Polymer coating for the electrodes of a sensor device

Abstract

The present invention relates to the field of sensor devices for electrochemically analyzing samples using electrochemical or electrofield-effect-based detection methods. In particular, the present invention relates to an electrode comprising a functionalized surface that can be exposed to a fluid sample, the functionalized surface comprising at least one polymer capable of mediating a salting-out effect and at least one detection agent that binds to an analyte contained in the fluid sample, wherein the at least one polymer capable of mediating a salting-out effect and the at least one detection agent are distributed on the electrode surface such that the detection agent is present in substantially equal amounts per surface area over the entire electrode surface, and the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that it can reduce the ionic strength of the fluid near the detection agent and can bind to an analyte contained in the fluid sample. The present invention further relates to a method for manufacturing the functionalized surface on the electrode, to an analyte detector for determining at least one analyte comprising the electrode, and to the use of the analyte detector for determining at least one analyte in a fluid sample.

Description

[0001] The present invention relates to the field of sensor devices for analyzing samples using electrochemical or field - effect - based detection methods. In particular, the present invention relates to an electrode comprising a functionalized surface that can be exposed to a fluid sample, the functionalized surface comprising at least one polymer capable of mediating a salting - out effect and at least one detection agent that binds to an analyte contained in the fluid sample, wherein the at least one polymer capable of mediating a salting - out effect and the at least one detection agent are distributed on the electrode surface such that the detection agent is present in a substantially equal amount per surface area over the entire electrode surface, and the polymer capable of mediating a salting - out effect is arranged around the detection agent in an amount such that it can reduce the ionic strength of the fluid near the detection agent and can bind to an analyte contained in the fluid sample. The present invention further relates to a method for manufacturing the functionalized surface on the electrode, to an analyte detector for determining at least one analyte comprising the electrode, and to the use of the analyte detector for determining at least one analyte in a fluid sample.

[0002] In recent years, great progress has been made in applying field - effect - transistor (FET) - based sensors to biological detection. FET - based measurements require binding charged molecules to the sensor surface, changing its surface potential, and thus changing the channel current inside the transistor. This charge - based sensing mechanism makes FET measurements label - free and highly sensitive to a variety of biological targets. For example, FET - biosensors have been applied to detect various biomolecules such as nucleic acids (DNA and RNA), enzymes, protein disease markers, and even whole viruses, bacteria, and eukaryotic cells. Due to their versatility, high sensitivity, and rapid response, FET - based biosensors can be well applied to point - of - care (PoC) devices.

[0003] Despite the above advantages, to date, FET - based biosensing has been limited to measurements in low - ionic - strength solutions. This is because charge detection in high - ionic - strength environments is hindered by Debye screening, where charged molecules in an ionic solution attract counterions, forming a double - layer that effectively shields the charge on the molecule, i.e., a negatively charged molecule such as DNA is surrounded by cations via electrostatic interactions. Debye screening thus depends on the electrolyte concentration.

[0004] Under physiological conditions where the ionic strength > 100 mM, the Debye screening effect limits detection to within approximately 1 nm from the sensor surface. Thus, most FET - based biosensors operate under non - physiological ionic - strength conditions only by pre - desalting or diluting the sample. For applicability to PoC settings, ionic shielding must be reduced in a more direct and efficient way, as the sample - handling capacity is very limited at sites near the patient.

[0005] Several recent studies have reported several strategies to mitigate such Debye screening problems. It has been reported that using smaller receptors (such as aptamers or smaller antibody fragments) brings the target molecule closer to the electrode surface, thereby enhancing transistor-based detection of proteins. In addition, it has been demonstrated that by operating high-frequency FET measurements, the ion shielding effect can be mitigated.

[0006] Recently, it has been shown that co-immobilizing polyethylene glycol (PEG) on the sensor enables transistor-based detection of biomolecules in high ionic strength solutions (Gao 2015, Nano Lett. 15: 2143; Gao 2016, PNAS 113(51): 14633-14638 and WO2016 / 161246).

[0007] PEGs on electrode materials have also been described for different purposes. For example, they are described as linkers for probe immobilization and reducing non-specific binding (Shim 2002, Nano Lett. 2: 285). In addition, PEG can act as a stabilizer for Fab fragments (Yoshimoto 2010, J. Am. Chem. Soc. 132: 7982).

[0008] The technical problem underlying the present invention can be seen as providing means and methods meeting the above requirements. The said technical problem is solved by the claims and the embodiments characterized hereinafter.

[0009] Accordingly, the present invention relates to an electrode comprising a functionalized surface that can be exposed to a fluid sample, the functionalized surface comprising at least one polymer capable of mediating a salting-out effect and at least one detection agent that binds to an analyte contained in the fluid sample, wherein the at least one polymer capable of mediating a salting-out effect and the at least one detection agent are distributed on the surface of the electrode such that the detection agent is present in substantially equal amounts per surface area over the entire electrode surface, and the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that:

[0010] i) it is capable of reducing the ionic strength of the fluid in the vicinity of the detection agent, and

[0011] ii) it is capable of binding to an analyte contained in the fluid sample.

[0012] As used herein, the terms "having", "including", or "comprising" or any grammatical variations thereof are used in a non-exclusive manner. Thus, these terms can refer both to cases where no other features exist in the entity described in the context except for the features introduced by these terms, and to cases where one or more other features exist. For example, the statements "A has B", "A includes B", and "A comprises B" can refer both to cases where no other elements exist in A except for B (i.e., cases where A consists solely and exclusively of B), and to cases where one or more other elements such as element C, elements C and D, or even other elements exist in entity A in addition to B.

[0013] In addition, it should be noted that the terms "at least one", "one or more", or similar expressions (which mean that a feature or element may occur once or more than once, i.e., twice, three times, four times, five times up to an infinite number of times) are typically used only once when introducing the corresponding feature or element. In the following text, in most cases, when referring to the corresponding feature or element, the expressions "at least one" or "one or more" will not be repeated, even though the corresponding feature or element may occur once or more than once.

[0014] In addition, as used herein, the terms "preferably", "more preferably", "particularly", "more particularly", "specifically", "more specifically", "generally", "more generally", or similar terms are used in conjunction with optional features without restricting the possibility of alternatives. Thus, the features introduced by these terms are optional features and are not intended to limit the scope of the claims in any way. As those skilled in the art will recognize, the present invention can be implemented by using alternative features. Similarly, the features introduced by "in one embodiment of the present invention" or similar expressions are intended to be optional features, and not any limitation on alternative embodiments of the present invention, nor any limitation on the scope of the present invention, and not any limitation on the possibility of combining the features introduced in this way with other optional or non-optional features of the present invention.

[0015] The term "electrode" as used herein refers to a structure capable of making electrical contact. Electrodes are typically used for electrical contact with non-metallic parts of a circuit such as semiconductor elements, electrolytes, vacuum or air. Generally, an electrode comprises a conductive material. In particular, the electrodes used according to the present invention are solid carriers. The term "electrode" as used herein can generally refer to a functional element configured to perform current measurement and / or voltage measurement and / or configured to apply current and / or potential and / or voltage to an element in electrical contact with the electrode. In particular, an electrode can comprise a conductive and / or semiconductive material. As an example, an electrode can comprise at least one metallic material and / or at least one organic or inorganic semiconductive material having at least one conductive or semiconductive surface. This surface itself can constitute the electrode or a part of the electrode. As an example, an electrode can comprise at least one material, in particular at least one surface material, having a conductivity of at least 1000 S / m, for example at least 1000000 S / m, which is isotropic or anisotropic in at least one direction. Specifically, an electrode can comprise graphene, carbon nanotubes, silicon nanowires, molybdenum oxide, molybdenum disulfide, titanium oxide, zinc oxide, metal oxides, gallium nitride, gold, silicon, magnetic beads, nanoparticles or any combination of these materials.

[0016] Electrodes are typically used for electrical contact with non-metallic parts of a circuit such as semiconductor elements, electrolytes, vacuum or air. Generally, an electrode comprises a conductive material. In particular, the electrodes used according to the present invention are solid carrier structures.

[0017] The term "electrical contact" as used herein generally refers to the arrangement or configuration of at least two components, where at least one of the components is capable of electrically influencing at least one other component and / or at least partially controlling the electrical quality of another component, such as but not limited to its electrical conductivity and current flow, for example via electrostatic induction. In particular, an electrode can be in electrical contact with an element without directly physically contacting the element. Thus, even when isolated from the element, the electrode can still control the current flow in the element by applying a voltage. The isolation can be constituted, for example, by an oxide layer, which is typically the case for the gate of a metal-oxide-semiconductor field-effect transistor (MOSFET) - a subclass of insulated-gate field-effect transistor (IGFET), which will be described in more detail below. Thus, generally, in order to be in electrical contact with each other, the at least two components can be positioned in close proximity without directly physically contacting each other, but such that the components can electrically influence each other. However, in addition or alternatively, the at least two components can also be physically connected via at least one connecting element having at least a semiconductive or conductive nature, such as via at least one electrical conductor. Similarly, in addition or alternatively, the at least two components can be spaced components, or can be integrated with each other completely or partially. As an example, at least one electrode can be connected to a field-effect transistor via at least one connecting element (such as via at least one conductive lead), or can even be integrated completely or partially into the field-effect transistor.

[0018] The term "functionalized surface" as used herein refers to the surface of an electrode according to the present invention, which has certain desired physical and / or chemical properties. The surface of the electrode according to the present invention is functionalized in such a way that it contains at least one polymer capable of mediating a salting-out effect and at least one detection agent that binds to an analyte contained in a fluid sample. In addition, the polymer and the detection agent are distributed on the surface of the electrode such that the detection agent is present in substantially equal amounts per surface area over the entire electrode surface. Further, in the arrangement, the polymer capable of mediating the salting-out effect is present around the detection agent, and its amount is such that it can reduce the ionic strength of the fluid near the detection agent and can simultaneously bind to the analyte contained in the fluid sample. The arrangement can generally be a continuous layer, or can be a cluster arrangement, such as a dot-like arrangement around each detection agent.

[0019] The term "fluid sample" as used herein refers to any solution containing an analyte to be detected or suspected of containing such an analyte. The fluid sample can be an aqueous solution or can contain other solvents, including organic solvents. The fluid sample can have any source, for example, it can be a fluid sample of a naturally occurring fluid, including a fluid present in the environment or a fluid present in a living organism (i.e., a body fluid), or a fluid derived from a living organism, such as an extract. In addition, the fluid can be an artificial fluid, such as a fluid obtained by dissolving a compound to be analyzed in a suitable solvent, or a fluid obtained as a product of a chemical reaction. Generally, the fluid sample is a body fluid, a liquid or a dissolved environmental sample, or a solution of at least one chemical compound. More generally, the fluid is a high ionic strength fluid. The fluid sample thus contains a large amount of cations and anions, for example, from dissolved salts or buffers. Generally, it is to be understood that at high ionic strength, the ions are present in millimolar amounts, typically in amounts equal to or greater than 10 mM, 25 mM, 50 mM, 75 mM, 100 mM, 150 mM or 200 mM. Body fluids are generally selected from: blood, plasma, serum or any part thereof, saliva, tears, mucus, lymph fluid, cerebrospinal fluid, urine, feces, sweat, semen, synovial fluid.

[0020] As used herein, the term "analyte" refers to a molecule that may or may not be present in a fluid sample, the presence and / or amount of which will be detected. Depending on the type of detection agent to be used, analytes can be selected from a variety of molecules. In particular, analytes can be small molecules, peptides, proteins, oligonucleotides, polynucleotides such as RNA or DNA, polymers or other macromolecules, viruses or organisms, such as microorganisms including single-celled organisms such as bacteria, archaea, algal protozoa or fungi. However, generally analytes are small molecule compounds, such as substrates of enzymes in metabolic pathways, intermediates of such pathways, or products obtained through metabolic pathways. Thus, analytes according to the present invention are more typically likely to be metabolites. Metabolic pathways are well known in the art and can vary between species. Preferably, the pathways include at least the citric acid cycle, respiratory chain, photosynthesis, photorespiration, glycolysis, gluconeogenesis, hexose monophosphate pathway, oxidative pentose phosphate pathway, fatty acid production and β-oxidation, urea cycle, amino acid biosynthesis pathway, protein degradation pathways such as proteasomal degradation, amino acid degradation pathways, biosynthesis or degradation of the following: lipids, polyketides (including for example flavonoids and isoflavonoids), isoprenoids (including for example terpenes, sterols, steroids, carotenoids, xanthophylls), carbohydrates, phenylpropanoids and derivatives, alkaloids, benzene ring compounds, indoles, indole-sulfur compounds, porphyrins, anthocyanins, hormones, vitamins, cofactors such as prosthetic groups or electron carriers, lignin, glucosinolates, purines, pyrimidines, nucleosides, nucleotides and related molecules such as tRNA, microRNA (miRNA) or mRNA. Thus, small molecule compound metabolites generally consist of compounds of the following classes: alcohols, alkanes, alkenes, alkynes, aromatic compounds, ketones, aldehydes, carboxylic acids, esters, amines, imines, amides, cyanides, amino acids, peptides, thiols, thioesters, phosphates, sulfates, thioethers, sulfoxides, ethers or combinations or derivatives of the foregoing compounds. Small molecules in metabolites can be primary metabolites required for normal cell function, organ function or animal growth, development or health. In addition, small molecule metabolites further include secondary metabolites with essential ecological functions, such as metabolites that enable an organism to adapt to its environment. In addition, metabolites are not limited to the primary and secondary metabolites described above and further encompass artificial small molecule compounds. The artificial small molecule compounds are derived from exogenously provided small molecules that are taken or absorbed by an organism but are not primary or secondary metabolites as defined above. For example, artificial small molecule compounds can be metabolites obtained from drugs through the metabolic pathways of an animal.

[0021] As used herein, the term "polymer capable of mediating a salting-out effect" refers to a macromolecule comprising more than one monomer subunit. Generally, the monomer subunits of the polymer are chemically linked to each other, thereby forming the macromolecule. The polymer may be composed of chemically identical monomers, i.e., a homopolymer, or may be composed of chemically different monomers, i.e., a heteropolymer. Due to its chemical architecture, the polymer has various properties resulting from the monomer subunits and their properties as well as the macromolecular arrangement of the monomer subunits. The polymers contemplated by the present invention are capable of mediating a salting-out effect, i.e., when appropriately arranged (e.g., in the form of a layer), they will reduce the ionic strength of the surrounding ionic fluid in the polymer layer. Generally, the polymer achieves this by enriching the solvent (especially water molecules) in its vicinity while repelling ions away from the polymer. Since the ionic strength in the solvent in the region around the polymer is reduced, the electrostatic effects caused by the dissolved ions will be reduced near the polymer molecules. Therefore, if appropriately arranged (e.g., in the form of a layer on an electrode), the Debye length between the analyte and the electrode in the fluid can be increased. The Debye length of an ionic solution can be determined by the following formula:

[0022]

[0023] where:

[0024] is the Debye length;

[0025] is the dielectric constant;

[0026] is the permittivity of free space;

[0027] is the Boltzmann constant;

[0028] T is the temperature;

[0029] is the Avogadro number;

[0030] e is the elementary charge;

[0031] I is the ionic strength.

[0032] The ionic strength and Debye length in the fluid can be determined by a person skilled in the art without much ado. In addition, the effect of the polymer on the Debye length can be determined by a person skilled in the art by techniques well known in the art and described in the examples attached below.

[0033] Typically, the size of the at least one polymer capable of mediating a salting-out effect is such that the Debye length is increased and the ionic strength around at least one detection molecule is decreased. More typically, the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is from 1 to 100 kDa, from 10 to 100 kDa, from 10 to 50 kDa, from 10 to 25 kDa, from 10 to 20 kDa or from 10 to 15 kDa.

[0034] In addition, the at least one polymer capable of mediating a salting-out effect is typically selected from: poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), copolymers of the foregoing polymers, polysaccharides, polypeptides, polynucleotides, and polysiloxanes.

[0035] The foregoing polymer capable of mediating a salting-out effect will be immobilized on the surface of the electrode. Thus, the polymer can be directly or indirectly (i.e., via a linker) attached to the surface of the electrode. The type of attachment used depends on the electrode material and the polymer. Thus, a direct attachment between the polymer and the electrode can be achieved if appropriately selected. However, the attachment is typically an indirect attachment via a linker. A linker as mentioned in the present invention is a molecule, more typically a bifunctional molecule, which allows simultaneous direct binding to both the electrode surface and the polymer. The binding mentioned herein can be covalent or non-covalent and can thus be permanent or reversible depending on the surrounding environment and the conditions employed. It is well known to those skilled in the art how and under what conditions direct or indirect binding can be achieved and what type of linker is suitable.

[0036] Typically, the polymer capable of mediating a salting-out effect is immobilized on the surface of the electrode via a linker or directly attached via a functional group without an additional linker. More typically, the linker is a linker having the following structure:

[0037] A - B - C

[0038] where A is a first functional group selected from the group consisting of: thiol, silane, phosphonic acid, aromatic molecule (such as pyrene), carboxyl, amine, NHS ester, maleimide,

[0039] Wherein B is a short organic chain selected from the group consisting of: low molecular weight polymers selected from the above groups, hydrocarbons (such as alkyl, alkenyl, alkynyl, phenyl), oxygen-containing groups (such as ether), haloalkanes (such as chloro), nitrogen-containing groups (such as amide), sulfur-containing groups (such as sulfoxide), phosphorus-containing groups, boron-containing groups or short inorganic polymer chains (such as Si-based (siloxane), P-based, B-based, S-based) and wherein C is a second functional group selected from the group consisting of: carboxyl, amine, NHS ester, maleimide.

[0040] The term "detection agent" as used in the present invention refers to a molecule that binds to an analyte to be detected when present in a fluid sample. Generally, the detection agent binds specifically and selectively to the analyte, i.e., does not cross-react with other analytes that may or may not be present in the sample. Specific binding can be tested by various well-known techniques. Depending on the type of analyte to be detected, the detection agent can be selected from different molecular classes. It is well known to those skilled in the art which type of analyte can be detected by which molecule used as a detection agent. Generally, the detection agent is selected from: antibodies and their fragments, nucleic acids, aptamers, peptide nucleic acids (PNA), receptor or ligand proteins or peptides, and enzymes.

[0041] The antibodies and their fragments mentioned herein encompass all types of antibodies that typically bind specifically to the analyte to be detected. Generally, the antibodies according to the present invention are monoclonal antibodies, polyclonal antibodies, single-chain antibodies, chimeric antibodies or any fragment or derivative of such antibodies that are still capable of binding the analyte. Such fragments and derivatives encompassed by the term antibody as used herein include bispecific antibodies, synthetic antibodies, Fab, F(ab) 2 ', Fv or scFv fragments, or chemically modified derivatives of any of these antibodies. Antibodies or their fragments can generally be obtained by using methods well known in the art. Monoclonal antibodies can be prepared, for example, by techniques including fusing mouse myeloma cells to spleen cells derived from an immunized mammal (preferably an immunized mouse) according to the Köhler & Millstein technique.

[0042] The nucleic acids mentioned herein refer to all types of deoxyribonucleic acid (DNA) and ribonucleic acid (RNA) and their chemically modified derivatives. These molecules are well known in the art. Generally, nucleic acids as detection agents can be used to detect other nucleic acids as analytes. In such cases, the detection agent nucleic acid is partially or completely complementary to the analyte nucleic acid to be detected or a portion thereof. Generally, such nucleic acids can have the size of oligonucleotides, i.e., a length of 5 to 35 nucleotides, more preferably a length of 10 to 25 nucleotides, or can be larger nucleic acid probes with a length of 100 to 1,000 nucleotides, more preferably a length of 300 to 600 nucleotides.

[0043] The aptamers referred to in the present invention include nucleic acid aptamers and peptide aptamers. In addition to their ability to base pair, nucleic acids can also be used as aptamers for detecting other analytes due to their ability to form three-dimensional structures that specifically bind to target molecules such as small molecules, proteins, nucleic acids, and even cells, tissues, and organisms. Nucleic acid aptamers can be engineered to bind various molecular targets through multiple rounds of in vitro selection or through systematic evolution of ligands by exponential enrichment (SELEX) technology. Peptide aptamers are artificial peptides that are selected or engineered to bind specific target molecules. These peptides typically consist of one or more peptide loops of variable sequences displayed by a protein scaffold. They are usually isolated from combinatorial libraries and are typically subsequently improved by directed mutagenesis or multiple rounds of mutagenesis and selection of the variable regions.

[0044] Peptide nucleic acid (PNA) is a synthetic polymer that has a nucleic acid-like backbone consisting of repeating N-(2-aminoethyl)-glycine units linked by peptide bonds. Various purine and pyrimidine bases are attached to the backbone by methylene bridges and carbonyl groups. PNA and nucleic acids have similar biological properties, and thus, PNA can be used as a detection agent like nucleic acids or aptamers.

[0045] The receptor or ligand proteins or peptides according to the present invention refer to proteins or peptides that can specifically recognize other proteins or peptides. Generally, receptor and ligand peptides or proteins can specifically interact with other molecules such as other proteins or peptides. Thus, the receptor or ligand can be used as a detection agent for such interacting proteins or peptides or even other molecules that interact with them. It is to be understood that according to the present invention, the receptor or ligand proteins or peptides can also comprise parts of the whole biologically active receptor or ligand and generally include parts that contain their binding domains. Receptor or ligand proteins and peptides can be naturally occurring receptors or artificially produced receptors. Typical artificial peptides as detection agents also include cyclic peptides.

[0046] Also generally suitable as detection agents are enzymes. Enzymes are proteins or peptides that specifically bind to a molecule (substrate) and are capable of enzymatically converting the molecule into other molecules (products). Thus, enzymes can generally specifically bind to substrates and can thereby be used to detect such substrates present as analytes in fluid samples. The analytes recognized by enzymes generally include small molecules, peptides, or proteins. However, some enzymes can also recognize large molecules such as polymers. Suitable enzymes and their substrates are well known in the art.

[0047] Typically, the at least one detection agent is immobilized on the surface of the electrode via a linker. More typically, the linker is selected from: low molecular weight (MW) polymers having an MW of 0.01 to 5, 0.01 to 1.0, 0.01 to 0.5, or 0.1 to 0.5 kDa, and the polymers are typically selected from: poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), and copolymers of the foregoing polymers. Using such low molecular weight polymers will enhance the conformational stability, flexibility, and capacity of the binding of the detection agent to the analyte. It will also reduce non-specific adsorption of species other than the analyte. Alternatively, the linker for the detection agent can also be the linker as described above for the polymer.

[0048] In particular, it is contemplated according to the present invention that the linker for the detection agent is a polymer as described above. Additionally, it is particularly contemplated that the polymer capable of mediating a salting-out effect is of the same class as the polymer serving as the linker polymer, i.e., both are poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), or a copolymer of the foregoing polymers. More typically, the linker polymer is a low molecular weight (MW) polymer having an MW of 0.01 to 5, 0.01 to 1.0, 0.01 to 0.5, or 0.1 to 0.5 kDa, while the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is 1 to 100 kDa, 10 to 100 kDa, 10 to 50 kDa, 10 to 25 kDa, 10 to 20 kDa, or 10 to 15 kDa.

[0049] The polymer and the detection agent will be distributed on the surface of the electrode such that the detection agent is present in substantially equal amounts per surface area over the entire electrode surface. It is to be understood that for performing analyte detection measurements using the electrode of the present invention, it is advantageous for the detection agent to be uniformly present over the entire surface of the electrode. Such uniform presence can be achieved by applying the detection agent or the linker for the detection agent and the polymer to the surface of the electrode in the form of a homogeneous solution. Thereby, all areas of the electrode will statistically receive substantially equal amounts of the detection agent or the linker and the polymer. How such uniform coating can be achieved is also well established in the art and is described in more detail elsewhere herein.

[0050] In addition, the polymer should be arranged around the detection agent in an amount such that it can reduce the ionic strength of the fluid near the detection agent but still be able to bind to the analyte contained in the fluid sample. Generally, the detection agent is surrounded by polymer molecules such that the detection agent cannot freely contact the fluid sample. This enables the polymer to reduce the ionic strength near the detection agent, increasing the Debye length. It should be understood that attention should be paid to the fact that the size of the detection agent and the size of the polymer are such that the detection agent is within the polymer layer or cluster on the surface and cannot approach or can only approach the free fluid sample limitedly. Thus, the size of the polymer is greater than or equal to the size of the detection agent, or in other words, the detection agent is preferably embedded in the polymer layer or cluster. In the following table, the detection agents and the polymers that can be used for coating are listed:

[0051] Table 1: Polymers for individual detection agents

[0052] Detection agent The thickness of the polymer layer Available polymers dsDNA (e.g., with a length of 15 nt) Equal to or greater than 5 nm 5, 10, 20, 30, 40 kDa PEG Nanobody Equal to or greater than 4 nm 5, 10, 20, 30, 40 kDa PEG Antibody fragment Fab Equal to or greater than 7 nm 10, 20, 30, 40 kDa PEG Antibody fragment F(ab)2 Equal to or greater than 7 nm 10, 20, 30, 40 kDa PEG Monoclonal antibody Equal to or greater than 15 nm 20, 30, 40 kDa PEG

[0053] The thickness of the polymer layer and the detection agent layer can be measured, calculated, and / or predicted by techniques well known in the art. Thus, when considering the arrangement of the polymer and the detection agent as separate layers, based on prediction and / or calculation, a polymer of appropriate size can be selected for a given detection agent without further hassle.

[0054] In addition, the detection agent molecules are generally completely surrounded by polymer molecules on the surface. To achieve an appropriate arrangement for analyte detection, the detection agent and / or its linker are applied simultaneously with the at least one polymer. Generally, the molar ratio of at least one polymer capable of mediating the salting-out effect present on the functionalized surface to at least one detection agent is 1:100 to 100:1, 1:50 to 50:1, 1:20 to 20:1, 1:10 to 10:1, 1:5 to 5:1, 2:10 to 8:1, 3:10 to 7:1, 4:10 to 6:1, or 5:10 to 5:1. Such molar ratios enable the aforementioned favorable arrangement of the detection agent and the polymer to be formed on the surface of the electrode. In particular, a suitable layer or cluster arrangement of the polymer and the detection agent can be obtained by applying the polymer and the linker (usually a polymer linker) in the form of a solution in a suitable solvent to the electrode. The solution generally contains the polymer and the linker in a molecular weight ratio of 1:10 to 1:50, more usually 1:20. It should be understood that for larger detection agents (such as antibodies, receptors, or enzymes), a larger molecular ratio is used, i.e., the amount of the linker is increased.

[0055] Advantageously, in the studies on which the present invention is based, the degree of increase in the Debye length was quantified by systematically comparing double-stranded DNA (dsDNA) detection at different salt concentrations in the presence and absence of PEG. dsDNA is an ideal biomolecule for studying the effect of Debye screening on FET-based measurements because it has a uniform surface charge and an easily adjustable length. In addition, it has been shown that PEG immobilized with dsDNA can locally desalt the region adjacent to the sensor surface, reducing the local ionic strength by at least a factor of 10. This results in a reduction of the Debye screening effect in the vicinity around the dsDNA molecule and enables improved FET-based detection in physiological ionic strength solutions. Thereby, according to the present invention, it has been found that a polymer capable of mediating a salting-out effect surrounding a detection agent on an electrode surface can increase the Debye length of the electrode measurement. Thus, measurements can also be effectively carried out in solutions with high ionic strength, which would typically impede field-effect-based detection of analytes in such solutions. In particular, it has been found that due to salting out, the ionic strength of the solution in the polymer layer is reduced. However, the detection agent immobilized on the electrode must be surrounded in the spatial arrangement of the polymer molecules, which enables the analyte to bind to the detection molecule and at the same time enables the reduction of the ionic strength in the layer as described. The proper arrangement of the polymer and the detection agent can be achieved by using a predetermined mixture of the detection agent or a linker for the detection agent and the polymer molecule or a linker for the polymer. Thanks to the findings on which the present invention is based, a method for manufacturing a functionalized electrode surface is also provided and is described in more detail elsewhere herein. It is particularly advantageous to use low molecular weight polymers as linkers for detection agents such as antibodies, peptides, receptors or enzymes, because the conformational stability, flexibility and capacity of the binding of the analyte to the detection agent will be increased and non-specific adsorption of species other than the analyte will be reduced.

[0056] All definitions and explanations given for the above terms apply, mutatis mutandis in detail, to all the following embodiments.

[0057] The present invention further relates to a method for manufacturing a functionalized surface on an electrode, comprising the following steps:

[0058] a) applying the linker and the polymer to the electrode under conditions allowing covalent or non-covalent immobilization of the linker and at least one polymer capable of mediating a salting-out effect on the surface of the electrode; and

[0059] b) applying the at least one detection agent to the electrode on which the linker and the polymer are immobilized under conditions allowing covalent or non-covalent attachment of at least one detection molecule to the electrode via the immobilized linker; and

[0060] wherein the conditions allow for the distribution of the at least one polymer capable of mediating a salting-out effect and the at least one detection agent on the surface of the electrode such that the detection agent is present in an equal amount per surface area across the entire electrode surface and the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that:

[0061] i) it is capable of reducing the ionic strength of the fluid in the vicinity of the detection agent, and

[0062] ii) it is capable of binding an analyte contained in the fluid sample.

[0063] The linker and the polymer are applied to the electrode under conditions that allow for covalent or non-covalent immobilization of the linker and the at least one polymer capable of mediating a salting-out effect on the surface of the electrode. Thus, depending on the type of immobilization of the molecule, the molecule can be covalently linked to the electrode surface or attached by a non-covalent mechanism. Generally, covalent binding of the linker or polymer can be achieved through functional groups present in the linker or polymer molecule. The functional groups are capable of forming a covalent bond with the electrode surface. The functional groups can also be present in the polymer to allow its covalent attachment to the electrode. Generally, non-covalent binding of the linker or polymer can also be achieved through functional groups present in the linker or polymer molecule. The functional groups are capable of forming a non-covalent bond with the electrode surface, for example via electrostatic interactions, hydrophobic interactions, π-interactions, hydrogen bonds or van der Waals forces. Typical functional groups can be thiols, silanes, phosphonic acids, aromatic molecules (such as pyrene), carboxyl groups, amines, NHS esters, maleimides.

[0064] Generally, the at least one polymer capable of mediating a salting-out effect for the method of the present invention is selected from: poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), and copolymers of the foregoing polymers.

[0065] Generally, the linker for the method of the present invention is a low molecular weight (MW) polymer having an MW of 0.01 to 5, 0.01 to 1.0, 0.01 to 0.5 or 0.1 to 0.5 kDa, and the polymer is poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), or a copolymer of the foregoing polymers.

[0066] The linker and the polymer are applied under conditions that allow a spatial arrangement of the linker and the polymer to form, which reflects the envisioned arrangement of the detection molecule and the polymer. This can be achieved by applying a predetermined mixture of the molecules to the surface of the electrode. Typically, the linker and the polymer molecules are mixed in a predetermined molar ratio in the application solution. In a suitable solvent, the molar ratio of the polymer to the linker for the detection agent is typically 1:100 to 100:1, 1:50 to 50:1, 1:20 to 20:1, 1:10 to 10:1, 1:5 to 5:1, 2:10 to 8:1, 3:10 to 7:1, 4:10 to 6:1, or 5:10 to 5:1. The solvent is in contact with the surface of the electrode, and conditions are applied that allow the linker and the polymer to be fixed on the surface. In addition, the linker and the polymer are typically applied such that they are present in equal amounts per surface area over the entire surface of the electrode.

[0067] In particular, the linker for the detection agent is a low molecular weight polymer as described above, and more typically a polymer of the same class as the polymer capable of mediating a salting-out effect, i.e., both are poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), or a copolymer of the foregoing polymers. In the method of the present invention, a suitable layer arrangement of the polymer and the detector can typically be obtained by applying the polymer and the foregoing polymer linker in the form of a solution in a suitable solvent, the solution containing the polymer and the linker in a molecular weight ratio of 1:10 to 1:50, more typically 1:20. Advantageously, molecules of the same polymer class are used as the linker and the polymer capable of mediating a salting-out effect, since the same immobilization reaction can be used to immobilize the polymer and the polymer linker onto the surface of the electrode.

[0068] In the next step, the at least one detection molecule is applied under conditions that allow the at least one detection molecule to be covalently or non-covalently linked to the electrode via the immobilized linker. The linker typically contains other functional groups capable of forming a covalent or non-covalent bond with the detection molecule. Typical functional groups include those specified elsewhere herein. Once the detection agent is applied to the electrode containing the linker and the polymer molecules on its surface, the envisioned spatial arrangement is created. In particular, the at least one polymer capable of mediating a salting-out effect and the at least one detection agent are present in equal amounts per surface area over the entire surface of the electrode. In addition, the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that it can reduce the ionic strength of the fluid near the detection agent and can bind the analyte contained in the fluid sample.

[0069] The present invention also provides an analyte detector for determining at least one analyte, which comprises the electrode of the present invention or an electrode obtainable by the method of the present invention, wherein the electrode is in electrical contact with a transducer or is part of a transducer.

[0070] As used herein, the term "analyte detector" refers to a device adapted to detect at least one analyte in a fluid sample. The analyte detector comprises the electrode of the present invention as an analytical component. In addition, the device generally comprises a loading element for introducing the fluid sample into the device and a contact element for allowing the electrode to contact the fluid sample. Further, the electrode is in electrical contact with a transducer or is part of such a transducer. Generally, the analyte device according to the present invention comprises a field effect transistor (FET), and the electrode of the present invention is used as a gate or as a channel in the transistor. When used as a gate or as a channel, the electrode of the present invention allows for the specific and effective detection of analytes in a fluid sample, for example, according to the analyte detector of the present invention.

[0071] Alternatively or in addition, the analyte device according to the present invention may comprise an electrochemical measurement device, and the electrode of the present invention may participate in an electrochemical measurement. The term "electrochemical measurement device" generally may refer to any device configured to perform at least one electrochemical measurement. For this purpose, the at least one electrochemical measurement device may comprise one or more electrical devices configured to perform at least one electrochemical measurement. As an example, the electrochemical measurement device may comprise at least one electrode, at least one power source, such as at least one power source selected from the following: a constant voltage power source, a variable voltage power source, a constant current power source, a variable current power source, a frequency generator for generating a periodic electrical signal. In addition, the electrochemical measurement device may comprise at least one electrical measurement device configured to measure at least one electrical signal or electrical measurement variable, such as at least one electrical measurement device selected from the following: a voltage measurement device, a current measurement device, a potentiostat. Other measurement devices are also feasible.

[0072] As used herein, the term "electrochemical measurement" generally may refer to the measurement of at least one measurable property of a redox reaction. As an example, the measurable property of the electrochemical measurement and / or the redox reaction may refer to current, voltage, potential, mass (e.g., the mass deposited on an electrode), impedance (in particular the real part and / or the imaginary part of the impedance).

[0073] Specifically, electrochemical measurements can be carried out in the presence of electroactive species. The term "electroactive species" as used herein generally can refer to, for example, compounds that facilitate or enhance or catalyze redox reactions by promoting electron transfer. Electroactive species can be dissolved in a fluid sample and / or can be immobilized on the surface of an analyte detector, where the surface is exposed to the fluid sample. In particular, the surface can be the aforementioned sensing surface and / or the aforementioned surface of the multipurpose electrode. Preferred examples of electroactive species are redox mediators, especially redox couples, such as but not limited to: potassium ferricyanide / potassium ferrocyanide; hexaammine ruthenium(II) chloride / hexaammine ruthenium(III) chloride; ferrocenemethanol. Further preferred examples of electroactive species are reducing agents, such as but not limited to ascorbic acid, glutathione, lipoic acid, uric acid, oxalic acid, tannins, and phytic acid. Electroactive species can facilitate or enhance the measurement of at least one measurable property of a redox reaction.

[0074] In addition or alternatively, electrochemical measurements can refer to directly or indirectly detecting the electrical properties of an element (such as an electrode), where the electrical properties are affected or acted upon by a chemical reaction and / or by electron transfer and / or by the binding of atoms or molecules. Specifically, the electrical properties may be affected or acted upon by a chemical reaction that includes a change in the oxidation state of at least one of its participants. The electrical properties can be, for example, the electrochemical potential of the element and / or a change in the electrochemical potential, the potential of the element and / or a change in the potential, the voltage applied to the element and / or a change in the voltage, and / or the amount of charge accumulated on the element. The direct or indirect measurement of the electrical properties of the element can be based on the electric field effects caused and / or affected by the electrical properties of the element. Thus, as a specific example, a chemical reaction can change the oxidation state of an element. The change in the oxidation state of the element can be measured in an electrochemical measurement via an electric field effect (possibly caused by the element's oxidation state). Thus, detecting a change in the oxidation state of an element can be done, for example, using a field effect transistor. In particular, the electrical properties of the element can affect the current between the source electrode and the drain electrode of a field effect transistor by affecting the gate voltage of the field effect transistor.

[0075] As used herein, the term "transducer" refers to any type of functional component or arrangement of functional components configured to convert energy from one form to another and / or configured to convert an input signal (in particular, an electrical signal such as an applied current, voltage, or potential) into a corresponding output signal (where the form of the output signal is different from the form of the input signal). Thus, a transducer can be configured to generate an electrical signal of another form (e.g., a current) as an output signal after receiving an electrical signal (e.g., an applied voltage or potential) as an input signal. Specifically, the transducer can be or can include a field effect transistor and / or an electrochemical measurement device. In the case where the transducer functions as a field effect transistor, the input signal can be a voltage or potential applied through a gate to a channel of the field effect transistor, which can be changed by binding charged species on the surface of the sensor, and the output signal can be a drain current I d between the source and drain electrodes of the field effect transistor. In the case where the transducer functions as an electrochemical measurement device, the input signal can be at least one measurable characteristic of a redox reaction.

[0076] The present invention generally contemplates using the analyte detector of the present invention to determine at least one analyte in a fluid sample. Typically, the determination of at least one analyte in a fluid involves diagnostic purposes, environmental monitoring and control, food safety, quality control, or manufacturing processes.

[0077] The detection of analytes typically plays a role in a variety of different processes. It can be used to diagnose changes in analytes in a subject, which in turn can help diagnose diseases or other medical conditions. In addition, the analysis of analytes can be used to monitor the environment, such as for detecting changes in the degree of pollution. However, production, food safety, and general quality control processes typically require the detection of analytes. Thus, the analyte detector according to the present invention can be used in any such process and generally allows for its further automation.

[0078] In addition, the present invention contemplates a method for determining at least one analyte, comprising the steps of:

[0079] (a) contacting a fluid sample suspected of containing at least one analyte with an electrode of the present invention or the analyte detector of the present invention; and

[0080] (b) performing an electrochemical measurement with the electrode or detector, thereby determining at least one analyte.

[0081] As used herein, the term "determination" refers to quantitative determination, i.e., the determination of quantity, as well as qualitative determination, i.e., the determination of the presence or absence of an analyte. Such determinations can be made using, for example, the analyte detector of the present invention, by performing electrochemical measurements as described elsewhere herein. Based on the results of the electrochemical measurements, at least one analyte can be determined as described above.

[0082] Electrochemical measurements can in principle be carried out under standard conditions that allow the analyte to bind to the detector on the electrode. Typically, such standard conditions include a temperature above the freezing point and below the boiling point of the fluid sample. In some applications, the temperature will be within the range of room temperature. However, electrochemical measurements are more typically carried out at a temperature of from 30 °C to 40 °C, preferably at a temperature of at least 30 °C, at least 32 °C, at least 35 °C or at least 37 °C. In the studies on which the present invention is based, it has been found that if the measurement is carried out at about 37 °C, a strong signal enhancement can be achieved, resulting in a detection limit that is three orders of magnitude lower compared to 21 °C.

[0083] For example, electrochemical measurements can be carried out in the presence of electroactive species in order to enhance the signal. The electroactive species can be dissolved in the analyte solution or immobilized on a surface. Typical electroactive species are: redox couples such as potassium ferricyanide / potassium ferrocyanide, hexammine ruthenium (II) chloride and hexammine ruthenium (III) chloride, ferrocenemethanol; reducing agents such as ascorbic acid, glutathione, lipoic acid, uric acid, oxalic acid, tannins, phytic acid. Typically, electrochemical measurements or transistor-based measurements are carried out in the presence of a secondary receptor that binds to the analyte on the surface in order to enhance the signal. The secondary receptor can enhance the signal and / or selectivity itself, or can be labeled with additional molecules such as enzymes. For example, the signal enhancing molecule can cause a change in the concentration of a species (e.g. a proton or an electron) whose concentration can be directly measured by the sensor through its interaction with the substrate.

[0084] The present invention further discloses and provides a computer program that includes computer-executable instructions for implementing the foregoing method according to the present invention in one or more of the embodiments disclosed herein when the program is executed on a computer or a computer network. Specifically, the computer program can be stored on a computer-readable data carrier. Thus, specifically, one, more than one or even all of the method steps c) and d) described above can be implemented and / or controlled and / or evaluated by using a computer or a computer network, preferably by using a computer program.

[0085] The present invention further discloses and provides a computer program product having program code means for implementing the foregoing method according to the present invention in one or more of the embodiments disclosed herein when the program is executed on a computer or a computer network. Specifically, the program code means can be stored on a computer-readable data carrier.

[0086] Furthermore, the present invention discloses and provides a data carrier on which a data structure is stored, which, after being loaded into a computer or a computer network, such as into the working memory or main memory of a computer or a computer network, can execute the foregoing method according to one or more embodiments disclosed herein.

[0087] The present invention further provides and discloses a computer program product having program code means stored on a machine-readable carrier for implementing the foregoing method according to one or more embodiments disclosed herein when the program is executed on a computer or a computer network. As used herein, a computer program product refers to a program as a salable product. The product can generally exist in any form, such as in paper form, or on a computer-readable data carrier. Specifically, the computer program product can be distributed over a data network.

[0088] Finally, the present invention provides and discloses a modulated data signal containing instructions readable by a computer system or a computer network for implementing the foregoing method according to one or more embodiments disclosed herein.

[0089] Preferably, with reference to the computer-implemented aspects of the present invention, one or more method steps or even all method steps of the foregoing method according to one or more embodiments disclosed herein can be implemented by using a computer or a computer network. Thus, generally, any method step including providing and / or operating data can be implemented by using a computer or a computer network. Generally, these method steps can include any method steps, generally except for method steps that require manual work (such as certain aspects of providing samples and / or performing actual measurements).

[0090] Specifically, the present invention further discloses:

[0091] - A computer or a computer network comprising at least one processor, wherein the processor is adapted to implement the method according to one of the above embodiments,

[0092] - A computer-loadable data structure adapted to implement the method according to one of the above embodiments when the data structure is executed on a computer,

[0093] - A computer program, wherein the computer program is adapted to implement the foregoing method according to one of the embodiments described above in this specification when the program is executed on a computer,

[0094] - A computer program comprising program means for implementing the foregoing method according to one of the embodiments described above in this specification when the computer program is executed on a computer or a computer network,

[0095] - A computer program comprising a program tool according to the foregoing embodiments, wherein the program tool is stored on a computer-readable storage medium,

[0096] - A storage medium on which a data structure is stored, and wherein the data structure is adapted to implement the foregoing method according to one of the above embodiments after being loaded into the main memory and / or working memory of a computer or a computer network, and

[0097] - A computer program product having program code means, wherein the program code means can be stored or is stored on a storage medium so as to implement the foregoing method according to one of the above embodiments when the program code means is executed on a computer or on a computer network.

[0098] The following exemplary embodiments further illustrate the invention, but are not to be construed in any way as limiting:

[0099] 1. An electrode comprising a functionalized surface that can be exposed to a fluid sample, the functionalized surface comprising at least one polymer capable of mediating a salting-out effect and at least one detection agent that binds to an analyte contained in the fluid sample, wherein the at least one polymer capable of mediating a salting-out effect and the at least one detection agent are distributed on the surface of the electrode such that the detection agent is present in substantially equal amounts per surface area over the entire electrode surface, and the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that:

[0100] i) it can reduce the ionic strength of the fluid near the detection agent, and

[0101] ii) it can bind to an analyte contained in the fluid sample.

[0102] 2. The electrode of embodiment 1, wherein the electrode comprises graphene, carbon nanotubes, carbon, silicon nanowires, molybdenum oxide, molybdenum disulfide, titanium oxide, zinc oxide, metal oxides, gallium nitride, gold, silver, platinum, silicon, magnetic beads, nanoparticles, or any combination of these materials.

[0103] 3. The electrode of embodiment 1 or 2, wherein the size of the at least one polymer capable of mediating a salting-out effect is such that the Debye length is increased and the ionic strength around the at least one detection molecule is reduced.

[0104] 4. The electrode of embodiment 3, wherein the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is 1 to 100 kDa, 10 to 100 kDa, 10 to 50 kDa, 10 to 25 kDa, 10 to 20 kDa, or 10 to 15 kDa.

[0105] 5. The electrode according to any one of embodiments 1 to 4, wherein the at least one polymer capable of mediating a salting-out effect is selected from: poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), copolymers of the foregoing polymers, polysaccharides, polypeptides, polynucleotides, and polysiloxanes.

[0106] 6. The electrode according to any one of embodiments 1 to 5, wherein the at least one polymer capable of mediating a salting-out effect is immobilized on the surface of the electrode via a linker or directly attached via a functional group without an additional linker.

[0107] 7. The electrode of embodiment 6, wherein the linker is a linker having the following structure:

[0108] A - B - C

[0109] wherein A is a first functional group selected from the following categories: thiol, silane, phosphonic acid, aromatic molecule (such as pyrene), carboxyl, amine, NHS ester, maleimide,

[0110] wherein B is a short organic chain selected from the following categories: low molecular weight polymers selected from the above categories, hydrocarbons (such as alkyl, alkenyl, alkynyl, phenyl), oxygen-containing groups (such as ether), haloalkanes (such as chloro), nitrogen-containing groups (such as amide), sulfur-containing groups (such as sulfoxide), phosphorus-containing groups, boron-containing groups, or short inorganic polymer chains (such as Si-based (siloxane), P-based, B-based, S-based) and

[0111] wherein C is a second functional group selected from the following categories: carboxyl, amine, NHS ester, maleimide.

[0112] 8. The electrode according to any one of embodiments 1 to 7, wherein the at least one detection agent specifically binds to an analyte contained in a fluid sample.

[0113] 9. The electrode of any one of embodiment 8, wherein the at least one detection agent is selected from: antibodies and their fragments, nucleic acids, aptamers, peptide nucleic acids (PNA), receptor or ligand proteins or peptides, and enzymes.

[0114] 10. The electrode according to any one of embodiments 1 to 9, wherein the size of at least one detection agent is such that the detection agent is surrounded by at least one polymer capable of mediating a salting-out effect.

[0115] 11. The electrode according to any one of embodiments 1 to 10, wherein the at least one detection agent is immobilized on the surface of the electrode via a linker.

[0116] 12. The electrode of embodiment 11, wherein the linker is the linker as defined in claim 7.

[0117] 13. The electrode of embodiment 11, wherein the linker is selected from: low molecular weight (MW) polymers having an MW of 0.01 to 5, 0.01 to 1.0, 0.01 to 0.5 or 0.1 to 0.5 kDa, said polymers being poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP) and copolymers of the foregoing polymers.

[0118] 14. The electrode of embodiment 13, wherein the polymer capable of mediating a salting-out effect is of the same class as the polymer that is the linker polymer.

[0119] 15. The electrode of any one of embodiments 1 to 14, wherein the molar ratio of at least one polymer capable of mediating a salting-out effect present on the functionalized surface to at least one detection agent is from 1:100 to 100:1, 1:50 to 50:1, 1:20 to 20:1, 1:10 to 10:1, 1:5 to 5:1, 2:10 to 8:1, 3:10 to 7:1, 4:10 to 6:1 or 5:10 to 5:1.

[0120] 16. A method of fabricating a functionalized surface on an electrode, comprising the steps of:

[0121] a) applying the linker and the polymer to the electrode under conditions that permit covalent or non-covalent immobilization of the linker and at least one polymer capable of mediating a salting-out effect on the surface of the electrode; and

[0122] b) applying the at least one detection agent to the electrode on which the linker and the polymer are immobilized under conditions that permit covalent or non-covalent attachment of at least one detection molecule to the electrode via the immobilized linker; and

[0123] wherein the conditions permit distribution of the at least one polymer capable of mediating a salting-out effect and the at least one detection agent on the surface of the electrode such that the detection agent is present in an equal amount per surface area over the entire electrode surface and the polymer capable of mediating a salting-out effect is arranged around the detection agent in an amount such that:

[0124] i) it is capable of reducing the ionic strength of the fluid in the vicinity of the detection agent, and

[0125] ii) it is capable of binding an analyte contained in a fluid sample.

[0126] 17. The method of embodiment 16, wherein the linker is a low molecular weight (MW) polymer having an MW of 0.01 to 5, 0.01 to 1.0, 0.01 to 0.5, or 0.1 to 0.5 kDa, and the polymer is poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), or a copolymer of the foregoing polymers.

[0127] 18. The method of embodiment 16 or 17, wherein the at least one polymer capable of mediating a salting-out effect is selected from: poly(ethylene glycol) (PEG), poly(ethylene oxide), poly(propylene glycol), polyglycerol, polyacrylamide (PAM), polyethyleneimine (PEI), polymethacrylate or another acrylic polymer, poly(vinyl alcohol) (PVA), poly(vinylpyrrolidone) (PVP), and copolymers of the foregoing polymers.

[0128] 19. The method of embodiment 17 or 18, wherein the polymer capable of mediating a salting-out effect has the same class as the polymer used as the linker polymer.

[0129] 20. The method of any one of embodiments 16 to 19, wherein the molar ratio of the at least one polymer capable of mediating a salting-out effect present on the functionalized surface to the at least one detection agent is 1:100 to 100:1, 1:50 to 50:1, 1:20 to 20:1, 1:10 to 10:1, 1:5 to 5:1, 2:10 to 8:1, 3:10 to 7:1, 4:10 to 6:1, or 5:10 to 5:1.

[0130] 21. An analyte detector for determining at least one analyte, comprising the electrode of any one of embodiments 1 to 15 or an electrode obtainable by the method of any one of embodiments 16 to 20, wherein the electrode is in electrical contact with a transducer or is part of a transducer.

[0131] 22. Use of the analyte detector of embodiment 21 for determining at least one analyte in a fluid sample.

[0132] 23. A method for determining at least one analyte, comprising the steps of:

[0133] (a) contacting a fluid sample suspected of containing at least one analyte with the electrode of any one of claims 1 to 15 or the analyte detector of claim 21; and

[0134] (b) performing an electrochemical measurement with the electrode or detector, thereby determining at least one analyte.

[0135] 24. The method of embodiment 23, wherein the electrochemical measurement is carried out at a temperature of 30 °C to 40 °C, preferably at a temperature of at least 30 °C, at least 32 °C, at least 35 °C or at least 37 °C.

[0136] 25. The use of embodiment 22 or the method of embodiment 23 or 24, wherein the determination of at least one analyte in the fluid relates to diagnostic purposes, environmental monitoring and control, food safety, quality control or manufacturing processes.

[0137] 26. The electrode of any one of embodiments 1 to 15, the method of any one of embodiments 16 to 20, or the use of embodiment 22, or the method of embodiment 23 or 24, or the use or method of embodiment 25, wherein the fluid sample is a liquid or a dissolved environmental sample or a solution of at least one chemical compound.

[0138] 27. The electrode of any one of embodiments 1 to 15, the method of any one of embodiments 16 to 20, or the use of embodiment 22, or the method of embodiment 23 or 24, or the use or method of embodiment 25 or 26, wherein the fluid is a high ionic strength fluid.

[0139] 28. The electrode, method or use of embodiment 27, wherein the fluid is a body fluid.

[0140] 29. The electrode, method or use of embodiment 28, wherein the body fluid is selected from: blood, plasma, serum or any part thereof, saliva, tears, mucus, lymph fluid, cerebrospinal fluid, urine, feces, sweat, semen, synovial fluid.

[0141] All references cited in this specification are hereby incorporated by reference in their entirety and in the specific disclosure aspects specifically mentioned in this specification.

[0142] Drawings of the specification

[0143] Figure 1 : Surface characterization by quartz crystal microbalance. (a) Schematic diagram of a QCM chip. (b - c) Representative data sets showing typical frequency and dissipation changes vs. time for different harmonic peaks (n = 5, 7, 9) and surface modifications: dsDNA (b), PEG (c), and a mixture of both (d). Generally, the frequency decreases upon molecular adsorption while the dissipation increases. (e - f) Summary of dissipation and frequency changes obtained from at least 3 chips. For PEG, the dissipation increases by more than 3 times because it is softer than DNA (e). This is also seen as a "broadening" of the harmonic peaks for the PEG - modified surface (f).

[0144] Figure 2: FET measurement. (a) Schematic diagram of the extended gate FET configuration, where the gold surface of the QCM chip is connected to a commercially available MOSFET that is spatially separated. (b - d) Typical transfer curves recorded using different surfaces at different concentrations of PBS buffer: bare gold (b), dsDNA (c), and DNA + PEG mixture (e). (d) Potential change of three different surfaces vs. PBS concentration. Due to non - specific ion adsorption (background), ΔV becomes more positive for gold. In the case of DNA on the surface, the DNA shielding effect is superimposed, resulting in a less significant shift towards positive values. The opposite trend is observed for the mixed layer.

[0145] Figure 3 : (a) FET data from Figure 2 after subtracting the gold background. As the PBS concentration increases, the potential shifts towards more negative values because the DNA shielding by electrolyte ions is stronger. A clear signal enhancement is observed for the PEG + DNA mixture vs. bare DNA. (b) Proposed model: Adding PEG increases the effective Debye length. As a result, a larger portion of DNA can be seen at a given ionic strength, leading to a larger signal.

[0146] Figure 4 : (A) Schematic diagram of the measurement setup. A semiconducting CNT (carbon nanotube) network is arrayed between interdigitated Au electrodes (channel length = 20 µm, channel width = 2 mm). In addition, the contacts are passivated with SU - 8 photoresist to avoid leakage current. SU - 8 is a negative epoxy - type near - UV photoresist based on EPON SU - 8 epoxy resin (from Shell Chemical), which was originally developed and patented by IBM (U.S. Patent No. 4882245). To supply different liquids to the sensing surface, a microfluidic PDMS chamber with a PTFE tube is used. An Ag / AgCl reference electrode is placed in the middle of the microfluidic channel. (B) Photograph of the measurement setup. (C) Typical transfer curves in two polarization directions. The hysteresis is very small. (D) AFM image of the electric - field - arrayed CNT network. The horizontal stripes on the top and bottom are gold electrodes for aligning the CNTs. The same electrodes are subsequently used as contacts for the source and drain electrodes of the electrical measurement.

[0147] Figure 5: Comparison of GFP detection using PEGylated (A - C) and non - PEGylated CNT FETs (D - F). The surface of CNTs was modified with a mixture of pyrene - butyric acid (PBA) and pyrene - poly(ethylene glycol) (A) or with PBA only (D). Then, a camelid nanobody (VHH) specific for green fluorescent protein (GFP) was immobilized on both surfaces and exposed to GFP solution to evaluate VHH - GFP binding. B and E show the transfer curves measured in 100 mM Tris buffer at different GFP concentrations. A shift to the right can be seen in both cases, with a stronger response in the PEGylated case (B). C and F summarize the potential shift ΔV obtained as a function of GFP concentration C GFP in solutions with 1 mM and 100 mM ionic strength. ΔV was read at a constant I SD value, as indicated by the horizontal lines in B, E. The signal in C is up to 3 times larger than the signal in F, which is attributed to local buffer dilution caused by PEG.

[0148] Figure 6 : Two different surface modifications on gold were compared - using 10 kDa PEG (A) and without 10 kDa PEG (D). Both surfaces were functionalized with anti - TSH antibody fragments using short 0.5 kDa PEG linkers. Transistor transfer curves for different BSA and TSH concentrations in 10 mM buffer are plotted in (B) and (E) for the surface with PEG and without PEG, respectively. The measured voltage shifts are shown in (C) for the surface with PEG and in (F) for the surface without PEG. The shifts are plotted vs. the concentration of TSH (solid symbols) or BSA (open symbols) in 10 mM (triangles) and 150 mM buffer (squares). A significant signal enhancement was observed in the case with PEG, with a notable response even in 150 mM buffer. In all cases, the contribution of non - specific adsorption of BSA was small.

[0149] Figure 7 : TSH calibration curves measured in horse serum at 21 °C and 37 °C. Error bars represent the standard deviation of 4 different chips. Examples

[0150] The present invention will be described by way of examples. However, these examples are only illustrative of the invention and should not be construed in any way as limiting the scope of the invention.

[0151] Example 1: Polyethylene glycol reduces the role of Debye screening in transistor - based DNA detection

[0152] QCM-D measurements: Measurements were carried out using the Q-sense E4 from Biolin Scientific AB (Stockholm, Sweden) with their electrochemical module. QCM-D simultaneously monitors the shifts in oscillator frequency (Δf) and energy dissipation (ΔD) at different overtones on a 4.95 MHz, gold-coated QCM-D sensor. In addition, the electrochemical module of Q-sense allows simultaneous FET measurements. Before use, the gold-coated sensor was cleaned by UV / ozone treatment for 10 minutes and then immersed in a solution containing 1 part H 2 O 2 (30%) + 1 part NH 4 OH + 5 parts deionized (DI) water for 10 minutes. Subsequently, the sensor was thoroughly washed with DI water and dried with N 2 gas. All experiments were carried out using freshly cleaned sensors.

[0153] Extended gate FET measurements: The electrochemical module used in the QCMD experiments allows the gold-coated sensor to be electrically connected to the gate terminal of a commercial MOSFET. An Ag / AgCl reference electrode (WPI, Dri-REF™, custom length) was placed near the surface of the sensor in the outlet flow channel. Electrical measurements were carried out using a two-channel source meter (Keithley 2636B).

[0154] dsDNA preparation and immobilization: Complementary single-stranded DNAs (i): 5´-CAATGCAGATACACTTTTTT-C3H6-SH-3´ and (ii): 5´-AGTGTATCTGCATTG-3´ were purchased from STAB vida, Portugal. Strand (i) was thiolated to facilitate immobilization on the gold surface. Methoxypolyethylene glycol thiol (mPEG-SH) with an average molecular weight of 10 kDa was purchased from Nanocs, Inc. (Boston, MA). Unless otherwise stated, all other reagents were purchased from Sigma-Aldrich.

[0155] (A) Formation of dsDNA, (B) PEG, and (C) mixed dsDNA-PEG layers: The formation of all three different surfaces was monitored by both QCM-D and FET measurements. Before each experiment, thiolated double-stranded DNA (dsDNA) was formed by mixing equimolar solutions of strands (i) and (ii) (10 μL each of 100 μM stock solutions stored at -20 °C) in 100 mM PBS at room temperature for 30 minutes. And before each experiment, a thiolated mPEG solution (5 mM) was prepared by dissolving PEG powder in water. (A) dsDNA SAM: A monolayer of dsDNA was formed on the gold-coated QCM-D sensor by injecting 1 μM thiolated dsDNA at 100 μL / min for 10 minutes in buffer A (1 M NaCl, 1 mM EDTA in 10 mM Tris pH 7), followed by incubation for 1 hour. Subsequently, the surface was washed with buffer A at 100 μL / min for 10 minutes. After washing away the excess dsDNA, PBS solutions of different concentrations (200 mM, 100 mM, 50 mM, 10 mM, and 1 mM) were injected into the chamber (100 μL / min for 10 minutes), and the FET response was recorded until a stable signal was obtained. (B) PEG SAM: Similar to A, except that 1 μM thiolated mPEG (average MW = 10 kDa) was injected to form a PEG monolayer. (C) Mixed dsDNA-PEG layer: First, a dsDNA monolayer was formed by incubating the sensor surface with 1 μM thiolated dsDNA. Subsequently, the excess dsDNA was washed away by flowing buffer A at 100 μL / min for 10 minutes. PEG was added to the dsDNA monolayer by injecting 1 μM mPEG-SH in buffer A into the chamber and incubating for 20 minutes, and then washing away the unbound PEG molecules by flowing buffer A at 100 μL / min for 10 minutes. Similar to A and B, PBS solutions of different concentrations were then injected into the chamber to monitor the FET response of the mixed layer under different ionic strength conditions.

[0156] Surface characterization: The formation of monolayers of dsDNA, PEG, and mixed dsDNA + PEG layers on gold-coated quartz crystals was characterized using a dissipative quartz crystal microbalance (QCMD). In QCMD, changes in the resonance frequency and dissipation energy of an oscillating piezoelectric quartz crystal are measured in response to mass adsorption on the crystal surface ( Figure 1 a). These are then used to obtain the viscoelasticity, as well as the density, mass, and thickness of the adsorbed layer. Figure 1 b - d show the fixed different harmonic peaks (n = 3 (not shown inFigure 1 The changes Δf in resonance frequency and ΔD in dissipation factor for 5, 7, 9) in b-d). The mean values ΔD and Δf calculated from three independent trials for each layer are shown separately in Figure 1 the bar graphs of e and f. Notably, even though the detected Δf for PEG (─Δf = 35 - 45 Hz) is less than twice that of dsDNA (─Δf = approximately 25 Hz), the dissipation change is approximately 6-fold. This is consistent with the fact that dsDNA behaves like a more rigid layer on the surface, while PEG is known to form a highly hydrated brush-like structure upon immobilization. This is also evident in the spread of Δf and ΔD based on PEG immobilization compared to dsDNA where all harmonics change with the same amplitude.

[0157] Voigt viscoelastic modeling was used to fit the QCMD data for all three different layers and obtain parameters that describe the physical properties of the formed layers. The fitted parameters are listed in Table 1.

[0158] Table 1: Parameters of Voigt viscoelastic modeling derived from QCMD results

[0159]

[0160] The measured thickness of dsDNA in 1 M NaCl is approximately 5 nm, which is close to the expected length of 15 base-pair dsDNA (5 nm) plus linker (1 nm). This indicates that dsDNA molecules are vertically immobilized on the surface rather than lying flat on the surface. When 1 μM PEG was used for immobilization, the PEG (10 kDa) layer was thicker at approximately 9 nm. The viscoelastic modeling results also showed that both the shear viscosity and shear modulus of dsDNA (4.7 + / - 0.7 mPa·s and 1.3 + / - 0.6 MPa, respectively) are higher than those of PEG (1.8 + / - 0.3 mPa·s and 0.26 + / - 0.05 MPa, respectively). These viscoelastic parameters indicate that dsDNA forms a more rigid layer on gold compared to PEG. Our results are consistent with previous studies which have shown that dsDNA forms a rather rigid film, while PEG is known to form a soft brush-like structure on the surface.

[0161] In a mixed layer where two components are added sequentially, each step can be modeled independently. As shown in Table 1, since dsDNA was immobilized first, the fitted parameters are consistent with those of the layer with only dsDNA. Based on these results, the surface coverage was estimated to be 60%. Subsequently, PEG was added, which is electrically neutral and may potentially bind to the void spaces between dsDNA molecules. The changes in both dissipation and frequency upon addition of PEG demonstrated that PEG was successfully incorporated into the dsDNA layer, resulting in a mixed film.

[0162] FET measurements: The QCMD device used to characterize all three different surfaces has an electrochemical module that allows for simultaneous FET measurements on the same chip using an extended gate FET (EGFET) configuration. In this particular EGFET setup, the gold surface of the QCM chip is electrically connected to the gate terminal of a commercial MOSFET ( Figure 2 a). The sensing surface with the biological solution is separated from the MOSFET readout transistor. This reduces the complexity of the sensing chip and protects the readout transistor from contact with the solution while maintaining the charge sensitivity of a conventional ion-sensitive FET (Tarasov 2016, 2D Mater. 2: 044008; Tarasov 2016, Biosens. Bioelectron. 79: 669). To study the effect of PEG on the Debye screening of dsDNA via EGFET measurements, the transfer curves obtained from the gold-immobilized dsDNA were compared with the mixed dsDNA+PEG layer in solutions of different ionic strengths (1 - 200 mM PBS) ( Figure 2 ). As a negative control, the response of the gold-only surface was also measured while changing the ionic strength ( Figure 2 b). As the PBS concentration increases, the transfer curve of the MOSFET shifts towards more positive values ( Figure 2 b and 2d). This positive potential shift means that negative ions adsorbed on the gold surface are detected, as previously observed in the case of Cl - ions and gold (Tarasov 2012, ACS Nano 6: 9291).

[0163] In the presence of dsDNA, the transfer curve still shifts to the right as the PBS concentration increases, but to a lesser extent compared to the gold-only surface ( Figure 2 c and Figure 2 d). In the absence of anions adsorbed onto the gold, the expected effect of increasing ionic strength on dsDNA detection should shift the potential towards more negative values. However, due to this background anion adsorption, a positive potential shift is still observed as the ionic strength increases. Thus, differential signals (ΔV = ΔV dsDNA - ΔV 仅有Au ) were obtained at different PBS concentrations in order to observe the effect of ionic screening on dsDNA measurements. On the other hand, in the mixed layer, the presence of PEG makes the detection of the negative charges on dsDNA more prominent. At this time, increasing the PBS concentration shifts the potential to the left ( Figure 2 e and Figure 2 d). When subtracting the background, the shift of the signal is thus higher in the presence of PEG than in its absence.

[0164] These transistor-based measurements showed that once PEG was present on the sensor surface, there was an enhanced signal detected by dsDNA even at 200 mM PBS. This indicates that at the concentrations used in our experiments, PEG does increase the Debye length even at physiological salt concentrations. This is likely due to the salting-out effect that PEG exhibits in certain aqueous salt solutions, characterized by a partial compositional spatial separation of the PEG-rich phase from the salt-rich phase in the ternary water + PEG + salt system. It is believed that this tendency of PEG for "salting-out" ions is the reason for the reduced effective ion concentration in the region closely surrounding the PEG molecules. Thus, this "local desalination" region will exhibit a relatively higher Debye length relative to the apparent ionic strength.

[0165] Figure 3 a shows the signal enhancement achieved when PEG was mixed with dsDNA on the sensor surface. The downward and upward-pointing triangles are the ionic strength dependencies of the FET signals of dsDNA in the absence and presence of PEG, respectively. This is the same data as shown in Figure 2 d, normalized by subtracting the background signal measured with bare gold. This data confirms that the signal was enhanced by at least 3 - 4 times when PEG was present (mixed). Using the Figure 3 salt dilution model schematically shown in b, we then compared the normalized data for dsDNA detection in the absence and presence of PEG. Figure 3 The upper dashed line in a represents the dependence of the Debye length on the total ionic strength. Taking into account that PEG reduces the ion concentration near dsDNA, we calculated the expected signal after a series of ionic strength dilutions. At dilutions of at least 10-fold, the expected signal ( Figure 3 lower dashed line in a) now overlaps with the signal obtained in the presence of PEG ( Figure 3 upward triangle in a). This means that using our experimental conditions, we can achieve at least a 10-fold dilution of the ion concentration in the mixed layer with PEG, and thus enhance the dsDNA signal by at least 3 times.

[0166] It has been demonstrated that the effects of ionic shielding on transistor-based dsDNA detection can be mitigated by adding PEG to the sensor surface. Measurements using EGFETs on gold electrodes showed that PEG enhanced the signal from dsDNA by at least 3-fold, even in high ionic strength solutions (up to 200 mM PBS). This reduction in ionic shielding can be attributed to the tendency of PEG to exclude ionic species in its immediate surrounding area (salting-out effect). This means that in the presence of sufficient PEG molecules on the sensor surface, localized desalted regions within the PEG layer are created. Thus, even in the presence of high ionic strength solutions, the region within the PEG layer on the sensor surface has an effectively lower ionic strength environment than the region above the PEG layer. This increases the Debye length in the region immediately adjacent to the sensor surface, thus increasing the detection range to much greater distances from the surface.

[0167] The application of PEG can also be extended to other types of biosensors. The key part is that the amount of PEG and receptors on the surface should be adjusted to maximize the receptor density and desalting effect at the same time. In addition to alleviating ion shielding, PEG has also been shown to be effective in preventing nonspecific interactions and making biomolecules on the sensor surface more stable. With all these advantages, incorporating PEG on the surface of transistor-based biosensors may have a huge impact in pushing FET biosensors towards PoC applications.

[0168] Example 2: Bioassay in high ionic strength solution using PEGylated carbon nanotube transistors with nanobody receptors

[0169] FET-based biosensors are made of high-quality, ordered semiconducting carbon nanotube networks, which provide sensitive and stable transducers as well as scalable manufacturing processes. A combined surface functionalization scheme is proposed to overcome Debye screening: 1) short nanobodies (V H H) act as a receptor to enable analytes to bind closer to the surface; and 2) add a polyethylene glycol (PEG) layer to increase the effective Debye length. Using green fluorescent protein (GFP) as a model system, a three-fold signal enhancement was demonstrated with the PEGylated surface in high ionic strength solutions. The mechanism was explained by local buffer dilution of PEG mediated by the salting-out effect. The sensor achieved a sub-pM detection limit with a dynamic range of more than 4 orders of magnitude. Moreover, if additional surface passivation was implemented, the sensor was highly specific with negligible contribution from nonspecific adsorption.

[0170] Immobilization of Nanobodies (Formation of Mixed Self-Assembled Monolayers (SAMs)): After fabricating the CNT FET as described previously (Rother 2016, ACS Appl. Mater. Interf. 8: 5571), the CNTs were washed with ethanol for 1 h to remove residual polymers / contaminants from the CNT sorting process. Subsequently, the CNT electrodes were treated with 1 mM pyrenebutyric acid (PBA, linker) + 0.25 mM pyrene-PEG (10 kDa) in ethanol for 1 h. Subsequently, it was rinsed briefly with ethanol and H 2 O and then kept overnight in 100 mM Tris pH 7.4. Subsequently, after a brief rinse with distilled water, an aqueous solution of 100 mM N-hydroxysuccinimide (NHS) and 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) was pumped into the channel for 20 min to activate the surface. Thereafter, 10 µM GFP-specific VHH nanobody or 10 µM bovine serum albumin (BSA) in 0.2 M phosphate buffer (pH 7) was introduced for 1 h for specific and non-specific adsorption experiments, respectively. Subsequently, the surface was rinsed with 100 mM Tris for 15 min to inactivate possible remaining active sites. Finally, in the case of the optimized CNT FET-based GFP sensing assay, the surface was exposed to a 100 nM BSA solution in 100 mM Tris for 30 min to reduce non-specific binding.

[0171] GFP Binding Measurement: The electrodes were rinsed with the indicated concentration of green fluorescent protein (GFP) in 1 mM Tris pH 7.4 for 15 min and three consecutive measurements were performed. Then, the electrodes were rinsed with the same concentration of GFP in 100 mM Tris pH 7.4 for 15 min and three consecutive measurements were performed again. These measurements and treatments were repeated over the entire concentration range from 0 to 100 nM GFP.

[0172] As Figure 4 shown in Figure 4 A, the CNT FET was operated in a liquid-gated configuration. A photograph of the measurement setup is provided in Figure 4 B. The CNTs were sorted by means of a polymer and aligned in an electric field according to a previously published protocol (Rother 2016, ACS Appl. Mater. Interf. 8: 5571). Typical transistor transfer curves are presented in Figure 4 C. All devices exhibited bipolar behavior with small hysteresis, a current on / off ratio exceeding 10 4 and a steep subthreshold swing below 110 mV / dec. AFM images of a typical device are shown in Figure 4 D.

[0173] Using pyrene butyric acid (PBA) as a linker molecule, the GFP-specific VHH was conjugated to the amino groups of the carbon nanotubes. To study the effect of PEG on the signal, the VHH was immobilized on surfaces coated with PBA+PEG ( Figure 5 A) and on control CNT samples modified only with PBA ( Figure 5 D). The two sensor surfaces were then exposed to various GFP concentrations dissolved in 1 mM or 100 mM Tris buffer. For the PEGylated surface, the measurement results are shown in Figure 5 B, C, and for the non-PEGylated surface, the measurement results are shown in Figure 5 E, F. In both cases, in response to increasing GFP concentration, the transfer curves shifted towards more positive values, and the PEGylated surface reacted more strongly ( Figure 5 B, E). Figures C and F compare the responses of the two sensors as a function of GFP concentration. Importantly, compared to the non-PEGylated surface, the signal of the PEGylated sensor showed a three-fold enhancement in 100 mM buffer (25 mV vs. 8 mV for 100 nM GFP). The signal enhancement observed in 1 mM buffer was less dramatic and corresponded to an increase of approximately two-fold (47 mV vs. 25 mV for 100 nM GFP). These results clearly demonstrate a strong positive effect of PEG on the maximum achievable sensor response.

[0174] Example 3: Transistor-based Detection of Thyroid Stimulating Hormone in Serum

[0175] Here, a surface chemistry approach combines antibody fragments as short biological receptors with desalted polyethylene glycol (PEG) molecules to enable direct label-free selective immunosensing in serum. Here, thyroid stimulating hormone (TSH) was selected as a representative analyte - a relevant and well-characterized immunosensing parameter with demanding sensitivity requirements. This approach was demonstrated with an extended gate configuration consisting of a gold sensing surface electrically connected to a commercial MOSFET transducer. This setup has the advantages of simple chip fabrication and well-established thiol-gold chemistry for attaching linker molecules. To evaluate the effect of desalted PEG on the Debye length, two different surface modifications were compared - with and without the addition of 10 kDa PEG. The sensing surface without the SH-PEG-COOH (0.5 kDa) self-assembled monolayer (SAM) is hereinafter referred to as the "Mono SAM" configuration. To study the system in the presence of PEG, SH-PEG-COOH (0.5 kDa) and SH-PEG (10 kDa) were combined and referred to as the "mixed SAM" configuration ( Figure 6 ).

[0176] Anti-TSH antibody fragments (F(ab')2) were immobilized on the sensor surface, and TSH detection in buffer and serum was demonstrated. PEG was co-immobilized on the surface to increase the Debye length. The PEG-to-linker ratio was chosen as 1:20 to incorporate the fragments. With this ratio, a three-fold signal enhancement was achieved in high ionic strength buffer compared to a control surface without PEG ( Figure 6 ). Additionally, non-specific adsorption was tested by exposing the chip to BSA and was shown to be very low (<10% of the specific signal, Figure 6 ). Subsequently, measurements were repeated in serum with at least 3 devices. The calibration curve is presented in Figure 7 . If the measurements were performed at 37 °C, a strong signal enhancement was achieved, resulting in a three orders of magnitude lower detection limit compared to 21 °C. This improvement is mainly attributed to the faster kinetics of binding at higher temperatures.

[0177] Cited references

[0178] Gao 2015, Nano Lett. 15: 2143;

[0179] Gao 2016, PNAS 113(51): 14633-14638;

[0180] WO2016 / 161246;

[0181] Rother 2016, ACS Appl. Mater. Interf. 8: 5571;

[0182] Shim 2002, Nano Lett. 2: 285;

[0183] Tarasov 2012, ACS Nano 6: 9291;

[0184] Tarasov 2016, 2D Mater. 2: 044008;

[0185] Tarasov 2016, Biosens. Bioelectron. 79: 669;

[0186] Yoshimoto 2010, J. Am. Chem. Soc. 132: 7982. Sequence Listing <110> ROCHE DIAGNOSTICS GMBH F. HOFFMANN-LA ROCHE AG ROCHE DIAGNOSTICS OPERATIONS, INC. <120> Polymer coating for an electrode of a sensor device <130> RD33974PC <140> PCT / EP2018 / 054280 <141> 2018-02-21 <150> EP17157373.6 <151> 2017-02-22 <160> 2 <170> BiSSAP 1.3.6 <210> 1 <211> 20 <212> DNA <213> Artificial sequence <220> <223> DNA (i) <400> 1 caatgcagat acactttttt 20 <210> 2 <211> 15 <212> DNA <213> Artificial sequence <220> <223> DNA (ii) <400> 2 agtgtatctg cattg 15

Claims

1. A method for fabricating a functionalized surface on an electrode, comprising the steps of: a) applying the linker and the polymer to the electrode under conditions that permit covalent or non-covalent immobilization of the linker and at least one polymer capable of mediating a salting-out effect on the surface of the electrode; and b) applying at least one detection agent to the electrode on which the linker and the polymer are immobilized under conditions that permit covalent or non-covalent attachment of at least one detection molecule to the electrode via the immobilized linker; and wherein the conditions permit distribution of the at least one polymer capable of mediating a salting-out effect and the at least one detection agent on the surface of the electrode such that the detection agent is present in an equal amount per surface area over the entire electrode surface and the polymer capable of mediating a salting-out effect is disposed around the detection agent in an amount such that: i) it can reduce the ionic strength of the fluid near the detection agent, and ii) it can bind an analyte contained in a fluid sample; wherein the linker is a low molecular weight (MW) polymer having an MW of 0.01 kDa to 5 kDa; and wherein the size of the polymer capable of mediating a salting-out effect is greater than or equal to the size of the detection agent, wherein the polymer capable of mediating a salting-out effect is poly(ethylene glycol) (PEG), wherein the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is 10 kDa to 100 kDa, and wherein the polymer capable of mediating a salting-out effect has the same class as the polymer serving as the linker polymer.

2. The method of claim 1, wherein the linker is a low molecular weight (MW) polymer having an MW of 0.01 kDa to 1.0 kDa, 0.01 kDa to 0.5 kDa, or 0.1 kDa to 0.5 kDa.

3. An electrode comprising a functionalized surface exposable to a fluid sample, the functionalized surface comprising at least one polymer capable of mediating a salting-out effect and at least one detection agent that binds an analyte contained in the fluid sample, wherein the at least one polymer capable of mediating a salting-out effect and the at least one detection agent are distributed on the surface of the electrode such that the detection agent is present in a substantially equal amount per surface area over the entire electrode surface and the polymer capable of mediating a salting-out effect is disposed around the detection agent in an amount such that: i) it can reduce the ionic strength of the fluid near the detection agent, and ii) it can bind an analyte contained in a fluid sample; wherein the at least one detection agent is immobilized on the surface of the electrode via a linker; wherein the linker is a low molecular weight (MW) polymer having an MW of 0.01 kDa to 5 kDa; and wherein the size of the polymer capable of mediating a salting-out effect is greater than or equal to the size of the detection agent, wherein the polymer capable of mediating a salting-out effect is poly(ethylene glycol) (PEG), wherein the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is 10 kDa to 100 kDa, and wherein the polymer capable of mediating a salting-out effect has the same class as the polymer serving as the linker polymer.

4. The electrode of claim 3, wherein the size of the at least one polymer capable of mediating a salting-out effect is such that the Debye length is increased and the ionic strength around the at least one detection molecule is decreased.

5. The electrode of claim 4, wherein the molecular weight (MW) of the at least one polymer capable of mediating a salting-out effect is from 10 kDa to 50 kDa, from 10 kDa to 25 kDa, from 10 kDa to 20 kDa or from 10 kDa to 15 kDa.

6. The electrode of any one of claims 3 to 5, wherein the at least one polymer capable of mediating a salting-out effect is immobilized on the surface of the electrode via a functional group without an additional linker.

7. The electrode of any one of claims 3 to 5, wherein the at least one detection agent specifically binds to an analyte contained in a fluid sample, and wherein the at least one detection agent is selected from: antibodies and fragments thereof, nucleic acids, aptamers, peptide nucleic acids (PNA), receptor or ligand proteins or peptides, and enzymes.

8. The electrode of any one of claims 3 to 5, wherein the size of the at least one detection agent is such that the detection agent is surrounded by the at least one polymer capable of mediating a salting-out effect.

9. An analyte detector for determining at least one analyte, comprising the electrode of any one of claims 3 to 8 or an electrode obtained by the method of any one of claims 1 to 2, wherein the electrode is in electrical contact with a transducer or is part of a transducer.

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