Patent Description:
The devices and methods of the present invention, as an example, may be used for diagnostic purposes, e.g. in clinical or laboratory analytics or for home monitoring purposes. The devices and methods of the present invention specifically may be used for detecting one or more analytes in body fluids or other liquids. As an example, DNA detection may be named. Other applications and uses, however, are feasible.

A wide variety of analyte detectors for detecting at least one analyte in at least one fluid sample have been described. Analyte detectors configured for reliably detecting chemical and/or biological species in a qualitative and/or quantitative manner can be used for various purposes such as, but not limited to, diagnostic purposes, monitoring of environmental contamination, food safety evaluation, quality control or manufacturing processes. Such analyte detectors may for instance rely on transistor-based measurements for identification the at least one analyte. Transistor-based analyte detectors have been adapted to allow the detection of a wide range of analytes including biomolecules such as proteins, antibodies, antigens, DNA, and chemical species such as ionic species and electrolytes.

A number of studies describe the use of transistor-based analyte detectors in the identification of antigens, antibodies or other proteins: Elnathan et al. (<NPL>) describe the detection of proteins in untreated serum and blood samples in the sub-pM concentration range using a nanowire-based field-effect transistor (FET) device combined with size-reduced antibody fragments. The use of size-reduced antibody fragments permits the biorecognition event to occur in closer proximity to the nanowire surface, falling within the charge-sensitive Debye screening length. In a study published by Gao et al. (<NPL>) incorporation of a porous and biomolecule permeable layer on a FET-based nanoelectric sensor is described.

The polymer layer increases the effective screening length in the region immediately adjacent to the FET-based sensor surface and thereby enables the detection of biomolecules in high ionic strength solutions in real-time. The same study also reports that silicon nanowire field-effect transistors with additional polyethylene glycol (PEG) modification can readily detect prostate specific antigen (PSA) in solutions with phosphate buffer concentrations as high as <NUM>. (<NPL>) present a simple and sensitive method for real-time detection of a prostate cancer marker (PSA-ACT complex) through label-free protein biosensors based on a carbon nanotube field-effect transistor (CNT-FET). Tarasov et al. (<NPL>) use gold-coated graphene FETs to measure the binding affinity of a specific protein-antibody interaction. In a different study, Tarasov et al. (<NPL>) employ an extended-gate field-effect transistor for direct potentiometric serological diagnosis using the model pathogen Bovine Herpes Virus-<NUM> (BHV-<NUM>). To demonstrate the sensor capabilities as a diagnostic tool, BHV-<NUM> viral protein gE is expressed and immobilized on the sensor surface to serve as a capture antigen for BHV-<NUM>-specific antibody (ant-gE). The gE-coated immunosensor was shown to be highly sensitive and selective to anti-gE and significantly faster than Enzyme-Linked ImmunoSorbent Assay (ELISA) that is typically performed by centralized laboratories.

Other studies explore the potential of transistor-based analyte detectors for the identification of nucleic acids, such as DNA or RNA, or possible components thereof, such as aden-osinmonophosphate (AMP). In <CIT> a biosensor is disclosed that can convert biological interactions into electrical and optical signals to sense a material to be analyzed. The biosensor includes a substrate, a source electrode and a drain electrode formed on one surface of the substrate, a carbon nanotube connecting the source and the drain electrodes, a metal gate covering the carbon nanotube, a recognition component immobilized on the metal gate, and a passivation layer covering the source and drain electrodes. In one embodiment, the recognition component may be a single-stranded oligonucleotide such as DNA or RNA. In the case of DNA, the biosensor has a recognition DNA immobilized on the surface of the metal gate. Electrical and/or optical signals are generated as a result of hybridization between the recognition DNA and a target DNA. Zayats et al. (<NPL>) present research applying aptamers for the label-free reagent-less analysis of small molecules. They demonstrate that the small substrate-induced separation of a duplex nucleic acid that includes the aptamer strand, on an ion-sensitive field-effect transistor (ISFET) or on an electrode, forms a substrate-aptamer complex that can be electrically characterized. In particular, an amine-functionalized nucleic acid that acts as aptamer was immobilized on the gate surface and further hybridized with a short nucleic acid. The addition of adenosine displaces the short nucleic acid and assembles the aptamer into the hairpin configuration that binds adenosine mono-phosphate (AMP).

Understanding and controlling the behavior of the analyte detector is crucial for its targeted use. Transistor-based analyte detectors can also respond to chemicals species such as ionic species and electrolytes. Tarasov et al. (<NPL>) use silicon nanowires coated with highly pH-sensitive hafnium oxide (HfO<NUM>) and aluminum oxide (Al<NUM>O<NUM>) in silicon nanowire field-effect transistor to determine their response to changes in the supporting electrolyte concentration. Wipf et al. (<NPL>) modify individual nanowires with thin gold films as a novel approach to surface functionalization for the specific detection of electrolyte ions by ion-sensitive field-effect transistors in a differential setup. They find that a functional self-assembled monolayer does not affect the unspecific response of gold to pH and background ionic species, which represents a clear advantage of gold compared to oxide surfaces.

Thus, transistor-based analyte detectors have been adapted in numerous ways to detect a multitude of analytes. The advances established in the field of transistor-based analyte detectors are in part due to advances in surface functionalization techniques, in particular those applicable to nano-devices. Shim et al. (<NPL>) study the adsorption behavior of proteins on the side of single-walled carbon nanotubes. They report that the functionalization of single-walled carbon nanotubes by co-adsorption of a surfactant and polyethylene glycol is found to be effective in resisting non-specific adsorption of streptavidin. In <CIT> a method is disclosed for immobilizing nucleic acid and a method for manufacturing a biosensor using the same method. The method provided enables high-density absorption when immobilizing nucleic acid probes onto a solid support by suppressing electrostatic repulsion among the nucleic acids. A nucleic acid immobilization method to immobilize a nucleic acid onto a solid support, includes: preparing a solution containing a probe molecule which includes a nucleic acid, a spacer molecule, and at least one kind of a divalent cation; and contacting the solution with the solid support for incubation. Yoshimoto et al. (<NPL>) examine the adsorption behavior of antibody fragments directly immobilized on a gold surface through S-Au linkage. They report that the conformational and/or orientation change of antibody fragments was suppressed by a coimmobilized mixed polyethylene glycol layer. Yoshimoto et al. expect their findings to be useful for the improvement of the antibody fragment method and, thus, for the construction of high-performance immunosensor surfaces.

However, analyte detectors able to detect at least one analyte in at least one fluid sample may also be based on electrochemical measurements. For details of electrochemical test elements and potential test chemicals useful in such test elements, which may also be used within the present invention, reference may be made to <NPL>. Further, impedance biosensors are a class of electrical biosensors able to detect unlabeled DNA and protein targets by monitoring changes in surface impedance when a target molecule binds to an immobilized probe. The challenges caused by the affinity capture step and other challenges unique to impedance readout are discussed in <NPL>. Furthermore, in their fundamental study from <NUM>, Severinghaus and Bradley (<NPL>) describe an apparatus to permit rapid and accurate analysis of oxygen and carbon dioxide tensions in gas, blood or any liquid mixture using an oxygen electrode and a carbon dioxide electrode. (<NPL>) report on a miniature Clark-type oxygen sensor that has been integrated with a microstructure using a novel fabrication technique. Moreover, analyte detectors may also combine functional elements as reported by Zhu et al. (<NPL>), who present a graphene enabled, integrated optoelectromechanical device and demonstrate its utility for biomolecular sensing. They demonstrate a novel nanoscale sensing device with optical, electronic and mechanical functional elements integrated on the same chip. By having each element target a different concentration regime, the sensitivity-dynamic range trade-off of traditional single-mode sensors can be significantly mitigated.

<CIT> discloses a system for detecting a target and a method for detecting a target. The system includes a field effect transistor, having a gate, a source, and a drain; a potentiostat, having a working electrode, a counter electrode, and a reference electrode; wherein the working electrode is coupled with a detection region, and the counter electrode is coupled with the gate; wherein the detection region, the gate, and the reference electrode are arranged in an ion fluid; wherein the potentiostat is configured to generate redox in the ion fluid by an electrochemical method to detect the target.

<NPL>, describes a label-free sensor for fast bacterial detection based on metal-oxide-semiconductor field-effect transistors (MOSFETs). The electrical charge of bacteria binding to the glycosylated gates of a MOSFET enables quantification in a straightforward manner and at a higher sensitivity than is achieved with electrochemical impedance spectroscopy (EIS) and matrix-assisted laser desorption ionization time-of-flight mass spectroscopy (MALDI-ToF) on the same modified surfaces.

<NPL>, describes the use of a separative extended gate field-effect transistor (SEGFET) as an immunosensor for the label-free recognition of dengue virus nonstruc-tural protein <NUM> (NS1). NS1 is detected in a concentration range of <NUM> to <NUM>µg mL-<NUM>, indicating that the system is promising for the early and simple diagnosis of dengue.

<CIT> discloses a floating gate based sensor apparatus including at least two separate electrical bias components with respect to a floating gate based sensor surface within the floating gate based sensor apparatus. By including the at least two electrical bias components, the floating gate based sensor apparatus provides enhanced capabilities for biomaterial and non-biomaterial detection and manipulation while using the floating gate based sensor apparatus.

<NPL>, describes the application of silicon nanowire field-effect transistors (SiNWFET) devices for noninvasive, real-time monitoring of interfacial effects during cell growth and differentiation using cultured rat adrenal pheochromocytoma (PC12) cells. Monitoring of cell adhesion during growth and morphological changes during neuronal differentiation was performed by measuring the non-Faradaic electrical impedance of the cell-SiNW FET system using a precision LCR meter.

<NPL>, describes the fabrication and characterization of graphene based field-effect transistors (GFETs) and introduces the new developments in physical, chemical, and biological electronic detection using GFETs. Further, several perspectives and current challenges of GFETs development are presented, and some proposals are suggested for further development and exploration.

<CIT> discloses a bio material receiving device including a thin film transistor (TFT) including a drain electrode, and a nano well accommodating a bio material. The drain electrode includes the nano well. The TFT may be a bottom gate TFT or a top gate TFT. A nano well array may include a plurality of bio material receiving devices. In a method of operating the bio material receiving device, each of the bio material receiving devices may be individually selected in the nano well array. When the bio material is accommodated in the selected bio material receiving device, a voltage is applied so that another bio material is not accommodated.

<NPL>], describes the fabrication and characterization of a combined pO<NUM>, pCO<NUM> and pH chemical sensor, designed for blood gas monitoring. Classical electrochemical principles are used in a miniaturized planar-type structure. Both amperometric (pO<NUM>) and potentiometric devices (pCO<NUM>, pH) are integrated on a <NUM> x <NUM> chip. The transducer part of the chip is realized using standard silicon technology. Polyacrylamide and polysiloxane layers, which are used as hydrogel and gas-permeable membrane, respectively, are deposited and patterned by photopolymerization. Thus, the whole sensor is fabricated on wafer level using IC-compatible processes. The characterization has been performed in aqueous solutions and in blood used for transfusion. For this purpose, the chip is mounted into a flow-through cell.

<NPL>, describes how tailoring the sensing surface of a transistor-based biosensor with short specific biological receptors and a polymer polyethylene glycol (PEG) can strongly enhance the sensor response. In addition, the sensor performance can be dramatically improved if the measurements are performed at elevated temperatures (<NUM> instead of <NUM>). With this novel approach, highly sensitive and selective detection of a representative immunosensing parameter-human thyroid-stimulating hormone-is shown over a wide measuring range with subpicomolar detection limits in whole serum. This allows direct immunodetection in whole serum using transistor-based biosensors, without the need for sample pretreatment, labeling, or washing steps. The presented sensor is low-cost, can be easily integrated into portable diagnostics devices, and offers a competitive performance compared to state-of-the-art central laboratory analyzers.

<NPL>, describes combining highly stable FETs based on single-walled semiconducting carbon nanotube (SWCNTs) networks with a novel surface functionalization comprising: <NUM>) short nanobody (VHH) receptors, and <NUM>) a polyethylene glycol (PEG) layer. These measures overcome the two major challenges that have limited the use of nanomaterial-based field-effect transistors (FETs) in physiological samples: screening of the analyte charge by electrolyte ions (Debye screening) and non-specific adsorption. Nanobodies are stable, easy-to-produce biological receptors that are very small (~<NUM>-<NUM>), thus enabling analyte binding closer to the sensor surface. Despite their unique properties, nanobodies have not been used yet as receptors in FET based biosensors. The addition of PEG strongly enhances the signal in high ionic strength environment. Using green fluorescent protein (GFP) as a model antigen, high selectivity and sub-picomolar detection limit with a dynamic range exceeding <NUM> orders of magnitude is demonstrated in physiological solutions. In addition, long-term stability measurements reveal a low drift of SWCNTs of <NUM> mV/h. The presented immunoassay is fast, label-free, does not require any sample pretreatment or washing steps.

In the field of analytics, generally, one major technical challenge typically resides in the selection of appropriate methods and devices for the specific analyte to be detected. Even more, in some cases, several types of analytes in one and the same sample may have to be detected. As discussed above, a wide variety of detectors having differing sensitivities and measurement principles is available. Transistor-based detectors, as an example, are highly sensitive to the analyte charge. Analyte detectors based on electrochemical measurements are usually sensitive to the current, impedance or potential changes resulting from electrochemical reactions involving the analyte. Thus, typically, for each analyte to be detected, a specific detector has to be selected, having properties suited for the analyte. The measurement setup, consequently, typically is highly specific for the analyte to be tested for, and the setup, in total, typically lacks versatility. Further, each measurement principle typically has its own drawbacks, technical limitations and inaccuracies. Consequently, the choice of a measurement principle also implies the choice of the technical drawbacks involved with this measurement principle. A combination of measurement principles, however, typically leads to a complex setup and evaluation. There is, consequently, a general need for electronic sensors in physiological liquids which generally provide a high versatility and selectivity and which provide a more universal sensor layout as compared to the methods, measurement principles and devices known in the art.

It is therefore an objective of the present invention to provide an analyte detector and a method for detecting at least one analyte in at least one fluid sample which at least partially avoid the drawbacks and disadvantages of known measurement principles. Specifically, it is desirable to provide an analyte detector and a method which allow for a high versatility, selectivity and sensitivity and, still, which provide a more universal sensor layout as compared to known means and methods.

This problem is solved by an analyte detector for detecting at least one analyte in at least one fluid sample and a method for detecting at least one analyte in at least one fluid sample with the features of the independent claims. Preferred embodiments, which might be realized in an isolated fashion or in any arbitrary combination, are listed in the dependent claims.

In a first aspect of the present invention, an analyte detector for detecting at least one analyte in at least one fluid sample is disclosed. The analyte detector is defined in claim <NUM>.

As used herein, the term "analyte detector" may generally refer to an arbitrary device configured for an analytical examination of the sample. The analyte detector may be configured for conducting at least one analysis, such as a medical analysis, of the sample. As generally used within the present invention, the terms "analysis", "analytical examination" and "determination of one or more analytes" are used synonymously and are understood to describe a qualitative and/or a quantitative detection of the at least one analyte. In particular, said terms may be understood as a determination of the concentration or amount of the respective analyte, where the sole determination of the absence or presence of the analyte may also be regarded as an analytical examination. Thus, specifically, the analyte detector may be configured for qualitatively and/or quantitatively detecting one or more analytes, specifically in one or more samples. The detection of the at least one analyte may take place at a high degree of sensitivity.

As further used herein, the term "analyte" generally may refer to an arbitrary chemical or biological substance or species, such as an ion, an atom, a molecule or a chemical compound. The analyte specifically may be an analyte which may be present in a bodily fluid or a body tissue. The term analyte specifically may encompass atoms, ions, molecules and macromolecules, in particular biological macromolecules such as nucleic acids, peptides and proteins, lipids, sugars, such as glucose, and metabolites. Further examples of potential analytes to be detected will be given in further detail below.

As used herein, the term "fluid sample" generally may refer to a liquid or gas. The fluid sample may have a defined or definable volume. Further, the fluid sample may be comprised in a defined or definable space or may also be present in an open space such as in an open surrounding. The fluid sample may be present in a static state or may flow continuously or discontinuously. The fluid sample may, as an example, be a pure liquid or a homogeneous or heterogeneous mixture, such as a dispersion, an emulsion or a suspension. Similarly, for gases, mixtures of gases or even mixtures of gases with liquids or solids may be used.

In particular, the fluid sample can contain atoms, ions, molecules and macromolecules, in particular biological macromolecules such as nucleic acids, peptides and proteins, lipids and metabolites, and also biological cells and cell fragments. Typical fluid samples to be examined are bodily fluids such as blood, plasma, serum, urine, cerebrospinal fluid, lachrymal fluid, cell suspensions, cell supernatants, cell extracts, tissue lysates or such likes. Fluid samples can, however, also be calibration solutions, reference solutions, reagent solutions or solutions containing standardized analyte concentrations, so-called standards.

As used herein, the term "electrode" may generally refer to a functional element configured to perform a current measurement and/or a voltage measurement and/or configured to apply a current and/or an electrical potential and/or a voltage to an element in electrical contact with the electrode. In particular, the electrode may comprise a conducting and/or a semiconducting material. As an example, the electrode may comprise at least one metallic material and/or at least one organic or inorganic semiconducting material, having at least one conducting or semiconducting surface. The surface itself may form the electrode or a part of the electrode. As an example, the electrode may comprise at least one material, specifically at least one surface material, having an electrical conductivity of at least <NUM>/m, e.g. at least <NUM>/m, either isotropically or anisotropically in at least one direction.

As used herein the term "in electrical contact" may generally refer to the arrangement or configuration of at least two components, wherein at least one of the components is able to electrically influence the at least one other component and/or to at least partially control an electrical quality of the other component such as, but not limited to, its conductivity and electrical current flow, for instance via field effects. In particular, an electrode may be in electrical contact with an element without being in direct physical contact with said element. Thus, an electrode may control the electrical current flow within an element by application of a voltage despite being insulated from said element. Insulation may, for instance, be constituted by an oxide layer as is typically the case for a gate electrode of a metal oxide semiconductor field-effect transistor (MOSFET) a subgroup of insulated-gate field-effect transistors (IGFET), which is described in more detail below. Thus, generally, for being in electrical contact with each other, the at least two components may be located in close proximity, without being in direct physical contact with one another, such that, however, the components may influence one another electrically. Additionally or alternatively, however, the at least two components may also be physically connected via at least one connecting element having at least semiconducting properties or electrically conductive properties, such as by at least one electrical conductor. Again, the at least two components are separate components or are fully or partially integrated into one another. As an example, the at least one multipurpose electrode may either be connected to the field-effect transistor via at least one connecting element, such as via at least one electrically conductive lead, or may even fully or partially be integrated into the field-effect transistor. Various possibilities are given.

As used herein, the term "multipurpose electrode" may generally refer to an arbitrary electrode configured to be able to form part of at least two different measurement devices. Thus, the multipurpose electrode may take part in analytical examinations based on at least two different methods, wherein each of the methods requires the use of at least one measurement device. The multipurpose electrode may, for instance, be configured to form part of at least both the field-effect transistor and the electrochemical measurement device. Thus, the multipurpose electrode may take part in analytical examinations based on at least one of the methods comprising the use of the field-effect transistor and the at least one other method comprising the use of the electrochemical measurement device.

As further used herein, the term "exposable" generally refers to the property of an element of providing at least one surface which may be brought into contact with the at least one substance to which the element is to be exposed. Thus, as an example, the at least one multipurpose electrode may provide at least one electrode surface accessible to the fluid sample. Specifically, as will be explained by exemplary embodiments below, the analyte detector may comprise at least one fluid channel, such as a fluid channel having an inlet port and an outlet port, through which the fluid sample may flow, wherein the at least one multipurpose electrode comprises at least one electrode surface accessible from the fluid channel, such that liquid flowing through or present in the fluid channel contacts the at least one electrode surface. Other options, however, are feasible.

As further used herein, the term "field-effect transistor" may generally refer to a functional element comprising at least one source electrode, at least one drain electrode and at least one gate electrode. The field-effect transistor further comprises at least one channel. As used herein, the term "channel" of the field-effect transistor may generally refer to a component able to conduct a current between the source electrode and the drain electrode. The channel may have at least one semiconducting material and/or at least one doped semiconducting material. The semiconducting material may be or may comprise at least one of an inorganic semiconducting material and an organic semiconducting material. Typically, a semiconducting material exhibits an electrical conductivity σ of <NUM>-<NUM> S/cm < σ < <NUM><NUM> S/cm.

In the field of organic semiconductors, however, due to the impact of the low charge carrier mobilities, due to the molecular orbitals and/or due to the low charge carrier densities, however, this description is often not fully applicable. Thus, organic conductive materials are often denoted as organic semiconductors, even though their conductivity may be higher than <NUM><NUM> S/cm, such as graphene.

In particular, the semiconducting material may comprise one, two or more regions, preferably two to ten regions, more preferably three regions, wherein each region may be n-type doped or p-type doped. Specifically the semiconducting material may comprise an inorganic and/or organic semiconducting material. The channel may be able to conduct a current between the source electrode and the drain electrode only under specific external conditions. The conditions may include a temperature of the channel and/or the voltage or electrical potential applied to the channel either directly or via the gate electrode or via an external electrode. In particular, the channel may be constituted by at least one semiconducting material, such as by at least one semiconducting layer. As an example, inorganic and/or organic semiconducting materials may be used. In the following, as a specific example, graphene is used as a semiconducting material, such as by using one or more graphene layers. The gate electrode may be in direct physical contact with the channel. In this configuration the field-effect transistor may generally be referred to as "non-insulated-gate field-effect transistor" (NIGFET). In particular, the gate electrode may be at least partially identical with the channel. Alternatively, the gate electrode may be in indirect physical contact with the channel, e.g. by using one or more electrically insulating materials interposed in between the gate electrode and the channel. In this configuration the transistor may generally be referred to as "insulated-gate field-effect transistor" (IGFET).

The insulated-gate field-effect transistor may be implemented as a "metal-insulator-semiconductor field-effect transistor" (MISFET). In this case, the gate electrode which may comprise at least one metal may be insulated from the channel which may comprise at least one semiconducting material. Specifically, the insulation of the gate electrode from the channel may be constituted by an oxide. In this configuration the field-effect transistor may generally be referred to as "metal-oxide-semiconductor field-effect transistor" (MOSFET). However, other materials for insulation of the gate electrode are feasible. The channel of the field-effect transistor may be in physical contact with an electrolyte solution, which may constitute or form part of the gate electrode. In this configuration an ionic double layer may form, that may serve as insulation of the gate electrode from the channel. In this configuration the field-effect transistor may be referred to as a "solution-gated or liquid-gated FET". The electrolyte solution may comprise substances that may influence the potential applied to the channel upon close proximity or adsorption to the channel and/or the insulation of the channel, thus allowing the detection of chemical species. In this configuration the field-effect transistor may be referred to as a "chemical field-effect transistor" or ChemFET. In particular a ChemFET may be configured for the detection of ionic species forming an "ion-sensitive field-effect transistor" (ISFET) that may be sensitive to H+ and/or other ionic species. A layer sensitive to ionic species, such as Al<NUM>O<NUM>, Si<NUM>N<NUM> or Ta<NUM>O<NUM>, may be in contact with the channel or may form part of the gate electrode of the ISFET and/or may form part of the channel and the gate electrode. In another configuration, the ChemFET may comprise a layer of immobilized enzymes as part of the gate electrode and/or the channel of the field-effect transistor. In this configuration the field-effect transistor may be referred to as an "enzyme field-effect transistor" (ENFET). Binding of the enzyme to the analyte may affect the potential applied to the channel and allow detection of the analyte. Thus, the ENFET is an example of a field-effect transistor-based biosensor (BioFET). As a BioFET the field-effect transistor may comprise a layer of immobilized biomolecules as biorecognition elements able to bind one or more species of molecules, specifically biomolecules, where the binding reaction may either directly or indirectly affect the potential applied to the channel.

The field-effect transistor may further be implemented as an "extended-gate field-effect transistor". As used herein, the term "extended-gate field-effect transistor" may generally refer to a field-effect transistor comprising a gate electrode configured to allow a spatial separation of the channel of the field-effect transistor from a process or reaction that sets or affects the potential of the gate electrode. Such an electrode may generally be referred to as an "extended gate electrode". Thus, the extended gate electrode of an extended-gate field-effect transistor may allow to physically separate the process of applying a potential to the channel and the process of applying a potential to the gate electrode.

The at least one field-effect transistor may comprise at least one substrate. The substrate may have purely mechanical properties and function, such as for carrying the above-mentioned components of the field-effect transistor. Alternatively, however, the substrate may also be fully or partially identical with one or more of the above-mentioned components. Thus, as an example, the at least one channel may fully or partially be embodied within the substrate.

The at least one field-effect transistor may further have at least one sensing surface. The at least one sensing surface, as an example, may be a surface of the field-effect transistor which may be exposed to the fluid sample. The sensing surface, as an example, may be a surface of the multipurpose electrode, e.g. the above-mentioned electrode surface. The sensing surface, however, may also be or comprise another surface, such as a surface of the channel of the field-effect transistor. Various embodiments are feasible and will be described in an exemplary fashion in further detail below.

As used herein, the term "electrochemical measurement" may generally refer to the measurement of at least one measureable characteristic of a redox reaction. The electrochemical measurement and/or the measurable characteristic of the redox reaction, as an example, may imply an electrical current, a voltage, an electrical potential, a mass, for instance a mass deposited on an electrode, an impedance, particularly the real part and/or the imaginary part of the impedance, a capacitance, a resistance or a phase shift. Specifically, the electrochemical measurement may be performed in the presence of an electroactive species. As used herein, the term "electroactive species" may generally refer to a compound that facilitates or enhances or catalyzes the redox reaction, for instance by facilitating an electron transfer. The electroactive species may be dissolved in the fluid sample and/or may be immobilized on a surface of the analyte detector, wherein the surface may be exposable to the fluid sample. In particular, the surface may be the above-mentioned sensing surface and/or the above-mentioned surface of the multipurpose electrode. Preferred examples of electroactive species are redox mediators, specifically redox couples, such as but not limited to: potassium ferricyanid/ potassium ferrocyanide; hexaammineruthenium (II) chloride/ hexaammineruthenium (III) chloride; ferrocene methanol. 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. The electroactive species may facilitate or enhance the measurement of the at least one measurable characteristic of the redox reaction. As used herein, the term electrochemical measurement device may generally refer to an arbitrary device configured to perform at least one electrochemical measurement.

The term "electrochemical measurement device" may generally refer to an arbitrary device configured for performing at least one electrochemical measurement. For this purpose, as will be outlined in further detail and in an exemplary fashion below, the at least one electrochemical measurement device may comprise one or more electrical devices configured for performing the at least one electrochemical measurement. As an example, the electrochemical measurement device may comprise at least one electrical source, such as at least one electrical source selected from the group consisting of: a constant voltage source, a variable voltage source, a constant electrical current source, a variable electrical current source, a frequency generator for generating periodic electrical signals. Further, the electrochemical measurement device may comprise at least one electrical measurement device configured for measuring at least one electrical signal or electrical measurement variable, such as at least one electrical measurement device selected from the group consisting of: a voltage measurement device, a current measurement device, a potentiostat. Other measurement devices are feasible. The field-effect transistor specifically may not be part of the electrochemical measurement device. Thus, in other words, the analyte detector may comprise the field-effect transistor and the electrochemical measurement device as separate devices, consisting of separate components, except for the multipurpose electrode, which may be part of both the field-effect transistor and of the electrochemical measurement device. Thus, generally, the field-effect transistor and the electrochemical measurement device may form separate components of the analyte detector, except for the multipurpose electrode, which may form part of both the field-effect transistor and the electrochemical measurement device. Specifically, the transistor measurement by using the field-effect transistor and the electrochemical measurement by using the electrochemical measurement device may be distinct and separate measurements. The electrochemical measurement may be made without making use of the field-effect transistor.

The electrochemical measurement and/or the field-effect transistor-based measurement may take place in the presence of at least two different species of biorecognition molecules, for instance at least two different species of receptor molecules, namely at least one first receptor molecule and at least one secondary receptor molecule. The first receptor molecule and the secondary receptor molecule may be able to bind the analyte directly or indirectly. The first receptor molecule and the secondary receptor molecule may bind the analyte simultaneously. The secondary receptor molecule may enhance the electrochemical measurement and/or the field-effect transistor-based measurement, for instance by enhancing a signal and/or a selectivity of the electrochemical measurement and/or of the field-effect transistor-based measurement. The secondary receptor may enhance the signal and/or the selectivity on its own. Additionally or alternatively the secondary receptor may be labelled with at least one additional molecule, such as but not limited to an enzyme. The secondary receptor may affect or enhance the detection of the analyte by the analyte detector through an interaction with the analyte, e.g. through binding the analyte. The direct or indirect interaction of the secondary receptor with the analyte may affect or enhance the electrochemical measurement and/or the field-effect transistor-based measurement for instance by affecting or enhancing or producing a change in a concentration of a chemical species, such as but not limited to protons and/or electrons. The change in a concentration of a chemical species may correspond to a concentration of the analyte in the fluid sample. Thus, the secondary receptor may contribute to a signal enhancement of the analyte detector.

As outlined above, the electrochemical measurement device is configured for performing the at least one electrochemical measurement by using the at least one multipurpose electrode. Thus, the multipurpose electrode takes part in the electrochemical measurement. As an example, the at least one multipurpose electrode may be in electrical contact with the electrochemical measurement device, such as with the at least one electrical source and/or the at least one electrical measurement device discussed above. The at least one multipurpose electrode may be part of the at least one electrochemical measurement device and/or may be connected to the electrochemical measurement device, such as via at least one electrical connecting element, e.g. via at least one lead.

The multipurpose electrode may be in electrical contact with a gate electrode of the field-effect transistor. In particular, the gate electrode may be in direct or indirect physical contact with at least one channel of the field-effect transistor, specifically with at least one semiconducting layer. There may, for example, be a dielectric layer between the gate electrode and the channel, for instance to avoid leak current. In the case of a liquid-gated field-effect transistor, an ionic double layer may constitute the dielectric layer. In the embodiments just described, the gate electrode is typically in indirect physical contact with the channel of the field-effect transistor, specifically with the at least one semiconducting layer.

The multipurpose electrode may be at least partially identical with at least one element selected from the group of the gate electrode of the field-effect transistor and the channel of the field-effect transistor. The field-effect transistor may comprise at least one channel. Specifically, the at least one channel may be fully or partially made of at least one semiconducting material. A complete field-effect transistor typically comprises a semiconducting channel, metal source, drain and gate electrodes. Specifically, the gate electrode may be replaced by a reference electrode in solution or by a pseudoreference electrode, such as a metal electrode in solution. The semiconducting layer may comprise at least one material selected from the group consisting of: inorganic elemental semiconductors, inorganic compound semiconductors, and organic semiconductors. Specifically, the semiconducting layer may comprise at least one material selected from the group consisting of: graphene, a layered semiconductor, carbon nanotubes, and semiconducting nanowires. Further, the semiconducting layer may comprise at least one surface accessible to the analyte. In particular, the at least one surface accessible to the analyte may be functionalized by metal particles, specifically be metal particles comprising one or more metals selected from the group consisting of: gold and platinum. However, the use of other metals or alloys is also feasible.

The analyte detector may comprise at least one graphene layer interconnecting at least two electrically conductive electrodes, wherein the graphene layer may be accessible to the analyte, wherein the multipurpose electrode may comprise at least one element of the group consisting of: at least one of the at least two electrically conductive electrodes, the graphene layer. As an example, the semiconducting layer comprising, for instance, graphene may be the multipurpose electrode or may be part of the multipurpose electrode. In particular the graphene layer may be the multipurpose electrode or may be part of the multipurpose electrode. In such an embodiment at least one other electrode, specifically the source and/or the drain electrode, may serve to make contact to the semiconducting layer comprising, for instance to the graphene layer. The graphene layer may be at least partially covered by metal particles, specifically by gold particles.

The at least one multipurpose electrode may be in electrical contact with one or both of a source electrode or a drain electrode of the field-effect transistor. The multipurpose electrode may, for example, comprise the channel of the field-effect transistor. In this embodiment the source electrode and the drain electrode may serve to make contact to the multipurpose electrode. Alternatively, the multipurpose electrode may be fully of partially identical to one or more of: the source electrode; the drain electrode; the gate electrode.

The analyte detector comprises at least one further electrode exposable to the fluid sample. The at least one further electrode may comprise at least one electrode selected from the group consisting of a counter electrode and a reference electrode, wherein the electrochemical measurement device is configured for performing the at least one electrochemical measurement using the multipurpose electrode and the further electrode. The analyte detector may comprise at least three electrodes exposable to the fluid sample, wherein at least one of the at least three electrodes may be the multipurpose electrode. The multipurpose electrode may comprise gold. In particular, the analyte detector may comprise at least three electrodes, wherein all three electrodes may be gold electrodes. The multipurpose electrode may comprise at least one functional component exposed to its surface, wherein the at least one functional component may be configured for directly or indirectly interacting with the analyte. Further, the functional component may comprise at least one receptor compound, the receptor compound being capable of binding the at least one analyte. Specifically, the receptor compound being capable of binding the at least one analyte may be selected from the group consisting of: antibodies and fragments thereof, aptamers, peptides, enzymes, nucleic acids, receptor proteins or binding domains thereof and hydrophilic polymers capable of mediating a salting-out effect.

In particular, the at least one electrochemical measurement may comprise at least one measurement selected from the group consisting of: a cyclic voltammetry measurement; an impedance measurement; a potentiostatic measurement; an amperometric measurement; an electrochemical impedance spectroscopy; voltammetry; amperometry; potentiometry; coulometry. As used herein, the term "electrochemical impedance spectroscopy" may generally refer to the measurement of an impedance between the working electrode and the counter electrode as a function of a frequency of an electrical signal applied, such as a voltage and/or current. As further used herein, the term "voltammetry" may generally refer to the measurement of the current between the working electrode and the counter electrode as a function of the voltage applied. As used herein, the term "amperometry" may generally refer to the measurement of the current between working electrode and reference electrode, e.g. as a function of voltage. As used herein, the term "potentiometry" may generally refer to the measurement of the potential difference between the working electrode and the reference electrode. As used herein, the term "coulometry" may generally refer to the determination of the amount of charge produced or consumed during electrolysis. This may, for instance, be done by the measurement of a current between two electrodes, e.g. as a function of time.

Further, the at least one electrochemical measurement device may comprise at least one device selected from the group consisting of: a voltage source, a current source, a voltage meter, a current meter, an impedance meter, an impedance spectrometer, a frequency analyzer, a potentiostat, a frequency generator.

Furthermore, the electrochemical measurement device may be configured for measuring one or more of the following: an absolute value of an impedance between at least two electrodes of the analyte detector as a function of frequency and voltage applied, at least one of the electrodes being the multipurpose electrode; a real part of an impedance between at least two electrodes of the analyte detector as a function of frequency and voltage applied, at least one of the electrodes being the multipurpose electrode; an imaginary part of an impedance between at least two electrodes of the analyte detector as a function of frequency and voltage applied, at least one of the electrodes being the multipurpose electrode; a phase shift between a signal applied to at least one first electrode of the analyte detector and a signal response of at least one second electrode of the analyte detector, at least one of the first and second electrodes being the multipurpose electrode; an electrical current through the multipurpose electrode as a function of a periodic voltage applied to the multipurpose electrode; an electrostatic potential of the multipurpose electrode; an electrical current through the multipurpose electrode; and a voltage between the multipurpose electrode and at least one further electrode, specifically at least one counter electrode and/or at least one reference electrode.

The analyte detector further comprises at least one controller, wherein the controller may be connected to the field-effect transistor and to the electrochemical measurement device and wherein the controller is configured for controlling at least one transistor measurement by using the field-effect transistor and for controlling the at least one electrochemical measurement by using the electrochemical measurement device. In particular, the controller may be configured for controlling the at least one transistor measurement by measuring a drain current of the transistor. Furthermore, the controller may be configured for sequentially triggering at least one measurement using the field-effect transistor and the at least one electrochemical measurement. The controller may also be configured for repeatedly performing a sequence of the at least one measurement using the field-effect transistor and the at least one electrochemical measurement.

The analyte detector may further comprise at least one fluid channel, wherein the at least one multipurpose electrode may be disposed to be in contact with the fluid sample within the fluid channel. The fluid channel may comprise at least one fluid inlet for providing the at least one fluid sample to the fluid channel and at least one fluid outlet for disposal of the fluid sample from the fluid channel. In particular, the analyte detector further may comprise at least one external reference electrode being in fluid contact with the fluid channel, specifically at least one Ag/AgCl reference electrode.

The at least one multipurpose electrode may be at least partially covered by a membrane which may be permeable by the analyte. In particular the membrane may be a polymer membrane. Further, a space in between the membrane and the at least one multipurpose electrode may be at least partially filled by an electrolyte, for example a hydrogel electrolyte.

The at least one transistor may be selected from the group consisting of: an ion-sensitive field-effect transistor (ISFET); a chemically sensitive field-effect transistor (ChemFET); a biological field-effect transistor (BioFET); an enzyme field-effect transistor (ENFET); an extended-gate field-effect transistor (EGFET); a solution-, electrolyte-, water-, or liquid-gated FET.

In a second aspect, a method for detecting at least one analyte in a fluid sample is disclosed. With respect to definitions and embodiments of the method, reference can be made to definitions and embodiments of the analyte detector described above. The method is defined in claim <NUM>.

In particular, the method comprises using an analyte detector as described above or as will be further described below. Thus, as outlined above, specifically, the transistor measurement and the electrochemical measurement are distinct and separate measurements.

Specifically, the electrochemical measurement may be made without making use of the field-effect transistor. Specifically, the transistor measurement using the field-effect transistor and the electrochemical measurement, e.g. using the electrochemical measurement device, may be triggered sequentially, e.g. by using the controller. A sequence of the at least one transistor measurement using the field-effect transistor and the at least one electrochemical measurement may be repeatedly performed, e.g. by the controller.

In method step c) at least one transistor measurement value may be generated. Further, in method step d) at least one electrochemical measurement value may be generated. Specifically, the transistor measurement value and electrochemical measurement value may be combined for one or both of quantitatively or qualitatively detecting the at least one analyte in the fluid sample. Furthermore, method step d) may comprise at least one measurement selected from the group consisting of: a voltammetry measurement; an impedance measurement; a potentiostatic measurement; an amperometric measurement; a coulometric measurement.

In a third aspect, a use of the analyte detector as described above or as will be further described below for the qualitative and/or quantitative determination of the at least one analyte in a fluid is disclosed. The use is defined in claim <NUM>. In particular, said fluid may be selected from the group of fluids consisting of: body fluids, liquid or dissolved environmental samples and solutions of mixtures of chemical compounds. Specifically, said qualitative and/or quantitative determination of the at least one analyte in a fluid may be involved in diagnostic purposes, environmental control, food safety, quality control or manufacturing processes.

The invention further discloses and proposes (not part of the claimed subject-matter) a computer program including computer-executable instructions for performing the method according to the present invention in one or more of the embodiments enclosed herein when the program is executed on a computer or computer network. Specifically, the computer program may be stored on a computer-readable data carrier. Thus, specifically, one, more than one or even all of method steps c) and d) as indicated above may be performed and/or controlled and/or evaluated by using a computer or a computer network, preferably by using a computer program.

The invention further discloses and proposes (not part of the claimed subject-matter) a computer program product having program code means, in order to perform the method according to the present invention in one or more of the embodiments enclosed herein when the program is executed on a computer or computer network. Specifically, the program code means may be stored on a computer-readable data carrier.

Further, the invention discloses and proposes (not part of the claimed subject-matter) a data carrier having a data structure stored thereon, which, after loading into a computer or computer network, such as into a working memory or main memory of the computer or computer network, may execute the method according to one or more of the embodiments disclosed herein.

The invention further proposes and discloses (not part of the claimed subject-matter) a computer program product with program code means stored on a machine-readable carrier, in order to perform the method according to one or more of the embodiments disclosed herein, when the program is executed on a computer or computer network. As used herein, a computer program product refers to the program as a tradable product. The product may generally exist in an arbitrary format, such as in a paper format, or on a computer-readable data carrier. Specifically, the computer program product may be distributed over a data network.

Finally, the invention proposes and discloses (not part of the claimed subject-matter) a modulated data signal which contains instructions readable by a computer system or computer network, for performing the method according to one or more of the embodiments disclosed herein.

Preferably, referring to the computer-implemented examples, one or more of the method steps or even all of the method steps of the method according to one or more of the embodiments disclosed herein may be performed by using a computer or computer network. Thus, generally, any of the method steps including provision and/or manipulation of data may be performed by using a computer or computer network. Generally, these method steps may include any of the method steps, typically except for method steps requiring manual work, such as providing the samples and/or certain aspects of performing the actual measurements.

Specifically, the present invention further discloses (not part of the claimed subject-matter):.

The analyte detector, the use of the analyte detector and the method for detecting at least one analyte in at least one fluid sample according to the present invention presents a variety of advantages over prior art analyte detectors, their use and methods for detecting at least one analyte in at least one fluid sample. Thus, the analyte detector employs the multipurpose electrode for detecting one and the same analyte via both the transistor-based measurement using the FET and the electrochemical measurement using the electrochemical measurement device. Herein, a measurement range and/or a range of detection may vary between the transistor-based measurement and the electrochemical measurement. Thus, the ability to detect the analyte via the multipurpose electrode with one transistor-based and one electrochemical method may enhance the measurement range of the analyte detector. Specifically, the measurement range of the analyte detector may thus be enhanced by one or even several orders of magnitude. Generally, the present invention thus may allow for providing a single device or analyte detector which combines at least two principles of measurement in one and the same device and which may have an extended measurement range over conventional devices providing only one of these principles of measurement.

Furthermore, the ability to detect the analyte via the multipurpose electrode with one transistor-based and one electrochemical method may increase a measurement accuracy of the analyte detector. Specifically, a measurement range and/or a range of detection of the transistor-based and the electrochemical method may at least partially overlap. Thus, an averaging of detection results of the analyte by the analyte detector in at least parts of the overlapping detection ranges may increase a measurement accuracy of the analyte detector. Further, the provision of at least two different measurement methods with at least partially overlapping measuring ranges in one and the same device, i.e. the analyte detector, may serve as a fail-safe and/or back-up mechanism and thus increase reliability of the analyte detector.

Further optional features and embodiments of the invention will be disclosed in more detail in the subsequent description of preferred embodiments, preferably in conjunction with the dependent claims.

<FIG>, <FIG> and <FIG> show each an exemplary schematic layout of an analyte detector <NUM> for detecting at least one analyte in at least one fluid sample <NUM>. The analyte detector <NUM> comprises at least one multipurpose electrode <NUM> exposable to the fluid sample <NUM>, at least one field-effect transistor <NUM> in electrical contact with the at least one multipurpose electrode <NUM>, and at least one electrochemical measurement device <NUM> configured for performing at least one electrochemical measurement using the multipurpose electrode <NUM>.

The analyte detector <NUM> further comprises at least one controller <NUM>. The controller <NUM> is connected to the field-effect transistor <NUM> and to the electrochemical measurement device <NUM> and is configured for controlling at least one transistor measurement by using the field-effect transistor <NUM> and for controlling at least one electrochemical measurement by using the electrochemical measurement device <NUM>. The controller <NUM>, as an example, may be or may comprise at least one computer or processor, e.g. for timing and/or triggering the measurements and/or for reading out and/or evaluating measurement results. The controller may further comprise additional elements, such as one or more of a voltage source, a current source, a voltage measurement device, a current measurement device, a frequency generator or the like, as the skilled person will know when designing electrochemical measurements or transistor measurements.

As shown in <FIG>, the analyte detector <NUM> is configured for performing at least one electrochemical measurement and at least one field-effect transistor-based measurement. The multipurpose electrode <NUM> may be in electrical contact with a gate electrode <NUM> of the field-effect transistor <NUM> as shown in <FIG>. As shown in <FIG>, <FIG> and <FIG>, the field-effect transistor <NUM> may further comprise at least one source electrode <NUM>, at least one drain electrode <NUM> and at least one channel <NUM>.

The field-effect transistor <NUM> may be selected from the group consisting of an ion-sensitive field-effect transistor (ISFET); a chemically sensitive field-effect transistor (ChemFET); a biological field-effect transistor (BioFET); an enzyme field-effect transistor (EN-FET); an extended-gate field-effect transistor (EGFET) <NUM> as shown in <FIG> and a solution-, electrolyte-, water- or liquid-gated FET as shown in <FIG> and <FIG>. The gate electrode <NUM> and the drain electrode <NUM> may comprise gold. The analyte detector <NUM> may further comprise a substrate <NUM> as shown in <FIG>, <FIG>, <FIG>. The substrate <NUM> may comprise at least one element of the group consisting of glass, plastic, paper and silicon. The substrate <NUM> may comprise at least two layers as shown in <FIG>.

The channel <NUM> may be fully or partially made of at least one semiconducting material. Specifically the channel <NUM> may comprise at least one semiconducting layer <NUM>, as shown in the liquid-gated FET depicted in <FIG>. The semiconducting material, specifically the semiconducting layer <NUM>, may comprise at least one material selected from the group consisting of: inorganic elemental semiconductors, inorganic compound semiconductors and organic semiconductors, specifically at least one material selected from the group consisting of graphene, a layered semiconductor, carbon nanotubes and semiconducting nanowires. The semiconducting layer <NUM> may comprise at least one surface <NUM> accessible to the analyte. The at least one surface <NUM> may be functionalized by metal particles, specifically by metal particles comprising one or more metals selected from the group consisting of gold and platinum, as shown in <FIG>. Other metals or alloys are possible.

The gate electrode <NUM> may be in direct or indirect physical contact with the at least one channel <NUM> of the field-effect transistor <NUM>, as shown in <FIG>. The multipurpose electrode <NUM> may be at least partially identical with the extended gate electrode <NUM> of the field-effect transistor <NUM>, as depicted in <FIG>. Additionally or alternatively, the multipurpose electrode <NUM> may be at least partially identical with the channel <NUM> and/or the gate electrode <NUM> of the field-effect transistor <NUM>, as e.g. shown in <FIG> and <FIG>.

In <FIG>, an exemplary embodiment of the analyte detector <NUM> is shown in a partial view. The electrochemical measurement device <NUM> and the optional controller <NUM> are not shown in this Figure. As shown in <FIG>, the analyte detector <NUM> may further comprise a chamber <NUM>. The chamber <NUM> may comprise or consist of polydimethylsiloxane (PDMS). Other materials, specifically other plastic materials, are feasible. As also shown in <FIG>, the analyte detector <NUM> may further comprise a passivation layer <NUM>. The passivation layer <NUM> may comprise SU-<NUM>. SU-<NUM> is a negative, epoxy-type, near-UV photoresist based on EPON SU-<NUM> epoxy resin (from Shell Chemical) that has been originally developed, and patented (<CIT>) by IBM. Other materials, specifically other photoresists, are feasible. As depicted in <FIG>, <FIG> and <FIG>, the analyte detector <NUM> may comprise at least one fluid channel <NUM>. The multipurpose electrode <NUM> may be disposed to be in contact with the fluid sample <NUM> within the fluid channel <NUM>. The fluid channel <NUM> may further comprise at least one fluid inlet <NUM> for providing the at least one fluid sample <NUM> to the fluid channel <NUM> and at least one fluid outlet <NUM> for disposal of fluid sample <NUM> as shown in <FIG>. The analyte detector <NUM> may further comprise at least one external reference electrode <NUM>, specifically at least one Ag/AgCl reference electrode, which may be in fluid contact with the fluid channel <NUM> as depicted in <FIG> and <FIG>. The fluid channel <NUM> and/or the fluid inlet <NUM> and/or the fluid outlet <NUM> may be at least partially confined by a plastic material, specifically polytetrafluoreth-ylene (PTFE). Other materials are feasible, specifically other plastic materials.

<FIG> shows an analyte detector <NUM> with the field-effect transistor <NUM> implemented as an extended-gate field-effect transistor <NUM>. Again, the electrochemical measurement device <NUM> and the optional controller <NUM> are not shown in this Figure. The extended-gate field-effect transistor <NUM> may comprise an extended gate electrode <NUM>. The extended gate electrode <NUM> may comprise a substrate <NUM>. The substrate <NUM> may comprise at least one material selected from the group consisting of glass, plastic, paper and silicon. The extended-gate field-effect transistor <NUM> may be integrated together with the extended gate electrode <NUM> on the same substrate <NUM>. As depicted in <FIG> the extended gate electrode <NUM> may comprise gold, in particular a gold layer <NUM>. The gold layer <NUM> may be exposable to the fluid sample <NUM>. Additionally or alternatively the extended gate electrode may also comprise other metals and/or semiconducting materials including graphene, which may have a surface <NUM> exposable to the fluid sample <NUM>.

The multipurpose electrode <NUM> may comprise gold, as shown in <FIG>. The analyte detector <NUM> comprises at least one further electrode which may comprise at least one electrode selected from the group consisting of a counter electrode <NUM> and a reference electrode <NUM>, wherein the electrochemical measurement device <NUM> may be configured for performing the at least one electrochemical measurement using the multipurpose electrode <NUM> and the further electrode. The analyte detector <NUM> may comprise at least three electrodes exposable to the fluid sample <NUM>, wherein at least one of the at least three electrodes is the multipurpose electrode <NUM>. All three electrodes may be gold electrodes.

The analyte detector <NUM> may further comprise a reference electrode <NUM>, in particular an Ag/AgCl electrode. Other combinations are feasible.

<FIG> shows a measurement diagram <NUM> recorded with the analyte detector <NUM> of the type schematically depicted in <FIG> plotting a current Id as a function of a voltage Vref. The voltage Vref may also be denoted by Vg. Thus, Vref and Vg are used synonymously throughout the Figures and the description of the embodiments. The current Id is plotted once using a y-axis with a linear scale (y-axis on the left-hand side of diagram <NUM> in <FIG>) and once using a y-axis with a logarithmic scale (y-axis on the right-hand side). The graph relating to the y-axis on the left-hand side is denoted by <NUM>. The graph relating to the y-axis on the right-hand side is denoted by <NUM>. As can be seen from <FIG>, the drain current Id varies as a function of the potential Vref of the reference electrode <NUM> in a non-linear fashion. <FIG> shows a partial view of the analyte detector <NUM> of the type schematically depicted in <FIG> comprising an extended gate electrode <NUM> as part of an extended-gate field-effect transistor <NUM>. The extended gate electrode <NUM> shown in <FIG> comprises gold, in particular a gold layer <NUM>. The gold layer <NUM> is exposable to the fluid sample <NUM>. The extended gate electrode <NUM> as shown in <FIG> further comprises a substrate <NUM> carrying the gold layer <NUM>. The substrate <NUM> has three further gold layers <NUM> that may serve as further electrodes. The further electrodes may serve as control electrodes. The further electrodes may be used as additional multipurpose electrodes <NUM>. The multipurpose electrodes <NUM> and the at least one additional multipurpose electrode <NUM> may all be functionalized in the same way. The at least two multipurpose electrodes <NUM> functionalized in the same way may be used to enhance a precision of the field-effect transistor-based measurement. Alternatively, the at least two multipurpose electrodes <NUM> may be functionalized in the different ways, for example to allow a reference measurement. The extended gate electrode <NUM> shown in <FIG> is in electrical contact with the channel <NUM> of the field-effect transistor <NUM>. The extended-gate electrode FET <NUM> partially shown in <FIG> further comprises a source electrode <NUM> (not shown) and a drain electrode <NUM> (not shown).

<FIG> show measurement diagrams <NUM> recorded with the analyte detector <NUM> of the type depicted in <FIG>. <FIG> show measurement diagrams <NUM> recorded in a field-effect transistor-based measurement plotting the drain current Id as a function of the gate voltage Vg for three different measurements corresponding to <NUM> minutes (<NUM>), <NUM> minutes (<NUM>) and <NUM> minutes (<NUM>) of incubation in MES (<NUM>-(N-morpholino)ethanesulfonic acid) buffer. In <FIG>, for the most part the three graphs of the three different measurements overlap, being indistinguishable or hardly distinguishable from one another. <FIG> each show a stable and reproducible baseline in buffer for the field-effect transistor-based measurement. <FIG> is a measurement diagram <NUM> of an electrochemical measurement plotting a magnitude of an impedance Z (y-axis on the left-hand side <NUM>) and a phase shift angle θ (y-axis on the right-hand side <NUM>) as a function of frequency f of an alternating voltage for three different measurements corresponding to <NUM> minutes (<NUM>), <NUM> minutes (<NUM>) and <NUM> minutes (<NUM>) of incubation in buffer. The measurement diagram <NUM> shows three impedance measurements and three phase shift angle θ measurements. Again, both for the impedance and for the phase shift measurement, for the most part the three graphs of the three different measurements overlap, being indistinguishable or hardly distinguishable from one another.

The multipurpose electrode <NUM> may comprise at least one functional component <NUM> exposed to its surface <NUM>, as shown in <FIG> and <FIG>, wherein the at least one functional component <NUM> may be configured for interacting with the analyte. The functional component <NUM> may comprise at least one receptor compound being capable of binding the at least one analyte. The receptor compound being able of detecting the at least one analyte may be selected from the group consisting of: antibodies and fragments thereof, aptamers, peptides, enzymes, nucleic acids, receptor proteins or binding domains thereof and hydrophilic polymers capable of mediating a salting out effect. <FIG> shows a measurement diagram <NUM> displaying data recorded with an analyte detector <NUM> comprising an extended-gate field-effect transistor <NUM> comprising an extended gate electrode <NUM> that has a gold layer <NUM> exposable to the fluid sample <NUM>. In this case the multipurpose electrode <NUM> may comprise the extended gate electrode <NUM> comprising the gold layer <NUM>. <FIG> shows a field-effect transistor-based measurement plotting the drain current Id versus the gate voltage Vg for three different measurement situations, namely using an extended gate electrode <NUM> with gold layer <NUM> of bare gold <NUM>, using an extended gate electrode <NUM> with a gold layer <NUM> after immobilization of double stranded DNA <NUM> on the gold layer <NUM> and using an extended gate electrode <NUM> with a gold layer <NUM> after dehybridization of the double stranded DNA <NUM> on the gold layer <NUM>. As can be seen in <FIG>, the field-effect transistor-based measurement can clearly distinguish between the presence of double stranded DNA molecules on the extended gate electrode <NUM> (graph <NUM>) and single stranded DNA molecules on the extended gate electrode <NUM> after dehybridization (graph <NUM>).

The analyte detector <NUM> may also be used for an electrochemical measurement, for example for an impedance measurement that may be able to distinguish between the presence of single stranded DNA (graph <NUM>) and absence of single stranded DNA (graph <NUM>) on the gold layer <NUM> of the multipurpose electrode <NUM> as can be seen in the measurement diagram <NUM> in <FIG>. The measurement diagram <NUM> in <FIG> plots the negative imaginary part of the impedance Z'' versus the real part of the impedance Z' for the two different measuring situations just described.

<FIG> illustrate again the ability of the analyte detector <NUM> to distinguish between the presence of single stranded DNA <NUM> and the presence of double stranded DNA <NUM> both in a field-effect transistor-based measurement (<FIG>) and in an electrochemical measurement (<FIG> shows a measurement diagram <NUM> plotting the drain current Id versus the gate voltage Vg for two different measuring situations, namely the presence of single stranded DNA <NUM> as probe DNA on the gold layer <NUM> of the extended gate electrode <NUM> of the extended-gate field-effect transistor <NUM> and the presence of double stranded DNA <NUM> as a result of hybridization of single stranded probe DNA with single stranded target DNA after the addition of <NUM> of single stranded target DNA. Thus, in this example, the single stranded probe DNA serves as the functional component <NUM>. <FIG> plots the magnitude of the impedance Z and the phase shift angle θ as a function of the frequency f of the alternating voltage for the same two measuring situations as just described. The two graphs corresponding to the phase shift angle θ in the presence of single stranded DNA <NUM> and in the presence of double stranded DNA <NUM> are clearly distinguishable from one another.

<FIG> show measuring diagrams <NUM> based on field-effect transistor-based measurements using an extended-gate field-effect transistor <NUM> with an extended gate electrode <NUM> comprising a gold layer <NUM> with single stranded probe DNA immobilized on the gold layer <NUM>. <FIG> plots the drain current Id as a function of the potential Vref of the reference electrode <NUM> in the presence of <NUM> target DNA <NUM> and in the absence of target DNA <NUM>. The field-effect transistor-based measurement is clearly able to detect a potential shift ΔV in the presence of target DNA <NUM> as indicated by the arrow. <FIG> plots a potential shift ΔV as a function of target DNA concentration for two different ionic strengths of buffer solution (<NUM> and <NUM>). The size of the potential shift ΔV increases with increasing target DNA concentration and with increasing Debye length in lower ionic strength buffer.

<FIG> show measurement diagrams <NUM> based on electrochemical measurements carried out using the electrochemical measurement device <NUM> of the analyte detector <NUM> comprising a multipurpose electrode <NUM> with a gold layer <NUM> modified with an aminothiophenol monolayer <NUM> carrying an anti-TSH antibody, where TSH stands for thyroid stimulating hormone. <FIG> shows a cyclic voltammetry measurement plotting a current Icv between the multipurpose electrode <NUM> and the counter electrode <NUM> as a function of the voltage V applied for four different measuring situations, namely with the gold layer <NUM> of the multipurpose electrode <NUM> being either bare gold <NUM> or gold covered with an aminothiophenol monolayer <NUM>, or gold covered with an aminothiophenol monolayer <NUM> being additionally modified with anti TSH antibodies <NUM> or in the additional presence of <NUM> pM of TSH <NUM> in <NUM> MES buffer (pH = <NUM>). <FIG> plots the reactance X, i.e. the imaginary part of the impedance Z, versus the resistance R, i.e. the real part of the impedance Z, for the same four measuring situations as just described for <FIG>. Both the cyclic voltammetry measurement shown in <FIG> and the impedance measurement shown in <FIG> show that the analyte detector <NUM> may be able to detect as little as <NUM> pM of TSH, demonstrating a potential high sensitivity of the analyte detector <NUM> and its potential use in medical applications. <FIG> shows a schematic view of the layered modifications of the multipurpose electrode <NUM> as used in the electrochemical measurements of <FIG>, comprising the anti-TSH antibody <NUM> as the functional component <NUM>. Specifically, the anti-TSH antibodies that are used and/or described in this experiment or other experiments or in this embodiment or in other embodiments, specifically in <FIG>, <FIG> and <FIG>, may be anti-TSH F(ab')<NUM>-fragments of an anti-TSH antibody that may also be denoted as anti-TSH F(ab')<NUM> fragments.

<FIG> show measurement diagrams <NUM> plotting the reactance X, i.e. the imaginary part of the impedance Z, versus the resistance R, i.e. the real part of the impedance Z. In the case of <FIG>, the multipurpose electrode <NUM> has a gold layer <NUM> modified with a monolayer <NUM> and anti-TSH antibodies <NUM> as shown in <FIG>. The impedance measurement carried out between the multipurpose electrode <NUM> and the counter electrode <NUM> can distinguish clearly between five different concentrations of TSH. In the case of <FIG>, the impedance measurement is carried out between the multipurpose electrode <NUM> having a gold layer <NUM> carrying an aminothiophenol monolayer <NUM> (self-assembled SAM) and the counter electrode <NUM> in the absence of further modifications of the gold layer <NUM> of the multipurpose electrode <NUM>, thus in the absence of anti-TSH antibodies, and in the absence of TSH. The aminothiophenol monolayer <NUM> may be a self-assembled monolayer (SAM). A drift of the six impedance measurements may be due to the time elapsed during a measurement process.

<FIG> shows a measurement diagram <NUM> based on field-effect transistor-based measurements using a field-effect transistor <NUM> with an extended gate electrode <NUM> that has a gold layer <NUM> modified with anti-TSH F(ab')<NUM> antibody fragments <NUM> as functional components <NUM>. The fragments <NUM> are immobilized on the extended gate electrode <NUM> via short (<NUM> kDa) bifunctional PEG linker molecules (thiol and carboxyl groups to attach PEG to the gold layer <NUM> and to the anti-TSH F(ab') antibody fragments <NUM>, respectively). Additionally, a long (<NUM> kDa) monofunctional (thiolated) PEG is added to the gold layer <NUM> as a desalting agent. On the vertical axis, a voltage shift ΔV is depicted, given in millivolts (mV), as a function of the concentration c of the compound (analyte TSH or control sample BSA), given in mol/l, BSA standing for bovine serum albumin. The results depicted in <FIG> clearly show that the analyte detector <NUM> is able to distinguish between the presence of TSH and BSA and between different concentrations of TSH when the extended gate electrode <NUM> is modified with anti-TSH F(ab')<NUM> antibody fragments (<NUM>).

The analyte detector <NUM> comprises a field-effect transistor <NUM>. The multipurpose electrode <NUM> may be at least partially identical with at least one element selected from the group consisting of the gate electrode <NUM> of the field-effect transistor <NUM> and the channel <NUM> of the field-effect transistor <NUM>. <FIG> shows a field-effect transistor <NUM> implemented as a liquid-gated FET, comprising a source electrode <NUM> and a drain electrode <NUM>, a channel <NUM> and a gate electrode <NUM>. In this case, the gate electrode <NUM> in <FIG> comprises the reference electrode <NUM> and the conductive electrolyte solution <NUM>. The multipurpose electrode <NUM> in <FIG> is at least partially identical with the channel <NUM>. The channel <NUM> in <FIG> comprises a graphene layer <NUM>. Thus, in this case, the semiconducting layer <NUM> is identical to the graphene layer <NUM>. The graphene layer <NUM> comprises a surface <NUM> accessible to the analyte. A field-effect transistor <NUM> depicted in <FIG> further comprises a reference electrode <NUM>. <FIG> shows a partial view of a field-effect transistor <NUM> of the type schematically depicted in <FIG> shows a measurement diagram <NUM> recorded with a field-effect transistor <NUM> of the type schematically depicted in <FIG> plotting the current Id as a function of the voltage Vref for different pH values.

The analyte detector <NUM> comprises at least one electrochemical measurement device <NUM> configured for performing at least one electrochemical measurement using the multipurpose electrode <NUM>. The electrochemical measurement device <NUM> is not depicted in this Figure and may be added in electrical connection to the multipurpose electrode <NUM>. The electrochemical measurement may comprise at least one measurement selected from the group consisting of: a cyclic voltammetry measurement; an impedance measurement; a potentiostatic measurement; an amperometric measurement; an electrochemical impedance spectroscopy; voltammetry; amperometry; potentiometry; coulometry. <FIG> show measurement diagrams <NUM> based on electrochemical measurements, namely the cyclic voltammetry measurement (<FIG>) and an impedance measurement (<FIG>), where the graphene layer <NUM> described in <FIG> served as multipurpose electrode <NUM> in the electrochemical measurement. The measurement diagram <NUM> shown in <FIG> plots a current ICV measured between the multipurpose electrode <NUM> and the counter electrode <NUM> as a function of a voltage V applied. The measurement diagram <NUM> shown in <FIG> plots the negative imaginary part of the impedance Z" (the imaginary part also often being referred to as the "reactance" X) versus the real part of the impedance Z' (also often referred to as the "resistance" R). <FIG> show measurement diagrams <NUM> based on field-effect transistor-based measurements, carried out using the same multipurpose electrode <NUM> comprising the graphene layer <NUM> that served a multipurpose electrode <NUM> in the electrochemical measurement depicted in the measurement diagrams <NUM> in <FIG>. The data depicted in the measurement diagram <NUM> in <FIG> were recorded with a field-effect transistor <NUM> implemented as an ion-sensitive field-effect transistor <NUM>. The data depicted in the measurement diagram <NUM> in <FIG> were recorded using a field-effect transistor <NUM> implemented as an extended-gate field-effect transistor <NUM>. Both <FIG> show the drain current Id plotted as a function of the gate voltage VG.

The graphene layer <NUM> may be at least partially covered by metal particles <NUM>, specifically by gold particles <NUM> as can be seen in <FIG>. <FIG> show measurement diagrams <NUM> corresponding to the measurement diagrams shown in <FIG> with the graphene layer <NUM> of the multipurpose electrode <NUM> being partially covered by gold particles <NUM>. <FIG> show the graphene layer <NUM> with gold particles <NUM> deposited by physical adsorption overnight (<FIG>) and by <NUM> minutes of electrodeposition (<FIG>).

<FIG> show measurement diagrams <NUM> recorded with the analyte detector <NUM> using a multipurpose electrode <NUM> comprising a graphene layer <NUM> partially covered by gold particles <NUM>. As shown in <FIG>, gold particles <NUM> may be deposited on the graphene layer <NUM> by electrodeposition. The duration of the electrodeposition may be varied. <FIG> shows several field-effect transistor-based measurements carried out using a gate electrode <NUM> comprising a graphene layer <NUM> partially covered by gold particles <NUM>, where the duration of deposition of gold particles <NUM> on the graphene layer <NUM> varied between <NUM> and <NUM> minutes. Similarly, <FIG> shows several impedance measurements carried out using the multipurpose electrode <NUM> comprising a graphene layer <NUM> partially covered by gold particles <NUM>, where the duration of deposition of gold particles <NUM> on the graphene layer <NUM> varied between <NUM> and <NUM> minutes. <FIG> shows the corresponding measurement carried out after a deposition duration of <NUM> minutes.

<FIG> shows a measurement diagram <NUM> depicting an amperometric measurement. A current I between the multipurpose electrode <NUM> and the counter electrode <NUM> is plotted as a function of time t for three different measurements, where each measurement uses a multipurpose electrode <NUM> comprising a graphene layer <NUM>, where the graphene layer <NUM> either has no metal particles <NUM> deposited onto it (<NUM>') or the graphene layer <NUM> has gold particles <NUM> deposited onto it by electrodeposition for <NUM> minutes (<NUM>') or for <NUM> minutes (<NUM>'). The three amperometric measurements displayed in <FIG> were carried out in the presence of <NUM> different concentrations of an electroactive species [Fe(CN)<NUM>]<NUM>- : <NUM>, <NUM> and <NUM>. The current increases with increasing [Fe(CN)<NUM>]<NUM>- concentration. The observed current changes are more pronounced for the graphene layer <NUM> having gold particles <NUM> deposited onto it, with a deposition time of <NUM> minutes resulting in a higher sensitivity than a deposition time of <NUM> minutes.

Both field-effect transistor-based measurements and electrochemical measurements may be carried out in the presence of polyethylene glycol (PEG), specifically in the presence of pyrene PEG (P-PEG) and/or thiolated PEG (S-PEG) as shown in the measurement diagrams <NUM> in <FIG> shows a measurement diagram <NUM> of a field-effect transistor-based measurement plotting the drain current Id as a function of the gate voltage Vg. The measurement was carried out using a field-effect transistor <NUM> comprising a graphene layer <NUM> partially covered by the gold particles <NUM> deposited onto the graphene layer <NUM> by electrodeposition for <NUM> minutes. The field-effect transistor-based measurement was carried out either in the absence of PEG <NUM> or in the presence of pyrene PEG <NUM> or in the presence of thiolated PEG <NUM>. Similarly, <FIG> shows a measurement diagram <NUM> of an impedance measurement carried out either in the absence of PEG <NUM> or in the presence of pyrene PEG <NUM> or in the presence of thiolated PEG <NUM>.

The analyte detector <NUM> may be able to detect TSH and/or distinguish between different concentrations of TSH via the field-effect transistor-based measurement, as shown in <FIG>, <FIG> and <FIG> and/or via the electrochemical measurement, specifically via the impedance measurement as shown in <FIG>. The multipurpose electrode <NUM> may comprise a graphene layer <NUM> as shown in <FIG>. The graphene layer <NUM> may be prepared by a graphene transfer onto the multipurpose electrode <NUM>. The graphene layer <NUM> may be modified by the addition of PEG, in particular by the addition of thiolated PEG (S-PEG), specifically by the addition of short (<NUM> kDa) bifunctional carboxylated thiol PEG (SH-PEG-COOH) <NUM> , which may at least partially serve as binding sites for the functional component <NUM>, which may comprise anti-TSH antibodies <NUM>, in particular F(ab)'<NUM> TSH-antibody fragments, as shown in <FIG>. The graphene layer <NUM> may also be modified by co-immobilization of a long (<NUM> kDa) monofunctional methoxy-terminated thiol PEG (SH-PEG-OCH<NUM>) <NUM> which may increase the effective Debye length. The thiol groups of the PEGs may be attached to the pyrene linkers on the graphene layer <NUM> via maleimide chemistry. <FIG> and <FIG> show measurement diagrams <NUM> that are recorded using the multipurpose electrode <NUM> of the type depicted in <FIG>. In <FIG> the current Id is plotted as a function of the voltage Vref of the reference electrode <NUM>. <FIG> plots the voltage shift ΔVCNP as a function of the concentration c of TSH, which serves as the analyte in this experiment, or the concentration of BSA, which serves as a control sample in this experiment. Herein, CNP stands for charge neutrality point. The shift ΔVCNP depicted in <FIG> may be calculated from the measurement diagram <NUM> in <FIG> as the difference between the x-coordinate of the minimum of a graph corresponding to a sample containing TSA (<NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>) and the x-coordinate of graph <NUM> corresponding to a sample without TSH. The x-axis in <FIG> displays the concentration c on a logarithmic scale. <FIG> and <FIG> show that the analyte detector <NUM> may be able to distinguish clearly between at least six different concentrations of TSH using the field-effect transistor <NUM>. The multipurpose electrode <NUM> of the type depicted in <FIG> may also be used for the impedance measurement depicted in <FIG> shows a measurement diagram <NUM> plotting the reactance X versus the impedance R. <FIG> shows that the analyte detector <NUM> may be able to reproducibly detect <NUM> of TSH <NUM> using the electrochemical measurement device <NUM>, in particular the impedance measurement.

The analyte detector <NUM> may also be configured for the analysis of a gaseous analyte. In particular the analyte detector <NUM> may be configured for the analysis of at least one blood gas, specifically of CO<NUM>. <FIG> shows a partial view of an analyte detector <NUM>, whose multipurpose electrode <NUM> is at least partially covered by a membrane <NUM>. The membrane <NUM> may comprise or consist of PDMS. The membrane <NUM> may be at least partially permeable by the analyte. The analyte detector <NUM> as shown in the sectional view in <FIG> may further comprise a space <NUM> between the membrane <NUM> and the multipurpose electrode <NUM>. The space <NUM> may be partially filled with an electrolyte <NUM> as also shown in <FIG>, specifically with a hydrogel electrolyte, more specifically with a bicarbonate buffer dissolved in an agarose hydrogel. The multipurpose electrode <NUM> shown in <FIG> may have a surface <NUM> that may be sensitive to a pH. <FIG> shows a measurement diagram <NUM> plotting the current Id on a logarithmic scale as a function of the voltage Vref for five different CO<NUM> partial pressures that are given in the unit of mm Hg. The measurement shown in <FIG> is recorded in an extended-gate field-effect transistor-based measurement with the analyte detector <NUM> of the type depicted in 21A. <FIG> shows a measurement diagram <NUM> plotting the voltage shift ΔV as a function of the CO<NUM> partial pressure with the CO<NUM> partial pressure given on a logarithmic scale on the x-axis. The data points displayed in the measurement diagram <NUM> are derived from the data displayed in <FIG> and <FIG> show that the analyte detector <NUM> may be able to distinguish between deionized water <NUM> and at least five different CO<NUM> partial pressures <NUM>, <NUM>, <NUM> and <NUM>.

<FIG> show measurement diagrams <NUM> of an electrochemical measurement in form of a capacitance measurement (<FIG>) and of a transistor-based measurement (<FIG>) of TSH molecules <NUM> both measurements using the same multipurpose electrode <NUM>. Thus, in this specific example, the multipurpose electrode <NUM> is used to detect the analyte TSH <NUM> using two different measurement techniques, one transistor-based method and one electrochemical method, that make use of one and the same multipurpose electrode <NUM>. The employed multipurpose electrode <NUM> is of the type illustrated in <FIG> and described in the corresponding text passage. In particular, the multipurpose electrode <NUM> used for TSH molecule <NUM> detection as shown in the measurement diagrams <NUM> of <FIG> comprises the graphene layer <NUM> modified by short (approx. <NUM> kDa) bifunctional carboxylated PEG <NUM>, which may at least partially serve as binding sites for the functional component <NUM>, which in this case comprise anti-TSH antibodies <NUM>, in particular F(ab') <NUM> TSH-antibody fragments. The graphene layer <NUM> is further modified by co-immobilization of a long (<NUM> kDa) monofunctional methoxy-terminated thiol PEG (SH-PEG-OCH<NUM>) <NUM> which may increase the effective Debye length. <FIG> shows the capacitance C in µF as a function of a frequency f in Hz of an alternating voltage applied for five different situations, in which the fluid sample is either buffer without TSH molecules <NUM> or the fluid sample is buffer comprising one of the following TSH concentrations: <NUM> TSH, <NUM> TSH, <NUM> TSH or <NUM> TSH. The different TSH concentrations are marked by reference numbers <NUM>, <NUM>, <NUM>, <NUM> and <NUM>, respectively. <FIG> shows the capacitance C in µF as a function of the concentration of TSH molecules <NUM> in nm for a voltage alternating at <NUM>-<NUM> Hz, wherein the concentration of TSH molecules <NUM> is plotted in a logarithmic scale. The capacitance C of the buffer without TSH molecules <NUM> is indicated in <FIG> by a straight line marked with the corresponding reference number <NUM>. In <FIG> the current Id is plotted as a function of the gate voltage Vg of the gate electrode <NUM> of the FET <NUM> for seven different situations, in which the fluid sample is either buffer without TSH molecules <NUM> and without BSA or the fluid sample is buffer comprising <NUM> BSA and no TSH molecules <NUM> or the fluid sample is buffer comprising one of the following TSH concentrations and no BSA: <NUM> fM TSH, <NUM> pM TSH, <NUM> pM TSH, <NUM> pM TSH or <NUM> TSH. The different concentrations of TSH <NUM> and BSA are marked by reference numbers <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM> and <NUM>, respectively. The measurement setup as shown in <FIG> was used.

<FIG> plots the voltage shift ΔVCNP as a function of the concentration of TSH <NUM>. Herein, CNP stands for charge neutrality point. The shift ΔVCNP may be calculated from the measurement diagram <NUM> in <FIG> as the difference between the x-coordinate of the minimum of a graph corresponding to a sample containing TSH (<NUM>, <NUM>, <NUM>, <NUM> and <NUM>) or BSA (<NUM>) and the x-coordinate of the graph corresponding to the sample <NUM> without TSH and BSA. The x-axis in <FIG> displays the concentration of TSH on a logarithmic scale. The shift ΔVCNP of sample <NUM> containing buffer comprising <NUM> BSA and no TSH molecules <NUM> is indicated in <FIG> by a straight dashed line marked with the corresponding reference number <NUM>.

<FIG> and <FIG> show that by making use of both an electrochemical measurement and a transistor-based measurement a measuring range of the analyte detector <NUM> may be enhanced significantly as compared to analyte detectors making use of just one measurement method. In this particular example, the measurement range of the electrochemical measurement performed as capacitance measurement covers mainly the nanomolar (nM) range while the measurement range of the transistor-based measurement covers mainly the picomolar (pM) range. Thus, the use of the transistor-based measurement in addition to the electrochemical measurement may enlarge the measurement range of the analyte detector <NUM> substantially, for example by one or even several orders of magnitude.

<FIG> shows a further example of a transistor-based measurement of TSH <NUM> carried out with a field-effect transistor <NUM> of the same type as used for acquiring the data displayed in <FIG>. In <FIG> the current Id is plotted as a function of the gate voltage Vg of the gate electrode <NUM> of the FET <NUM> for eight different situations, in which the fluid sample is either buffer without TSH molecules <NUM> or the fluid sample is buffer comprising one of the following TSH concentrations: <NUM> fM TSH, <NUM> pM TSH, <NUM> pM TSH, <NUM> pM TSH, <NUM> pM TSH, <NUM> TSH or <NUM> TSH. The different TSH concentrations are marked by reference numbers <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM>, <NUM> and <NUM>, respectively. <FIG> plots the voltage shift ΔVCNP as a function of the concentration of TSH <NUM> for the data displayed in <FIG>, which constitutes a first set of measurements. The ΔVCNP values calculated from the data in <FIG> are marked with reference sign <NUM>. The shift ΔVCNP may be calculated from the measurement diagram <NUM> in <FIG> as described above for <FIG>. <FIG> further plots ΔVCNP values originating from a separate FET measurement, which constitutes a second set of measurements. The first set and the second set of measurements were carried out with a separate FET <NUM>. For the second set of measurements solely the calculated ΔVCNP values are plotted. These ΔVCNP values are marked with reference sign <NUM>. For the TSH concentration of <NUM> fM, which is marked by reference number <NUM>, the measurements of the first set of measurements <NUM> and the second set of measurements <NUM> overlap. The reference point for the formation of ΔVCNP was determined using a real-time measurement in this case and is not shown here.

<FIG> illustrate again that the FET-based measurement of the analyte detector <NUM> covers a measurement range comprising the picomolar range.

<FIG> show measurement diagrams <NUM> of an electrochemical measurement in form of a cyclic voltammetry measurement (<FIG>) and of a transistor-based measurement (<FIG>) of glucose. Again, both measurements employ the same multipurpose electrode <NUM>. Thus, in this specific example, the multipurpose electrode <NUM> is used to detect the analyte glucose using two different measurement techniques, one transistor-based method and one electrochemical method, which make use of one and the same multipurpose electrode <NUM>. For glucose detection a multipurpose electrode <NUM> comprising a graphene layer <NUM> is employed, wherein the graphene layer <NUM> is modified such that glucose dehydrogenase (GDH) is immobilized on the graphene layer <NUM>. preferably via <NUM>-pyrenebutyric acid N-hydroxysuccinimide ester (PBA-NHS) as linker. Immobilization of PBA-NHS preferably occurs by applying PBA-NHS, e.g. in a concentration of <NUM>, in a solvent, preferably ethanol, to the graphene layer <NUM>. In a further step GDH may be added preferably in a concentration of <NUM>/mL. Furthermore, a redox mediator such as ferrocenemethanol (FcMeOH) may be added. The PBA-NHS molecules immobilized on the graphene layer <NUM> may bind the GDH molecules thus immobilizing them on the graphene layer <NUM>. In the presence of GDH, glucose may be oxidized to glucolactone while flavin adenine dinucleotide (FAD) may be reduced to FADH<NUM>. In the presence of a redox mediator such as FcMeOH, FADH<NUM> may be oxidized to FAD while reducing the redox mediator such as FcMeOH, which may deliver electrons to the multipurpose electrode <NUM>. For glucose detection the multipurpose electrode <NUM> may be employed as the gate electrode <NUM> of the FET <NUM> and allow glucose detection via the FET-based measurement of the analyte detector <NUM> as shown in <FIG>. For the detection of glucose using the electrochemical measurement, the multipurpose electrode <NUM> may further be employed as working electrode e.g. for a cyclic voltammetry measurement as shown in <FIG>. The above-described redox chain reaction may allow to detect, for example, a current and/or a current density as a function of glucose concentration and the potential applied. <FIG> plots the current density j in µA/cm<NUM> against the applied potential E vs. the reference electrode (Ag/AgCl) for six different situations, in which the fluid sample is either buffer without glucose or the fluid sample is buffer comprising one of the following glucose concentrations: <NUM>, <NUM>, <NUM>, <NUM> or <NUM>. The different glucose concentrations are marked by reference numbers <NUM>, <NUM>, <NUM>, <NUM>, <NUM> and <NUM>, respectively. The data displayed in <FIG> constitute a first set of measurements <NUM> of the electrochemical glucose detection. <FIG> plots ΔI / Iblank as a function of the concentration of glucose for the data displayed in <FIG>, which constitutes the first set of measurements <NUM> of the electrochemical glucose detection. The concentration of a molecule may throughout this document be denoted by the molecule in squared brackets. <FIG> further plots ΔI / Iblank values originating from a separate cyclic voltammetry measurement, which constitutes a second set of measurements <NUM>. The first set of measurements <NUM> and the second set of measurements <NUM> were carried out with separate electrochemical measurement devices <NUM>. For the second set of measurements <NUM> solely the calculated ΔI / Iblank values are plotted. Herein, I is a current value corresponding to a current density j, Iblank denotes the current value produced by the sample <NUM> containing buffer without glucose and ΔI is the difference between the current value of a sample containing glucose and the current value of the sample <NUM> containing buffer without glucose when both current values are taken at the peak value (~<NUM> V).

Claim 1:
An analyte detector (<NUM>) for detecting at least one analyte in at least one fluid sample (<NUM>), the analyte detector (<NUM>) comprising at least one multipurposeelectrode (<NUM>) exposable to the fluid sample (<NUM>), wherein the multipurpose electrode (<NUM>) is configured to take part in analytical examinations based on at least one of method comprising the use of a field-effect transistor (<NUM>) and the at least one other method comprising the use of an electrochemical measurement device (<NUM>), wherein the term analytical examination is understood to describe a qualitative and/or a quantitative detection of the at least one analyte, the analyte detector (<NUM>) further comprising at least one field-effect transistor (<NUM>) in electrical contact with the at least one multipurpose electrode (<NUM>), wherein for being in electrical contact with each other the at least one field-effect transistor (<NUM>) and the at least one multipurpose electrode (<NUM>) are separate components or are fully or partially integrated into one another, the analyte detector (<NUM>) further comprising at least one electrochemical measurement device (<NUM>) configured for performing at least one electrochemical measurement using the multipurpose electrode (<NUM>), wherein the analyte detector (<NUM>) comprises at least one further electrode exposable to the fluid sample (<NUM>), wherein the analyte detector further (<NUM>) comprises at least one controller (<NUM>), wherein the controller (<NUM>) is connected to the field-effect transistor (<NUM>) and to the electrochemical measurement device (<NUM>) and wherein the controller (<NUM>) is configured for controlling at least one transistor measurement by using the field-effect transistor (<NUM>) and wherein the controller (<NUM>) additionally is configured for controlling the at least one electrochemical measurement by using the electrochemical measurement device (<NUM>), wherein the controller (<NUM>) is further configured to combine the transistor measurement value and the electrochemical measurement value for one or both of quantitatively or qualitatively detecting the at least one analyte in the fluid sample.