation of a conductivity sensor, the parameter that is directly measured is the current response when a known voltage (or voltage profile) is applied across the sensor. Thus, in some embodiments, the method includes applying a voltage across the sensor, measuring the current flow through the sensor, and detecting the presence or absence of a bioanalyte on the sensor element based on the current flow. Conventional equipment for conductivity sensors, such as a potentiostat, can be used to apply the voltage and measure the current flow.

[0117] However, it is not excluded that a different electrochemical parameter corresponding to the sensor resistance may also be measured. For example, it is in principle possible to pass a given current through the sensor and measure the voltage required to achieve this current. The measured voltage then corresponds to the sensor resistance.

[0118] The substance in contact with the sensor can be any substance that contains or can contain a bioanalyte of interest, in some embodiments the substance is a liquid sample that is or includes a sample solution, e.g., a bodily fluid such as saliva, sweat, urine, tears, blood, plasma, interstitial fluid, or respiratory aerosols / droplets.

[0119] The presence or absence of a bioanalyte can be detected by comparing the measured electrochemical parameter to a reference value of that parameter for the sensor. If the measured parameter is a current response, the current flow or the electrical resistance of the sensor determined from the current flow can be compared to a predefined reference current flow or resistance of the sensor that corresponds to the presence or absence of a bioanalyte on the sensor element. For example, the current flow (or resistance) of the sensor after contact with the substance can be compared to the current flow (or resistance) of the sensor after contact with a reference solution that does not contain the bioanalyte.

[0120] In its simplest form, such a comparison can be used to determine the presence or absence of a bioanalyte in a substance. Alternatively, the current flow (or resistance, or other measured electrochemical parameter) of the sensor after contact with a sample solution containing the bioanalyte can be compared to a calibration curve that plots the current flow (or resistance, or other measured electrochemical parameter) of the sensor after contact with a series of reference solutions with known concentrations of the bioanalyte. In this way, the concentration of the bioanalyte in the sample solution can be calculated.

[0121] The method can optionally include one or more preparation steps between contacting the sensor element with the substance and applying the voltage. For example, if the substance is a sample solution, the sensor element can be incubated at defined conditions (e.g., temperature) for a defined time to allow binding of the bioanalyte (if present in the sample solution) to the bioanalyte binding sites. The sample solution can then be removed from the sensor and the sensor element can be dried before performing a conductivity measurement.

[0122] Alternatively, the sensor can be incorporated, for example, into a microneedle and used as an invasive sensor inserted into the human body for in situ detection of the bioanalyte. In another embodiment, the sensor is incorporated into a wearable device for monitoring the bioanalyte in human sweat.

[0123] A variety of bioanalytes corresponding to the bioanalyte binding sites described herein can be detected by the methods of the present disclosure. Thus, non-limiting examples of bioanalytes include viral proteins, cytokines, and proteins, including C-reactive protein (CRP).

[0124] How the sensor is manufactured The present invention also relates to a method of manufacturing a sensor for detecting a bioanalyte, the method comprising the steps of providing a substrate including a semiconductor portion, the semiconductor portion comprising a high resistivity non-oxide semiconductor. A pair of terminal electrodes are fabricated on the substrate in a spaced apart and opposing relationship such that the semiconductor portion of the substrate is disposed between and in electrical contact with the terminal electrodes, and a conductive path between the terminal electrodes passes through the semiconductor portion. A bioanalyte binding site is then immobilized on a surface of the semiconductor portion, thereby fabricating a sensor element including (i) a semiconductor portion and (ii) a bioanalyte binding site.

[0125] In one aspect of the invention, a substrate 102 is provided in step A, as shown in the schematic diagram of Figure 2. The substrate 102 includes a semiconductor portion 110 that includes a non-oxide semiconductor 112 as described herein. In the embodiment shown in Figure 2, the substrate 102 includes the semiconductor portion 110 as an integral part of the substrate, such that the remainder of the substrate is composed of the same high resistivity non-oxide semiconductor 112. Alternatively, the substrate 102 may include the semiconductor portion 110 formed as a separate thin surface layer on an underlying support layer, which may be composed of any suitable material capable of accepting and supporting the semiconductor layer 110.

[0126] In step B, a pair of terminal electrodes 104, 106 are fabricated on the substrate 102 in a spaced apart, opposing relationship. The electrodes are fabricated such that the semiconductor portion 110 of the substrate is disposed between the terminal electrodes 104, 106 and is in electrical contact with the terminal electrodes. Thus, the conductive path 120 between the terminal electrodes 104 and 106 passes through the semiconductor portion 110 and, in turn, through the non-oxide semiconductor 112.

[0127] In step C, bioanalyte binding moieties 114 are immobilized on surface 116 of the semiconductor portion, thereby creating sensor element 108. Although Figure 1 shows a single binding moiety, it will be appreciated that multiple bioanalyte binding moieties 114 can be immobilized on surface 116. Sensor element 108 includes semiconductor portion 110 and bioanalyte binding moieties 114. Thus, sensor 100 is fabricated after performing steps A, B, and C, as previously described herein with reference to Figure 1.

[0128] The substrate including the semiconductor portion may be according to any of the embodiments described herein in relation to the sensor of the present invention.

[0129] The terminal electrodes can be fabricated on the substrate by any suitable method. In some embodiments, the terminal electrodes are formed by microfabrication techniques. Gold terminal electrodes can be formed by evaporating a thin gold film (250 nm with a 100 nm chromium adhesion layer) onto the semiconductor layer using electron beam lithography. The as-deposited thin gold film is then patterned using standard photolithography and wet etching techniques to define a pair of terminal electrodes.

[0130] The bioanalyte binding moieties may be immobilized on the surface of the semiconductor portion by either physical absorption or chemical bonding. In a preferred form, the bioanalyte binding moieties are chemically bonded to the surface of the semiconductor portion.

[0131] Non-oxide semiconductors, including silicon semiconductors, typically contain surface functional groups, such as hydroxy groups, that are susceptible to covalent bond-forming reactions with surface modifiers, such as silanizing agents (surface modifiers that contain silanizing groups, such as alkoxysilanes). Thus, bioanalyte binding sites can be chemically attached to the semiconductor moiety by a process that includes: (i) silanizing the non-oxide semiconductor with a silanizing agent having a terminal functional group selected from the group consisting of epoxy groups, thiol groups, amino groups, carboxy groups, and hydroxy groups; and (ii) reacting a precursor that includes a bioanalyte binding site with the terminal functional group. As a result of this process, the binding site is immobilized on the surface of the semiconductor moiety by an organic linker that is a residue of the silanizing agent.

[0132] Suitable silanizing agents include (3-glycidyloxypropyl)trimethoxysilane (GPS), (3-mercaptopropyl)trimethoxysilane (MTS), (3-aminopropyl)triethoxysilane (APTES), and N-(2-aminoethyl)-3-aminopropyl-trimethoxysilane (AEAPTS).

[0133] In one exemplary embodiment, as shown in the schematic diagram of FIG. 3, the semiconductor portion 310 of the substrate beneath and between the gold (Au) terminal electrodes comprises a highly resistive non-oxide semiconductor, in this case a highly resistive intrinsic silicon wafer. The surface of the semiconductor portion is contacted with a silanizing agent 350, which may optionally be an epoxy-functionalized silanizing agent such as (3-glycidyloxypropyl)trimethoxysilane (GPS). The silanizing agent reacts with the surface hydroxy (-OH) functional groups of the semiconductor portion, thereby immobilizing the silanizing agent to the surface via covalent bonds and functionalizing the surface with pendant conjugated groups 352, in this case epoxy groups. A precursor molecule 354 containing a bioanalyte binding site 314 is then immobilized to the surface by a conjugation reaction of an epoxy-reactive functional group, in this case an amine (-NH2), present in the precursor molecule. That is, the bioanalyte binding site 314 is immobilized to the surface of the semiconductor portion 310 by an organic linking group 356, which is a residue of the silanizing agent 350.

[0134] The bioanalyte binding sites are typically immobilized on the surface of the semiconductor portion by immobilizing a pre-existing precursor that contains the bioanalyte binding site. The precursor can generally be any molecule or other entity (including biomolecules and non-biological entities) that contains a bioanalyte binding site according to any of the embodiments described herein in connection with the sensors of the present invention. In some embodiments, the precursor is a biomolecule that contains the bioanalyte binding site.

[0135] In some embodiments, the bioanalyte binding sites are initially present on a biomolecule or other entity that has been pre-functionalized with a surface-reactive functional group, such as a silanized group, and thus the bioanalyte binding site can be chemically attached to the semiconductor moiety by contacting the pre-functionalized biomolecule (or other entity) with the non-oxide semiconductor under conditions suitable to allow for covalent bond formation and thus surface immobilization. EXAMPLES

[0136] Materials and Methods High resistivity silicon wafers (diameter 100 mm) with resistivities of 1000–2000 ohm cm and 5000–10000 ohm cm were purchased from D&X Co., Ltd., Japan. Both types were single-sided polished silicon wafers. The orientation of the 1000–2000 ohm cm wafers was <100> The thickness was 500±10μm. The wafer orientation was 5000-10000 ohm·cm. <100> and the thickness was 450±25 μm.

[0137] Silicon wafer sensors were fabricated by patterning two terminal in-plane electrodes on a high resistivity silicon wafer using standard photolithography processes. The electrode gap was varied from 1-2 μm to 100 μm. However, this electrode gap was optimized to be 40 μm for best sensor performance. The electrode length was varied from 200 μm to 4000 μm. The optimal electrode length was chosen to be 4000 μm. Thus, the sensor element area (silicon substrate area between the electrodes) was 16 × 10-8 m 2 It was.

[0138] 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. SARS-CoV-2 molecularly imprinted polymer (MIP) was purchased from MIP Diagnostics Ltd. SARS-CoV-2 spike protein with His tag (S-RBD) was purchased from ThermoFisher Scientific and used as received.

[0139] The concentration of the anti-IL-6 stock solution as received was 48 mM. For use in immobilizing anti-IL-6 on the surface of silicon wafer sensors, the anti-IL-6 stock solution was diluted 1:10 in phosphate buffered saline (PBS, pH 7.4). 6 The concentration of the as-received anti-CRP stock solution was 4 μM. To immobilize anti-CRP, the as-received anti-CRP solution was diluted 1:50 in PBS (pH 7.4). The concentration of the as-received SARS-CoV-2 nanoMIP solution was 0.339 mg / mL and was used undiluted for the experiments.

[0140] A standard series of IL-6 solutions was prepared by completely dissolving as-received IL-6 powder in a known amount of autoclaved Milli-Q water and diluting with pH 7.4 PBS solution. The prepared IL-6 concentrations were 4 nM, 4 pM, and 4 fM. A standard series of CRP solutions was also prepared by diluting as-received CRP solutions in a given volume of pH 7.4 PBS. The prepared CRP concentrations were 13 nM, 13 pM, and 13 fM. The concentration of the as-received SARS-CoV-2 spike protein solution was 1 mg / mL, and a standard series of SARS-CoV-2 spike protein solutions was prepared by diluting with a given volume of pH 7.4 PBS solution. The standard series of SARS-CoV-2 spike protein consisted of 0.1mg / mL, 0.01mg / mL, 1μg / mL, 0.1μg / mL, 0.01μg / mL, 1ng / mL, 0.1ng / mL, 0.01ng / mL and 1pg / mL.

[0141] The conductance of the sensor was measured using a commercial current source meter (B2901A precision source / measure unit from Keysight Technologies). The sensor was placed on an LTS120 Linkam stage as a sensor holder for all measurements. Keysight Quick IV Measurement software was used for data acquisition.

[0142] The bias between the electrodes was maintained at 1.8 V. Sensor resistance measurements were taken after antibody immobilization and antigen immobilization. Data acquisition time for a given sensor was 1 min.

[0143] Example 1. Preparation of GPS silanized silicon wafer sensor: Silanization of the silicon wafer sensor surface with (3-glycidyloxypropyl)trimethoxysilane (GPS) (Sigma Aldrich) was performed after exposing the freshly prepared sensor device to O2 plasma (Plasma Cleaner PDC-002, Harrick Plasma) for 10 min to activate the hydroxyl groups on the silicon surface. Then, 20 μL of the freshly prepared GPS solution was drop-cast onto an Al foil, which was placed in a vacuum desiccator, resulting in GPS vapor within the desiccator. The O2 plasma cleaned silicon sensor was then exposed to this GPS vapor for 30–45 min in an LC200 glove box system. The silanized silicon wafer sensor was then rinsed thoroughly with Milli-Q water for 2 min to remove unbound silane groups from the surface. The cleaned sensor was then heated at 150 °C for 10 min to enhance the binding of the silane groups to the silicon wafer surface. These GPS silanized silicon wafer sensors, functionalized with surface epoxide functional groups chemically bonded to the substrate surface, were then used to immobilize various bioanalyte binding sites, including antibodies (containing antigen binding sites).

[0144] Example 2. Immobilization of biological antibodies and conductometric measurements of antigens: Immobilization of antibodies (IgG) onto GPS silanized silicon wafer sensors was performed as follows: 1:10 in a volume of 15 μL of freshly prepared 6 The diluted anti-IL-6 solution (i.e., at a concentration of 48 nM) was uniformly drop-cast onto the surface of a freshly GPS-silanized silicon wafer sensor and incubated for 1 h to immobilize the IL-6 antibody onto the surface of the sensor. This immobilization occurs by reaction of epoxide-reactive functional groups, such as amines, on the antibody with epoxide functional groups on the silanized silicon wafer surface. The sensor was then rinsed with a PBS solution at pH 7.4 to remove unbound antibody. The PBS-washed functionalized sensor was then dried in a stream of N2 gas. These anti-IL-6 antibody-immobilized sensors were used for IL-6 antigen concentration measurements. Following the same procedure, CRP-immobilized GPS-silanized silicon wafer sensors were prepared using 15 μL of freshly prepared 1:50 diluted anti-CRP solution (i.e., at a concentration of 80 nM).

[0145] Prior to the addition of antigen, the baseline conductance of the antibody-immobilized silicon wafer sensor was measured. A volume of 15 μL of antigen solution of known concentration (IL-6 concentrations 4 nM, 4 pM, 4 fM, CRP concentrations 13 nM, 13 pM, 13 fM) was drop-cast onto the surface of the antibody-immobilized silicon wafer sensor and incubated for 10 min. After such time had elapsed, the antigen solution remaining on the sensor was removed and the surface was dried under a flow of N2 gas. The sensor was then subjected to conductance measurements to measure the sensor resistance corresponding to each antigen solution concentration. Three individual sensors were used for a given antigen concentration and the average resistance change was calculated.

[0146] The results shown in Figures 4 and 5 were obtained from sensors fabricated on silicon wafers with resistivity of 1000-2000 ohm·cm.

[0147] The resistance change for both IL-6 and CRP antigens showed an increase with increasing antigen concentration. The resistance change was calculated by determining the difference in the resistance value of the sensor before (R0) and after (R) antigen immobilization (i.e., R-R0) relative to the resistance of the sensor before antigen immobilization (R0). Both IL-6 and CRP antigens showed a nonlinear increase in the resistance change as a function of antigen concentration. The contribution of the matrix to the resistance change was evaluated by determining the resistance change of PBS on the corresponding antibody-immobilized sensor. The resistance change of PBS on the sensors functionalized with IL-6 and CRP antibodies was only 1% and 6%, respectively. In contrast, at concentrations of IL-6 and CRP in sweat and saliva of healthy humans, the sensor produced a much higher resistance change than PBS. This indicates that the interference from PBS is negligible at clinically relevant concentrations of IL-6 and CRP.

[0148] The present non-oxide semiconductor sensor detected both IL-6 and CRP concentrations that differ from the reported concentrations of these two antigens in healthy human saliva and sweat. The reported IL-6 concentration is about 0.4 pM (10 ng / L) in healthy human sweat (Journal of Immunological Methods, 2006, 315, 99) and about 0.6 pM (16 ng / L) in healthy human saliva (BioMed Research International, 2018, 2018, 8531961). The reported CRP concentration is about 0.5 pM (12 ng / L) in healthy human sweat (Inflammatory Bowel Disease, 2020, 26, 1533) and about 12 pM (285 ng / L) in healthy human saliva (Journal of Immunological Methods, 2011, 373, 19). Typically, the concentrations of these two antigens in human sweat and saliva are elevated when inflammation occurs, compared with their concentrations in healthy body fluids. The sensor showed a resistance change of about 3% for an IL-6 concentration of 4 fM, and a resistance change of about 7% for a CRP concentration of 13 fM. These IL-6 and CRP concentrations are at least 1 / 100 times lower than the reported concentrations in healthy human saliva and sweat. This suggests that the conductivity sensor introduced in the present invention is extremely sensitive in detecting IL-6 antigen and CRP antigen in human body fluids.

[0149] Example 3. Immobilization of plastic antibody (MIP) and conductivity measurement of SARS-CoV-2 spike protein: SARS-CoV-2 molecularly imprinted polymers (MIPs), also known as plastic antibodies, were immobilized on non-oxide semiconductor sensors as follows. A volume of 15 μL of as-received SARS-CoV-2 nanoMIP solution (0.339 mg / mL) was uniformly drop-cast onto the surface of a freshly GPS-silanized silicon wafer sensor and incubated for 1 h to immobilize the SARS-CoV-2 nanoMIPs on the sensor surface. The sensor was then rinsed with a PBS solution at pH 7.4 to remove unbound nanoMIPs. The PBS-washed functionalized sensors were then dried in a stream of N2 gas. These SARS-CoV-2 nanoMIP-immobilized sensors were used for the concentration measurement of SARS-CoV-2 spike protein.

[0150] The baseline conductance of the nanoMIP-immobilized silicon wafer sensor was measured before the addition of SARS-CoV-2 spike protein. A volume of 15 μL of SARS-CoV-2 spike protein solutions of known concentrations (0.1 mg / mL, 0.01 mg / mL, 1 μg / mL, 0.1 μg / mL, 0.01 μg / mL, 1 ng / mL, 0.1 ng / mL, 0.01 ng / mL and 1 pg / mL) was drop-cast onto the surface of the nanoMIP-immobilized silicon wafer sensor and incubated for 10 min. After such time, the SARS-CoV-2 spike protein solution remaining on the sensor was removed and the surface was dried under a flow of N2 gas. The sensor was then subjected to conductance measurements to determine the sensor resistance corresponding to each SARS-CoV-2 spike protein solution concentration. Three individual sensors were used for a given antigen concentration and the average resistance change was calculated.

[0151] The results shown in Figure 6 were obtained from sensors fabricated on silicon wafers with resistivities of 1000–2000 ohm·cm.

[0152] The change in resistance for SARS-CoV-2 spike protein increased nonlinearly with increasing protein concentration. The change in resistance was calculated by determining the difference in the resistance of the sensor before (R0) and after (R) SARS-CoV-2 spike protein immobilization (i.e., R-R0) relative to the resistance of the sensor before SARS-CoV-2 spike protein immobilization (R0). The smallest positive change was observed at 0.1 ng / mL SARS-CoV-2 spike protein solution, suggesting that the detection limit of the proposed sensor is 0.1 ng / mL. The contribution of PBS was -48%, suggesting that there was no interference from PBS in the protein measurement.

[0153] definition Whenever a range is given herein, e.g., a temperature range, a time range, or a concentration range, all intermediate ranges and subranges, as well as all individual values ​​that fall within the given range, are intended to be included in the disclosure. Of course, any subrange contained in the description herein, or any individual value within a range or subrange, can be excluded from the claims herein.

[0154] All definitions and those used herein should, of course, be understood to take precedence over dictionary definitions, definitions in documents incorporated by reference, and / or ordinary meanings of the defined terms.

[0155] As used herein, the indefinite articles "a" and "an" should be understood to mean "at least one," unless expressly indicated otherwise.

[0156] 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., "one or more" of the elements so conjoined. Other elements other than the elements specifically identified by the "and / or" clause may optionally be present, whether related to the elements specifically identified or not. Thus, as a non-limiting example, a reference to "A and / or B", when used with an open-ended expression such as "comprising", may in one embodiment refer to only A (optionally including elements other than B), in another embodiment may refer to only B (optionally including elements other than A), and in yet another embodiment may refer to both A and B (optionally including other elements), and so forth.

[0157] While the invention has been described in conjunction with a limited number of embodiments, it will be apparent to those skilled in the art that many alternatives, modifications, and variations are possible in light of the foregoing description, and therefore, the present invention is intended to embrace all such alternatives, modifications, and variations that may fall within the spirit and scope of the invention disclosed.

[0158] When the terms "comprise", "comprises", "comprised" or "comprising" are used in this specification (including the claims), they should 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.

Claims

1. A sensor for detecting a bioanalyte, comprising: a substrate; a pair of terminal electrodes disposed on the substrate in a spaced-apart and opposing relationship to each other; a sensor element between the pair of terminal electrodes and in electrical contact with the pair of terminal electrodes, wherein the sensor element comprises: (i) a semiconductor portion of the substrate, including a high-resistivity non-oxide semiconductor, and a conduction path between the terminal electrodes penetrating the semiconductor portion; (ii) a bioanalyte binding site on the surface of the semiconductor portion; and a sensor in which binding of the bioanalyte to the bioanalyte binding site causes a change in the electrical resistance of the sensor.

2. The sensor according to claim 1, wherein the non-oxide semiconductor has a resistivity greater than 100 ohm·cm.

3. The sensor according to claim 1 or claim 2, wherein the non-oxide semiconductor has a resistivity in the range of about 500 ohm·cm to about 50,000 ohm·cm.

4. The sensor according to any one of claims 1 to 3, having an electrical resistance in the range of about 10 kiloohms to about 10,000 kiloohms.

5. The sensor according to any one of claims 1 to 4, wherein the non-oxide semiconductor is an intrinsic silicon semiconductor.

6. The sensor according to claim 5, wherein the silicon semiconductor is a float zone silicon semiconductor.

7. The sensor according to any one of claims 1 to 6, wherein the substrate includes the semiconductor portion as an integral part thereof.

8. The sensor according to any one of claims 1 to 7, wherein the bioanalyte binding site is chemically bonded to the semiconductor portion.

9. The sensor according to claim 8, wherein the bioanalyte binding site is chemically bonded to the semiconductor layer by a process including (i) silanizing the non-oxide semiconductor 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 carboxy group, and a hydroxy group, and (ii) reacting a precursor including the bioanalyte binding site with the terminal functional group.

10. The sensor according to any one of claims 1 to 9, wherein the bioanalyte binding site is present on a biomolecule or a molecularly imprinted polymer.

11. The sensor according to any one of claims 1 to 10, wherein the bioanalyte binding site is present on a biomolecule selected from the group consisting of proteins, peptides, lipopeptides, protein-binding carbohydrates, and protein-binding ligands.

12. wherein the biomolecule is a capture protein, The sensor according to claim 10 or claim 11, wherein 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.

13. The protein-binding scaffold is selected from the group consisting of adnectin, affilin, affibody, affimer molecule, affitin, alphabody, aptamer, anticalin, armadillo repeat protein-based scaffold, atrimer, avimer, designed ankyrin repeat protein (DARPins), finomer, inhibitory cystine knot (ICK) scaffold, Kunitz domain peptide, monobody and / or nanobody, The sensor according to claim 12, wherein the binding fragment of the antibody comprises Fab, (Fab')2, Fab', single-chain variable fragment (scFv), di- and tris scFv, single-domain antibody (sdAb), diabody, or a fusion protein containing a binding domain of an antibody.

14. The sensor according to any one of claims 1 to 13, wherein the bioanalyte binding site binds to interleukin-6 (IL-6) or C-reactive protein (CRP) or a viral protein.

15. A method for detecting a bioanalyte, comprising: a) contacting the sensor element of the sensor according to any one of claims 1 to 14 with a substance that may contain the bioanalyte; b) measuring an electrochemical parameter of the sensor corresponding to the resistance of the sensor; c) detecting the presence or absence of the bioanalyte on the sensor element based on the electrochemical parameter measured in step b). A method comprising the steps of

16. The method according to claim 15, wherein the substance is a sample solution, and the sample solution contains body fluid.

17. A method for manufacturing a sensor for detecting a bioanalyte, comprising: preparing a substrate including a semiconductor portion, wherein the semiconductor portion includes a high-resistivity non-oxide semiconductor. A step of fabricating a pair of terminal electrodes on the substrate in a relation of being spaced apart and facing each other, wherein the semiconductor portion of the substrate is disposed between the terminal electrodes and is in electrical contact with the terminal electrodes, and a conductive path between the terminal electrodes penetrates the semiconductor portion. A step of immobilizing a bioanalyte binding site on the surface of the semiconductor portion, thereby fabricating a sensor element including (i) the semiconductor portion and (ii) the bioanalyte binding site. A method comprising the above. **Claim 18** The non-oxide semiconductor has a resistivity greater than 100 ohm·cm. The method according to claim 17, wherein the non-oxide semiconductor is an intrinsic silicon semiconductor. **Claim 19** Immobilizing the bioanalyte binding site includes chemically bonding the bioanalyte binding site to the semiconductor portion. Chemically bonding the bioanalyte binding site to the semiconductor layer includes (i) silanizing the non-oxide semiconductor 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 hydroxy group, and (ii) reacting a precursor including the binding site with the terminal functional group. The method according to claim 18.

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