Organic electrochemical sensor with common electrode and measurement method

Abstract

The invention relates to an organic electrochemical sensor for detecting an analyte in a physiological sample, said sensor comprising at least two organic electrochemical transistors.

Description

TECHNICAL FIELD

[0001] The present invention relates to an organic electrochemical sensor. BACKGROUND

[0002] Heart failure is a chronic life-threatening disease that requires life-long management, 30% of patients are readmitted within 30 days of discharge. For most patients, the goal of treatment and medication is to improve the patient's quality of life, slow disease progression and avoid acute life-threatening events that require hospitalization. Optimization of drug dosage is individualized; it requires dose titration to balance between therapeutic efficacy and minimal side effects. Despite efforts to reduce hospitalization and readmission, these efforts focus on surrogate and late indicators of heart failure acute events, such as body weight and patient symptoms, rather than frequent blood testing to provide early indications that require adjustment of the patient's drug dosage.

[0003] Chronic kidney disease and kidney failure are also major public health problems affecting millions of people worldwide. These conditions are characterized by a gradual loss of kidney function, leading to the kidneys' inability to efficiently filter waste and excess water from the blood. If left untreated, these conditions can lead to end-stage renal disease (ESRD), requiring dialysis or kidney transplantation to survive. Prevention and management of kidney failure rely on awareness of underlying risk factors, early screening, and control.

[0004] However, known detection methods and devices for these diseases are not satisfactory.

[0005] Indeed, current blood tests rely only on clinical chemistry in a laboratory or point-of-care devices only available in specialized healthcare institutions. There is currently no solution to provide home monitoring of all key blood biomarkers to assess disease evolution in heart failure or kidney failure. Patients therefore need to regularly go to a hospital or a doctor's office for monitoring, which can be difficult and lead to a lack of adequate monitoring. In the worst case, patients are sent to a hospital only when symptoms appear, which means that it is already too late to detect worsening of heart failure or kidney failure without or with only minor sequelae.

[0006] The same problem arises in many other chronic diseases.

[0007] Organic electrochemical transistors (OECTs) have recently been developed. An OECT essentially comprises a source electrode and a drain electrode connected by a channel, and a gate electrode. By functionalizing the gate or the channel of an OECT, it can be used as a biosensor, i.e. it allows the detection and quantification of an analyte in a sample. However, known OECTs can only detect one analyte at a time. Therefore, multiple different functionalized OECTs are needed to diagnose a chronic disease such as heart failure or kidney failure, which leads to increased manufacturing costs and complexity for the patient.

[0008] Therefore, there is a need for a device suitable for detecting and measuring multiple analytes in a single sample with high sensitivity and high selectivity, in order to monitor chronic diseases such as heart failure or kidney failure. The device needs to be simple for the patient to use, while being compact and low cost to manufacture. SUMMARY

[0009] It is an object of the present invention to provide an organic electrochemical sensor to address one or more of the drawbacks of known devices, in particular for measuring and monitoring analytes of chronic diseases such as heart failure or kidney failure. The sensor comprises multiple source electrodes and drain electrodes connected by a channel, and multiple gate electrodes. Part of the channel and the gate electrodes are functionalized, such that different analytes can be measured in an un-preprocessed physiological sample, such as a single drop of blood.

[0010] To this end, the present invention relates to an organic electrochemical sensor for detecting an analyte in a physiological sample.

[0011] The sensor comprises:

[0012] - at least two groups of electrodes, each group of electrodes comprising a source electrode (S1, S2, S3, S4) and a drain electrode (D1, D2, D3, D4), the source electrode and the drain electrode of each group being connected by an organic conductive channel (C1, C2, C3, C4), and - at least two gate electrodes (G1, G2, G3, G4), wherein at least one channel (C1) and at least one gate electrode (G1) are not functionalized; wherein the channels (C2, C3, C4) and the gate electrodes (G2, G3, G4) fulfill at least two of the following three conditions: i. at least one channel (C2, C3, C4) or at least one gate electrode (G2, G3, G4) comprises a selective ion recognition layer, preferably a selective alkali metal ion recognition layer; ii. at least one of the channels (C2, C3, C4) or at least one of the gates (G2, G3, G4) comprises a Faraday recognition layer allowing the recognition of a specific analyte by a reduction or oxidation reaction at said channel or said gate; iii. at least one of the channels (C2, C3, C4) or at least one of the gates (G2, G3, G4) comprises an affinity recognition layer allowing the recognition of biomolecules in the sample by affinity; wherein any set of electrodes, channels (C1, C2, C3, C4) connected to said set of electrodes and any gate (G1, G2, G3, G4) form an organic electrochemical transistor.

[0013] According to an advantageous aspect of the application, the sensor comprises at least three gates (G1, G2, G3, G4).

[0014] In one embodiment, the channels (C2, C3, C4) and the gates (G2, G3, G4) satisfy said three conditions. In other words, the sensor comprises a selective ion recognition layer and a Faraday recognition layer and an affinity recognition layer.

[0015] Indeed, the sensor allows measuring at least two, preferably three, analytes in a single sample, these analytes having very different chemical properties. The measurement of at least two analytes allows for example the monitoring of the evolution of a chronic disease such as heart failure or kidney failure without the need for multiple samplings and / or analyses. Remote blood testing and direct measurement of biomarkers indicating a worsening of the health condition can provide data for sound decisions involving drug dose titration and management. Such monitoring enables practitioners to react quickly to changes in the patient's health status without the need for the patient to come to the office and before an emergency or hospitalization can be needed.

[0016] The use of functionalized channels and gates allows detecting analytes in a sample without pre-treatment. Indeed, the functionalization allows increasing the sensitivity to the target. Thus, the target can be detected in the original sample without pre-treatment such as increasing the concentration of said target. Thus, the sensor of the application is easy to use.

[0017] Moreover, the non-functionalized channels and gates allow performing reference measurements which are identical for different functionalized gates and channels. This reference measurement advantageously eliminates variations in the amount of offset between organic electrochemical transistors, thus enabling an optimized sensitivity.

[0018] According to an advantageous aspect of the application, the Faraday recognition layer is an enzymatic recognition layer, which allows the recognition of a specific analyte by an enzymatic reaction with the sample, preferably, the enzymatic recognition layer is configured to detect creatinine. Here, the enzymatic reaction comprises a reduction or oxidation step.

[0019] Due to the comorbidity nature of creatinine in heart failure events and the potential side effects of diuretics on the kidney of the patient, the detection and measurement of creatinine can detect and assess the risk of kidney injury or kidney failure.

[0020] In one embodiment, for each organic electrochemical transistor, at least one of the channels or gates constituting the organic electrochemical transistor is not functionalized.

[0021] In one embodiment, the sensor comprises: - at least two groups of electrodes, each group of electrodes comprising a source electrode and a drain electrode, the source electrode and the drain electrode of each group being connected by an organic conductive channel, and - at least three gates, wherein at least one channel and at least one gate are not functionalized; wherein at least one channel comprises a selective ion recognition layer, preferably a selective alkali metal ion recognition layer; wherein at least one gate comprises an enzymatic recognition layer, which allows the recognition of a specific analyte by an enzymatic reaction with the sample; wherein at least one gate comprises an affinity recognition layer, which allows the recognition of biomolecules in the sample by affinity; wherein any group of electrodes, the channel connected to the group of electrodes and any gate constitute an organic electrochemical transistor.

[0022] According to an advantageous aspect of the application, the affinity recognition layer comprises at least one of the following: a specific antibody, an aptamer, a nanobody or a molecularly imprinted polymer.

[0023] According to an advantageous aspect of the application, the affinity recognition layer is configured to detect NT-pro-BNP (N-terminal pro brain natriuretic peptide).

[0024] The detection and measurement of NT-pro-BNP is related to the physiological state of the myocardium and its pumping capacity. This biomolecule can therefore assess the probability of a heart failure event. It can also assess the probability of a kidney failure event.

[0025] According to an advantageous aspect of the application, at least one channel comprises an ion recognition layer configured to detect potassium ions.

[0026] Indeed, high levels of potassium are associated with attenuation of cardiac vascular elasticity, which can lead to arrhythmias and can lead to decompensation events. This ion can thus be used for the detection and monitoring of heart failure. It can also be used for the detection and monitoring of kidney failure.

[0027] According to an advantageous aspect of the application, the at least one channel comprises an ion recognition layer configured to detect sodium ions.

[0028] The measurement of sodium levels can monitor side effects that can be generated by drugs that prevent heart failure. Indeed, excessive intake of sodium is one of the main risk factors for the appearance of complications in patients with heart failure. The measurement of sodium levels is thus complementary to the direct detection of heart failure.

[0029] According to an advantageous aspect of the application, the enzyme recognition layer is configured to perform a multi-enzyme cascade.

[0030] The use of a multi-enzyme cascade can improve the sensitivity of the sensor to detect small molecules.

[0031] According to an advantageous aspect of the application, the organic electrochemical sensor further comprises a gate configured to directly detect creatine.

[0032] According to an advantageous aspect of the application, the Faradaic recognition layer for detecting creatine comprises a creatine kinase, it is thus an enzyme recognition layer in this case. It can distinguish between natural creatine present in the sample and creatine produced by the multi-enzyme cascade.

[0033] According to an advantageous aspect of the application, the at least two drain electrodes are common, and / or the at least two source electrodes are common.

[0034] This commonality allows to reduce the total number of electrodes in the sensor, thus increasing the compactness of the sensor and reducing the manufacturing costs. It also reduces the complexity of the sensor, which improves its robustness and simplifies the quality control.

[0035] In one embodiment, the organic electrochemical sensor comprises: one channel (C2) comprising a selective potassium ion recognition layer; one gate (G2) comprising an enzyme recognition layer configured to detect creatinine; one non-functionalized channel (C1); one non-functionalized gate (G1).

[0036] In one more specific embodiment, the organic electrochemical sensor comprises: one channel (C2) comprising a selective potassium ion recognition layer; one gate (G2) comprising an enzyme recognition layer configured to detect creatinine; one gate (G3) comprising an affinity recognition layer configured to detect NT-pro-BNP; one non-functionalized channel (C1 ); one non-functionalized gate (G1 ).

[0037] In a more particular embodiment, the organic electrochemical sensor comprises, one channel comprising a selective potassium ion recognition layer; one channel comprising a selective sodium ion recognition layer; one gate comprising an enzymatic recognition layer configured to detect creatinine; one gate comprising an affinity recognition layer configured to detect NT-pro-BNP; one non-functionalized channel; one non-functionalized gate. The organic electrochemical sensor can further comprise two non-functionalized channels (C1, C4) and one non-functionalized gate (G1 ).

[0038] The source electrodes (S1, S2, S3, S4) can comprise a pair of common source electrodes, each of the common source electrodes being used by two of the organic electrochemical transistors; the drain electrodes (D1, D2, D3, D4) can comprise a pair of common drain electrodes, each of the common drain electrodes being used by two of the organic electrochemical transistors. Comprising common source electrodes More particularly, the organic electrochemical sensor can further comprise one gate comprising an enzymatic recognition layer configured to detect creatine.

[0039] According to an advantageous aspect of the application, the size of the organic electrochemical sensor is less than 5 cm 2 .

[0040] Thanks to the proximity between the organic electrochemical transistors, it is possible to measure analytes in small volume samples.

[0041] The application also relates to an electronic device comprising an organic electrochemical sensor as described above.

[0042] Finally, the application also relates to a method for detecting analytes in a physiological sample, the method comprising the steps of: a. providing an organic electrochemical sensor as described above, b. placing the sample on the at least two channels and the at least three gates, c. for each organic electrochemical transistor comprising the enzymatic recognition layer or the affinity recognition layer: i. applying a voltage to the non-functionalized channel and to the non-functionalized gate, respectively; and to the non-functionalized channel and to the functionalized gate, ii. then measuring the output, which is a change in the drain electrode current value.

[0043] d. For each organic electrochemical transistor comprising the selective ion recognition layer: iii. applying a voltage to the unfunctionalized channel, to the unfunctionalized gate; and to the functionalized channel and the unfunctionalized gate, respectively, iv. then measuring the output, which is the difference in current values between the two drain electrodes.

[0044] It must be noted that the order of measurement of all OECTs of step c) is irrelevant. Similarly, the order of measurement of all OECTs of step d) is irrelevant. Finally, the order of steps c) and d) is irrelevant. In fact, all the required OECTs are used for measurement in any order. BRIEF DESCRIPTION OF DRAWINGS

[0045] Figure 1 The enzyme chain reaction for the detection of creatinine is shown. In the first step, creatinine is converted to creatine by creatininase (also called creatinine amino hydrolase or CI). In the second step, creatine is converted to sarcosine, and urea (not shown) by creatinase (also called creatine amino hydrolase or CII). In the third step, sarcosine is converted to hydrogen peroxide H2O2, and formaldehyde and glycine (not shown) by sarcosine oxidase (SOx).

[0046] Figure 2 An organic electrochemical sensor is shown, comprising 4 gates and 4 channels connecting 3 drain electrodes to 3 source electrodes.

[0047] Figure 3 An organic electrochemical sensor is shown, comprising 3 gates and 4 channels connecting 4 drain electrodes to 4 source electrodes. The drain and source electrodes are shared.

[0048] Figure 4 An organic electrochemical sensor is shown, comprising 4 gates and 4 channels connecting 4 drain electrodes to 4 source electrodes. One of the gates has a central position surrounded by the channels. The drain and source electrodes are shared.

[0049] Figure 5 An organic electrochemical sensor is shown, comprising 3 gates and 4 channels connecting 4 drain electrodes to 4 source electrodes. The drain and source electrodes are shared, except for the unfunctionalized channel.

[0050] Figure 6 An organic electrochemical sensor is shown, comprising 4 gates and 4 channels connecting 4 drain electrodes to 4 source electrodes. The drain and source electrodes are shared, except for the unfunctionalized channel.

[0051] Figure 7A-D is a set of graphs that illustrate the use of whole blood media. Figure 4 The organic chemical sensor measures the change in current. Figure 7A The diagram shows the current I as the concentration of NT-pro-BNP increases regularly from 0 pg / mL to 2500 pg / mL (indicated by the arrows). D,3 The change (in amperes) with time t (in seconds). Figure 7B The diagram shows the current I as the creatinine concentration increases regularly from 5 μM to 3000 μM (indicated by arrows). D,2 The change (in amperes) with time t (in seconds). Figure 7C The diagram shows the current difference I as the potassium ion concentration increases regularly from 2.0 mEq / L to 5.96 mEq / L (indicated by the vertical column). D,1-2 The change (in milliamperes) with time t (in seconds). Figure 7D The diagram shows the current difference I as the sodium ion concentration increases regularly from 100 mEq / L to 180 mEq / L (indicated by the vertical columns). D,1-3 The change (in milliamperes) with time t (in seconds).

[0052] definition In this invention, the following terms have the following meanings: "Analyte" refers to the substance to be detected. The analyte can be an ion or a biomolecule.

[0053] "Biomolecules" are a group of analytes produced by living organisms and essential for one or more typical biological processes. Biomolecules include large macromolecules such as proteins, carbohydrates, lipids, and nucleic acids, as well as small molecules such as vitamins and hormones.

[0054] "Detection" refers to identifying analytes in a sample (qualitative) or measuring the concentration of analytes in a sample (quantitative).

[0055] "Mutualization" refers to the use of the same electrode by two different OECTs so that a single electrical connection is used to apply voltage or measure current to the different electrodes. Specifically, the electrode can be the drain of one particular OECT and the source of another, with the same electrode used sequentially in both OECTs. In other words, the shared electrode is used by two organic electrochemical transistors.

[0056] "Unfunctionalized" refers to a channel or gate that does not contain any specific recognition layer. Unfunctionalized surfaces may be uncoated or coated with non-specific chemicals.

[0057] "Voltage applied to a channel" refers to the potential difference applied between the drain and source terminals connected by the channel.

[0058] "Voltage applied to a gate" refers to the potential difference applied between the gate and the source. Detailed Implementation

[0059] The following detailed description will be better understood when read in conjunction with the accompanying drawings. For illustrative purposes, an organic electrochemical sensor is shown in a preferred embodiment. However, it should be understood that this application is not limited to the precise arrangement, structure, features, embodiments, and aspects shown. The drawings are not drawn to scale and are not intended to limit the scope of the claims to the depicted embodiments.

[0060] This invention relates to an organic electrochemical sensor comprising an assembly constituting at least three organic electrochemical transistors (OECTs) in some arrangement. Known organic electrochemical transistors include a source electrode and a drain electrode connected by an organic conductive channel, and a gate electrode. The conductive channel may be an n-type or p-type polymer, such as PEDOT:PSS. The gate electrode may be organic (e.g., comprising PEDOT:PSS). The source electrode and drain electrode preferably comprise inorganic conductive compounds.

[0061] It should be noted that in the following instructions, the source electrode, drain electrode, and gate electrode may also be referred to simply as the source, drain, and gate.

[0062] The organic electrochemical sensor of the present invention includes at least two gates (G1, G2, G3, G4), preferably at least three gates (G1, G2, G3, G4), and at least two sets of electrodes. Each set of electrodes includes a source electrode (S1, S2, S3, S4) and a drain electrode (D1, D2, D3, D4) connected by conductive channels (C1, C2, C3, C4). The electrodes may be coated with an intermediate layer. For example, the intermediate layer is chosen to improve the adhesion between the electrode surface and the functional layer described below.

[0063] At least two of the entire set of channels (C2, C3, C4) and gates (G2, G3, G4) are functionalized, meaning they include a recognition layer on their surfaces. The recognition layer can be selected from... i. Selective ion recognition layer; ii. A Faraday recognition layer that allows for the identification of a specific analyte by reduction or oxidation of the analyte; or iii. Affinity recognition layer.

[0064] These identification layers can be disposed on the channels (C2, C3, C4) or gates (G2, G3, G4). In this invention, the channels (C2, C3, C4) and gates (G2, G3, G4) include at least two of the three types of identification layers. Preferably, the channels (C2, C3, C4) and gates (G2, G3, G4) include all three types of identification layers.

[0065] This functionalization allows for the specific sensing (e.g., detection, concentration measurement, etc.) of analytes (ions or biomolecules) in physiological samples placed on the sensor without any pretreatment of the sample. In practice, the gate electrode establishes an electrical connection with the channel via an electrolyte medium, which is the sample being analyzed and in which the analyte is sought. Thus, the sample behaves like an ion reservoir. When the sample is placed on the gate (G2, G3, G4) or channel (C2, C3, C4) including the recognition layer, ions from the ion reservoir are injected into or extracted from the gate (G2, G3, G4) (and correspondingly from the channel (C2, C3, C4)) due to the reaction or interaction between the recognition layer and the target analyte. This results in a change in the electron charge density of the channel, and consequently a change in the drain current, which can be measured using conventional electronic equipment. Therefore, the measured current ultimately indicates the presence of the target analyte interacting with the corresponding recognition layer.

[0066] Furthermore, at least one channel C1 and at least one gate G1 are unfunctionalized. "Unfunctionalized" means that the surfaces of channel C1 and gate G1 do not contain any specific recognition layer. However, the unfunctionalized surfaces may be coated with chemicals that are nonspecific to targeted monitoring. For example, gate G1 may include neutral proteins such as bovine serum albumin (BSA), Prionex™ reagent, casein, newborn calf serum (NBCS), or 2-mercaptoethanol to avoid any specific adsorption on gate G1. In another example, channel C1 may include an antibody that is nonspecific to the analyte to be monitored. Therefore, the unfunctionalized gate G1 and channel C1 provide a normalized (reference) signal, i.e., insensitive to any particular analyte. This reference signal allows for noise cancellation and improves sensor accuracy. Furthermore, the reference measurement advantageously eliminates offset variations between organic electrochemical transistors, thereby achieving optimized sensitivity. Preferably, in addition to functionalization, the unfunctionalized surface includes the same intermediate layer as the surface containing the recognition layer, allowing for direct comparison of measurements between the functionalized and unfunctionalized surfaces.

[0067] Therefore, the sensor of the present invention is configured such that any set of electrodes (one source (S1, S2, S3, S4) and one drain (D1, D2, D3, D4)), the channel (C1, C2, C3, C4) connected to the electrode set, and any gate electrode (G1, G2, G3, G4) constitute an organic electrochemical transistor (OECT). In other words, the source, drain, and gate electrodes are configured to achieve proper measurement of the sample through electrical connections with a set of source and drain electrodes and a gate. This effect cannot be achieved by simultaneously using multiple independent OECTs that are separately configured but submitted to the same sample. In fact, in the case of independent OECTs, the electrical connection of a set of source and drain electrodes from one OECT to a gate of another OECT means that there is a large distance between the gate and the channel connecting the source and drain electrodes. Therefore, the effects of changes in the state of the channel of the first OECT may not be well transmitted to the gate of the second OECT, for example, due to possible inhomogeneities in the sample or the resistivity of the electrolyte. The greater the distance between the channel and the gate, the worse the measurement accuracy. In other words, in the case of a standalone OECT, any electrode group, the channel connected to the electrode group, and any gate electrode do not constitute an OECT that allows the detection of analytes in a physiological sample. If both the channel (C1) and the gate (G1) are unfunctionalized, the OECT can be used for reference measurements of the sample. If the channel (C2, C3, C4) or the gate (G2, G3, G4) includes a recognition layer, the OECT is designed to measure a signal associated with a target in the recognition layer. In any case, in an OECT, one gate or one channel is unfunctionalized, and optionally both are unfunctionalized. The number of OECTs that can be formed is at least equal to the number of functionalized surfaces, wherein, optionally, at least one additional OECT for reference measurements is included. Therefore, the sensor of the present invention includes at least three OECTs.

[0068] It is important to note that the sensor of this invention is not a simple combination of multiple OECTs, each containing a functionalized gate or a functionalized channel. In fact, in this invention, the number of functionalized and unfunctionalized gates (G1, G2, G3, G4) and channels (C1, C2, C3, C4) is reduced to optimize device compactness. In other words, different organic electrochemical transistors can be formed that share the unfunctionalized gate G1 and channel C1 to reduce the total number of electrodes.

[0069] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - Including at least one channel (C2, C3, C4) surface and / or at least one gate (G2, G3, G4) surface of the selective ion recognition layer; and - Including at least one gate (G2, G3, G4) surface and / or at least one channel (C2, C3, C4) surface of the Faraday recognition layer.

[0070] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - Including at least one channel (C2, C3, C4) surface and / or at least one gate (G2, G3, G4) surface of the selective ion recognition layer; - Including at least one gate (G2, G3, G4) surface and / or at least one channel (C2, C3, C4) surface of the Faraday recognition layer; and - Including at least one gate (G2, G3, G4) surface and / or at least one channel (C2, C3, C4) surface of the affinity recognition layer.

[0071] This configuration allows for an optimal combination of a channel with high transconductance and a gate with high capacitance on a minimal sensor surface. This combination can be achieved using low-cost screen printing and inkjet printing manufacturing techniques.

[0072] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - Including at least two channel (C2, C3, C4) surfaces and / or at least two gate (G2, G3, G4) surfaces of the selective ion recognition layer; - Including at least one gate (G2, G3, G4) surface and / or at least one channel (C2, C3, C4) surface of the Faraday recognition layer; and - Including at least one gate (G2, G3, G4) surface and / or at least one channel (C2, C3, C4) surface of the affinity recognition layer.

[0073] The identification layer can completely cover the corresponding gate (G2, G3, G4) or the corresponding channel (C2, C3, C4).

[0074] The selective ion recognition layer allows for the selection of a specific ionic chemical substance, such as an atomic ion or a molecular ion. This is based on the electrical interaction between oppositely charged ions present in the sample and the functionalized surface. The ions can be alkali metal ions, such as sodium ions (Na+). + ), potassium ions (K) + ) or lithium ion (Li + ), or other ions, such as chloride ions (Cl ions) - ), calcium ions (Ca 2+ ), ... Measuring sodium ion concentration in blood is challenging because it has a very narrow concentration range and very high values ​​(the normal range in human blood is 135–145 mEq / L, see [reference]). Figure 5 (D). Therefore, both sensitivity resolution and saturation point must be very high. Thus, covering the channels (C2, C3, C4) with sodium ion recognition layers allows the OECT response to vary linearly with sodium concentration (e.g., 100–180 mEq / L). Measuring sodium concentration allows monitoring of potential side effects from medications used to prevent, for example, heart failure. Therefore, measuring sodium levels is complementary to the detection, prevention, and monitoring of heart failure.

[0075] Measuring potassium ion concentration is also challenging because its concentration in blood is relatively low compared to other electrolytes (normal range in whole blood is 3.5–5.1 mEq / L, see [reference]). Figure 5 (C) Therefore, high selectivity is essential. Thus, covering the channels (C2, C3, C4) with potassium ion recognition layers allows the OECT response to vary linearly with potassium concentration (e.g., 2.5–6 mEq / L). High potassium levels are associated with decreased elasticity of the heart and blood vessels, which can lead to arrhythmias and potentially decompensated events. Therefore, this analyte can be used for the detection, prevention, and monitoring of heart failure. In the case of chronic kidney disease, potassium imbalance is associated with disordered renal function and drug effects, which can lead to hypokalemia or hyperkalemia. This is why monitoring potassium levels is crucial to prevent further deterioration of the patient's health.

[0076] The Faraday recognition layer is the layer in which the analyte undergoes reduction or oxidation. The resulting current alters the electronic properties of the OECT sensor, enabling electrical measurements.

[0077] In one embodiment, the Faraday recognition layer is an enzyme recognition layer, in which a single enzyme or several enzymes are used to assist in the reduction or oxidation of the analyte, optionally acting in a multi-enzyme cascade reaction. The enzyme recognition layer allows the recognition of a specific analyte by an enzymatic reaction with the sample. More precisely, it allows the recognition of hydrolyzable molecules. For example, the enzyme recognition layer is configured to detect creatinine. For this purpose, the enzyme recognition layer may include creatinine anhydrase or creatinine amide hydrolase. Creatine anhydrase (also known as creatinine deaminase) and creatinine amide hydrolase are enzymes that hydrolyze creatinine to creatine. Due to its comorbid nature, the measurement of creatinine in the blood allows for the assessment of the risk of kidney failure. However, detecting creatinine through direct interaction with the recognition layer is challenging due to the lack of reliable recognition elements. Therefore, the enzyme recognition layer can be configured to perform a multi-enzyme cascade reaction. For example, a multi-enzyme cascade using three or four enzymes allows the degradation of creatinine to creatine, the first step of which is the hydrolysis of creatinine to creatine.

[0078] Figure 1An example of a multi-enzyme cascade is shown. This cascade can convert existing creatinine into H₂O₂, which can then be quantified electrochemically in conjunction with an enzyme to obtain a signal corresponding to the creatinine concentration. The multi-enzyme cascade may include a mediator. The mediator's role is to facilitate the transfer of any electrons generated due to the electrochemical reaction to the OECT electrode. The intermediate can be selected from the group consisting of 2,2'-azono-bis(3-ethylbenzothiazoline-6-sulfonic acid), ferrocene, ferrocene monocarboxylic acid, iron / ferrocyanide, 2,6-dichlorophenolindophenol, hexaaminoruthenium, 1,4-naphthoquinone, 7-hydroxy-3H-phenoxazin-3-one (also known as halogen), 1,1'-bis(cyanomethyl)-4,4'-bipyridine (also known as cyanomethyl viologen), 1,1'-bis(ethyl)-4,4'-bipyridine (also known as ethyl viologen), riboflavin 5'-monophosphate, 1,1'-ethylene-2,2'-bipyridine (also known as diquat), 1,1'-trimethylene-2,2'-bipyridine (also known as terbufos), or 4,4'-bipyridine-1,1'-bis(2-ethylsulfonate). Ferrocene is the preferred intermediate. Figure 1 In this example, the enzyme is HRP (horseradish peroxidase).

[0079] A challenge when using multi-enzyme cascade reactions is distinguishing between native creatine (present in blood samples) and creatine produced during the cascade, the latter of which must be measured to determine the concentration of creatinine. To overcome this challenge, one solution is to measure only the creatine produced by the degradation of creatinine. This can be achieved in two ways. First, at the entrance to the sample deposition area, the native creatine in the blood is degraded before reaching the enzyme recognition layer. For this purpose, the enzyme recognition layer for detecting creatine includes, for example, creatine kinase. Second, an additional reference electrode is included on the sensor, or a non-functionalized electrode is used as the reference electrode. Preferably, the reference electrode comprises the same intermediate layer as the electrode containing the enzyme recognition layer for detecting creatinine, but does not include the enzyme that allows creatinine to be hydrolyzed to creatine. Therefore, the signal difference between the electrode containing the enzyme recognition layer and the reference electrode allows for the measurement of creatinine produced by the multi-enzyme cascade reaction. Another solution to overcome this difficulty is that one of the gates G4 is configured to directly detect native creatine, thus subtracting the concentration of native creatine from the total concentration of creatine measured after the multi-enzyme cascade reaction to determine the concentration of creatine produced by the multi-enzyme cascade reaction. The sensor of this invention allows the detection of creatinine in whole blood media at concentrations ranging from 0 to 3000 μM. See [link to related document]. Figure 5 B.

[0080] Affinity recognition layers allow for the recognition of specific biomolecules in a sample through affinity. Therefore, affinity recognition layers differ from ion recognition layers. In other words, regardless of their functionalization, ion recognition layers are not affinity recognition layers.

[0081] More specifically, the affinity recognition layer allows the recognition of non-hydrolyzable molecules using specific antibodies (antibody fragments, antibody dendritic macromolecule conjugates, nanobodies) or aptamer interactions, or engineered binding proteins or nanobody interactions on the electrode surface, or molecularly imprinted polymers or deoxyribonucleic acid (DNA) or ribonucleic acid (RNA). This structural interaction is preferably a protein-protein interaction, where amino acids are configured to capture specific proteins. The molecule may contain binding domains or sequences configured to specifically interact with the biomolecule.

[0082] For example, the affinity recognition layer is configured to recognize the N-terminus of brain natriuretic peptide B (NT-pro-BNP). Detection of NT-pro-BNP is challenging because it is present in the blood at very low levels (between less than 100 pg / mL and 450 pg / mL in healthy patients, but potentially as high as 20,000 pg / mL in very severe cases). Using specific antibody interactions to recognize NT-pro-BNP advantageously enables OECT to respond linearly to peptide concentrations (e.g., from 0 to 300 pg / mL). The combination of two linear responses allows for detection in the range of 5 pg / mL to 30,000 pg / mL; see [link to relevant documentation]. Figure 5 A. An NT-pro-BNP level greater than 450 pg / mL reflects impaired myocardial pumping ability. An NT-pro-BNP level greater than 560 pg / mL is a predictor of mortality risk in chronic kidney disease.

[0083] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - At least one channel (C2, C3, C4) surface, comprising a selective ion recognition layer configured to measure potassium concentration; and - At least one gate (G2, G3, G4) surface, which includes a Faraday recognition layer for detecting creatinine.

[0084] This configuration can be used to monitor kidney failure.

[0085] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - At least one channel (C2, C3, C4) surface, which includes a selective ion recognition layer configured to measure potassium concentration; - At least one gate (G2, G3, G4) surface, comprising a Faraday recognition layer for detecting creatinine; and - At least one gate (G2, G3, G4) surface, which includes an affinity recognition layer for detecting NT-pro-BNP.

[0086] This configuration can be used to monitor kidney failure.

[0087] In one embodiment, the functionalized surface of the organic electrochemical sensor of the present invention is: - At least one channel (C2, C3, C4) surface, which includes a selective ion recognition layer configured to measure potassium concentration; - At least one channel (C2, C3, C4) surface, which includes a selective ion recognition layer configured to measure sodium concentration; - At least one gate (G2, G3, G4) surface, comprising a Faraday recognition layer for detecting creatinine; and - At least one gate (G2, G3, G4) surface, which includes an affinity recognition layer for detecting NT-pro-BNP.

[0088] This configuration can be used to monitor heart failure.

[0089] An electrical connection E can be arranged on the sensor to facilitate the application of voltage to different electrodes or the measurement of current.

[0090] In one embodiment, the source electrodes (S1, S2, S3, S4) are shared. In another embodiment compatible with the previous embodiment, the drain electrodes (D1, D2, D3, D4) are shared. The sharing (or common use) of the source electrodes (S1, S2, S3, S4) and drain electrodes (D1, D2, D3, D4) means that the source electrodes (S1, S2, S3, S4) or drain electrodes (D1, D2, D3, D4) are electrically connected, such that a voltage can be applied to the shared electrode or a current from the shared electrode can be measured through the same electrical connection E. For example, in… Figure 3 and 4 In this design, two pairs of source electrodes ((S1, S2) and (S3, S4)) and two drain electrodes (D2, D3) are shared. The drain and source electrodes can also be interdigitated. This sharing allows for a reduction in the total number of electrodes in the sensor (from 8 to 5 in this case), thereby improving sensor compactness and reducing manufacturing costs. This also allows for the use of small user samples. Furthermore, it reduces sensor complexity, improves its robustness, and simplifies its quality control.

[0091] Figure 2An organic electrochemical sensor is shown, comprising four gates (G1, G2, G3, G4) and three channels (C1, C2, C3) connecting three drains (D1, D2, D3) to three sources (S1, S2, S3), wherein the drains and sources are interdigitated electrodes. In this embodiment, one channel C2 includes a sodium ion recognition layer, another channel C3 includes a potassium ion recognition layer, one gate G2 contains an affinity recognition layer configured to recognize NT-pro-BNP, and another gate G3 contains an enzyme recognition layer configured to detect creatinine. The other channel C1 and the other two gates (G1, G4) are not functionalized.

[0092] Figure 3 An organic electrochemical sensor is shown, comprising three gates (G1, G2, G3) and four channels (C1, C2, C3, C4) (not shown, but located above the drain and source pairs) connecting four drains (D1, D2, D3, D4) to four sources (S1, S2, S3, S4), where the drains and sources are interdigitated electrodes. As described above, multiple source electrodes (S1, S2 and S3, S4) and drain electrodes (D2, D3) are shared. The channel geometry of this embodiment allows for improved output intensity and signal-to-noise ratio. In this embodiment, one channel C2 includes a sodium ion recognition layer, another channel C3 includes a potassium ion recognition layer, one gate G2 includes an affinity recognition layer configured to recognize NT-pro-BNP, and another gate G3 includes an enzyme recognition layer configured to detect creatinine. The other two channels (C1, C4) and the remaining gate G1 are not functionalized.

[0093] Figure 4 An organic electrochemical sensor is shown, comprising four gates (G1, G2, G3, G4) and four channels (C1, C2, C3, C4) (not shown, but located above the drain and source pairs) connecting four drains (D1, D2, D3, D4) to four sources (S1, S2, S3, S4). The drains and sources are interdigitated electrodes. As described above, multiple source electrodes (S1, S2 and S3, S4) and drain electrodes (D2, D3) are shared. One gate, G4, is located at the center surrounded by the channels. This geometry allows for increased sensitivity. In this embodiment, two gates (G1, G4) are unfunctionalized, compared to the previous embodiment.

[0094] Figure 5An organic electrochemical sensor is shown, comprising three gates (G1, G2, G3) and four channels (C1, C2, C3, C4) (not shown, but located above the drain and source pairs) connecting four drains (D1, D2, D3, D4) to four sources (S1, S2, S3, S4), where the drains and sources are interdigitated electrodes. As described above, the multiple source electrodes (S2 and S4) and drain electrodes (D2 and D3) are shared. More specifically, the sources (S1, S2, S3, S4) comprise a pair of shared sources, such that each of these shared sources is used by two of the organic electrochemical transistors; the drains (D1, D2, D3, D4) comprise a pair of shared drains, such that each of these shared drains is used by two of the organic electrochemical transistors. In this embodiment, one channel C2 includes a sodium ion recognition layer, another channel C3 includes a potassium ion recognition layer, one gate G2 includes an affinity recognition layer configured to recognize NT-pro-BNP, and another gate G3 includes an enzyme recognition layer configured to detect creatinine. The remaining channel C1 and the remaining gate G1 are not functionalized. Keeping the electrodes (S1 and D1) of an unfunctionalized channel separate, i.e., independent, allows for the avoidance of additional capacitance that could affect the measurement.

[0095] Figure 6 An organic electrochemical sensor is shown, comprising four gates (G1, G2, G3) and four channels (C1, C2, C3, C4) (not shown, but located above the drain and source pairs) connecting four drains (D1, D2, D3, D4) to four sources (S1, S2, S3, S4), where the drains and sources are interdigitated electrodes. As described above, multiple source electrodes (S2 and S4) and drain electrodes (D2 and D3) are shared. More specifically, the sources (S1, S2, S3, S4) comprise a pair of shared sources, such that each of these shared sources is used by two of the organic electrochemical transistors; the drains (D1, D2, D3, D4) comprise a pair of shared drains, such that each of these shared drains is used by two of the organic electrochemical transistors. One channel C1 and one gate G1 are unfunctionalized. Maintaining a separate, unfunctionalized channel for the electrodes (S1 and D1) avoids the need for additional capacitance that could affect the measurement. One of the gates, G4, is located at the center of the channel. This geometry allows for improved sensitivity.

[0096] This last embodiment is advantageous. In fact, the optimized geometry allows for a reduced surface area, resulting in a smaller required user sample volume and making the sensor suitable for use with capillary blood samples. Therefore, the placement of the active sensing element allows for rapid and specific detection of multiple analytes in only a single sample. Finally, this geometry also eliminates the need for microfluidic implementation, thus reducing fabrication complexity and cost.

[0097] In all of the above embodiments, the sensor may include the continuous layer described below.

[0098] An insulating substrate can be provided as a support. This substrate can be selected from plastics, paper, and glass. The substrate is preferably a flexible polymer substrate, such as polyethylene terephthalate (PET), polyimide (Kapton™), or a paper substrate.

[0099] An electrical layer can be deposited on the substrate. This electrical layer includes source electrodes (S1, S2, S3, S4), drain electrodes (D1, D2, D3, D4), gate electrodes (G1, G2, G3, G4), and channels (C1, C2, C3, C4), as well as electrical connections E and conductive traces T between them. The source electrodes (S1, S2, S3, S4), drain electrodes (D1, D2, D3, D4), and gate electrodes (G1, G2, G3, G4) are metallic or non-metallic electrodes deposited on the substrate. Gold and its alloys are particularly suitable for the electrodes. The electrical layer can also be a sandwich layer, such as Ag-C-Ag / AgCl, Ag-Ag / AgCl, or Ag-C. The source electrodes (S1, S2, S3, S4) and drain electrodes (D1, D2, D3, D4) are preferably printed using silver and carbon screen printing inks, which consist of a mixture of silver or carbon with various polymers to achieve optimal flowability. This allows the electrodes to contain an organic non-conductive matrix and conductive inorganic components. The ink can be a protective conductive ink to prevent the electrodes from being oxidized by the sample.

[0100] The channels (C2, C3, C4) may include a first intermediate conductive layer to connect the source (S2, S3, S4) and drain (D2, D3, D4). This intermediate layer ensures effective operation of the ion sensor, thereby improving the sensor's signal and sensitivity. An ion-selective recognition layer can then be deposited on the intermediate conductive layer. The ion-selective recognition layer may include an ion carrier (i.e., selective for sodium or potassium).

[0101] The gates (G2, G3, G4) may comprise a covalently self-assembled monolayer. A recognition layer may then be deposited on this monolayer.

[0102] In the case of enzyme recognition layers, the enzyme is either chemically cross-linked to the mediator solution before being deposited onto the monolayer, or it is non-covalently bound to the surface by drying the enzyme on the surface, then connecting the surface to the cellulose membrane and placing it on the electrode surface.

[0103] The sensor may also include a filter membrane above the recognition layer of the gates (G2, G3, G4). The membrane may also cover the entire electrode and recognition layer. This membrane acts as a filter between the placed sample and the electrode. In particular, this membrane is suitable for untreated blood samples, allowing only plasma to contact the recognition layer, thus reducing contamination and improving the OECT response and signal-to-noise ratio. Suitable selective membranes are commercially available, such as LF1, MF1, VF1, and VF2 from Waterman International LLC (Medstone, UK).

[0104] Finally, dielectric ink is printed onto the surface of the sensor, leaving only the surfaces of the channels and functional electrodes uncovered.

[0105] All of the above layers can be deposited by any suitable physical or chemical method, including inkjet printing, droplet coating, roll-to-roll, screen printing, spin coating, and vacuum physical and chemical deposition.

[0106] In one embodiment, the organic electrochemical sensor is less than 5 cm in size. 2 Preferably less than 3cm 2 Preferably less than 2cm 2 The dimension refers to the area of ​​the convex shape surrounding all the gates (G1, G2, G3, G4) and the OECT arrangement of the sensor. This dimension corresponds to the area of ​​the sample that needs to be spread out for proper analysis. Due to the proximity of the organic electrochemical transistors, this advantageously allows for the measurement of analytes in small-volume samples. For example, the sample volume can be approximately 100 μL (e.g., in a 2 cm³ sample). 2 (A sample layer up to 500 μm thick) or even less than 40 μL.

[0107] Electrodes can have various sizes, depending on size and precision limitations. The lateral dimensions (length and width) of the source electrodes (S1, S2, S3, S4) and drain electrodes (D1, D2, D3, D4) can range from 100 nm to 2 cm, preferably from 1 μm to 1.5 cm, and more preferably from 50 μm to 1 cm. The lateral dimensions (length and width) of the gate electrodes (G1, G2, G3, G4) can range from 100 nm to 5 cm, preferably from 1 μm to 3 cm, and more preferably from 50 μm to 2 cm.

[0108] The surface area of ​​each electrode is typically 1 mm. 2 Up to 100mm2 Within a certain range. The surface area of ​​the channel is typically 0.1 mm. 2 Up to 10mm 2 Within the range.

[0109] The organic electrochemical sensor of this invention is particularly suitable for testing blood, saliva, urine, and other biological fluids. In particular, it is suitable for monitoring chronic diseases in blood samples, such as heart failure or kidney failure.

[0110] The present invention also relates to an electronic device comprising an organic chemical sensor according to any of the above embodiments. The electronic device further includes all electrical components necessary for applying voltage to electrodes, measuring current, acquiring and analyzing signals. Advantageously, the electronic device includes a measuring device configured to be reversibly connected to the organic electrochemical sensor. The organic electrochemical sensor may be a disposable sensor, while the measuring device may be reusable. Thus, a reusable measuring device can be used to perform measurements from multiple organic chemical sensors.

[0111] The present invention also relates to a method for detecting analytes in physiological samples. Any embodiment of the above-described organic electrochemical sensor is suitable for this application. The first step of the method is to provide the organic electrochemical sensor of the present invention.

[0112] Physiological samples are placed on at least two channels (C1, C2) and at least three gates (G1, G2, G3) of the sensor.

[0113] For each organic electrochemical transistor including the enzyme recognition layer or the affinity recognition layer: - Apply voltage to the unfunctionalized channel C1 (i.e., apply voltage to the drain electrode D1 while the source electrode S1 is grounded) and the unfunctionalized gate G1, respectively; and apply voltage to the unfunctionalized channel C1 and the functionalized gates (G2, G3). Then measure the output, which is the change in the drain current value.

[0114] The order in which the voltage is applied is irrelevant: the voltage can be applied first to the unfunctionalized channel C1 and the unfunctionalized gate G1. Alternatively, the voltage can be applied first to the unfunctionalized channel C1 and the functionalized gates (G2, G3).

[0115] For each organic electrochemical transistor including the selective ion recognition layer: - Apply voltages to the unfunctionalized channel C1 and the unfunctionalized gate G1, respectively; and apply voltages to the functionalized channel C2 and the unfunctionalized gate G1. -Then measure the output, which is the difference in current values ​​between the two drain electrodes (D1, D2).

[0116] Similarly, the order in which the voltage is applied is irrelevant: the voltage can be applied first to the unfunctionalized channel C1 and the unfunctionalized gate G1. Alternatively, the voltage can be applied first to the functionalized channel C2 and the unfunctionalized gate G1.

[0117] Furthermore, the measurement order of OECT, which includes an enzyme recognition layer, an affinity recognition layer, or a selective ion recognition layer, is irrelevant.

[0118] The voltage applied to the channel and gate is preferably applied continuously for a short period of time. The continuous voltages do not necessarily have the same voltage value: their strength is appropriate for each OECT.

[0119] As mentioned above, the change in drain current is a function of the analyte concentration detected by the recognition layer.

[0120] The sensor is preferably a single-use sensor, so that different OECT detections of each detectable analyte can be performed once using the sensor.

[0121] The method for detecting analytes may also include a calibration phase prior to sample placement, which includes the following steps: - Calibration samples are placed on the at least two channels (C1, C2) and the at least three gates (G1, G2, G3), the calibration samples comprising analytes of known concentrations detectable by at least one of the organic electrochemical transistors of the organic electrochemical sensor. -For each organic electrochemical transistor that includes the enzyme recognition layer or the affinity recognition layer: i. Apply voltage to the unfunctionalized channel C1 (i.e., apply voltage to the drain electrode D1 while the source electrode S1 is grounded), the unfunctionalized gate G1, and apply voltage to the unfunctionalized channel C1 and the functionalized gates (G2, G3). ii. Then measure the first output, which is the change in the drain current value.

[0122] -For each organic electrochemical transistor including the selective ion recognition layer: i. Apply voltage to the unfunctionalized channel C1 and the unfunctionalized gate G1, respectively; and apply voltage to the functionalized channel C2 and the unfunctionalized gate G1, respectively. ii. Then measure the second output, which is the difference in current values ​​between the two drain electrodes (D1, D2).

[0123] - Increase the known concentration of the analyte in the sample. - Repeat steps i. to ii. at least once for each organic electrochemical transistor. - Determine the correspondence between each of the first and second outputs and the concentration of the analyte, and - Optionally, the organic electrochemical sensor may be flushed to make it usable for detecting the analyte in an unknown sample.

[0124] As previously mentioned, the voltages applied to the channel and gate are preferably applied continuously over a short period of time. These continuous voltages do not necessarily have the same value: their intensity is appropriate for each OECT.

[0125] Example The present invention is further illustrated by the following embodiments.

[0126] In these embodiments, such as Figure 4 As shown, the organic electrochemical sensor according to this disclosure includes: -A channel C2, connecting the source S2 and the drain D2, includes a selective potassium ion recognition layer; -A channel C3, connecting the source S3 and the drain D3, includes a selective sodium ion recognition layer; - A gate G2, including an enzyme recognition layer configured to detect creatinine; - A gate G3, including an affinity recognition layer configured to detect NT-pro-BNP; - A gate G4, including an enzyme recognition layer configured to detect creatine; - Two unfunctionalized channels (C1, C4) connect the source S1 and drain D1 and the source S4 and drain D4, respectively; - Two unfunctionalized gates (G1, G4). - The four sources (S1, S2, S3, S4) are arranged as two pairs of shared sources; and - Four drains (D1, D2, D3, D4), including a common drain pair.

[0127] The sensor is fabricated as follows.

[0128] A flexible PET polymer substrate is provided as a support. The substrate is cleaned to remove any dust and organic contaminants that may be present on its surface.

[0129] The electrical layer is a sandwich-shaped Ag-C or Ag-Ag / AgCl layer.

[0130] For the channels (C1, C2, C3, C4), the first intermediate conductive layer is a layer based on poly(3,4-ethylenedioxythiophene)polystyrene sulfonate (PEDOT:PSS). This PEDOT:PSS base layer is covered by a sodium polystyrene sulfonate (PSS:Na) layer, which is used for the channels (C2, C3) containing the ion recognition layer. This first intermediate conductive layer is formed into a solid layer through thermal crosslinking.

[0131] An ion-selective recognition layer is then deposited on the intermediate conductive layer between the two channels (C2, C3). The solvent from the ion-selective recognition layer is then evaporated to form a solid film.

[0132] The enzyme recognition layer is a mixture of four enzymes (creatine anhydride hydrolase CI, creatine amide hydrolase CII, sarcosine oxidase SOx, and HRP + ferrocene mediator) deposited non-covalently on the surface of the gate G2 electrode. The electrode is then dried at room temperature for 30 minutes or in an oven (37°C) for 15 minutes. Finally, a thin layer of sulfonated tetrafluoroethylene fluoropolymer-copolymer (Nafion™) is deposited and dried to block interference between the sample and the electrode. For the additional reference electrode, only a mixture of three enzymes (CI, SOx, and HRP) is deposited.

[0133] In this embodiment, the affinity recognition layer of the sensor uses a biotinylated capture antibody. Therefore, a layer of chitosan is deposited on the surface of gate G3 and dried at room temperature. Prior to deposition of the A / G proteins, the amine functional groups of the chitosan membrane are activated with glutaraldehyde for one hour. The electrode is then washed with PBS to remove unpaired proteins from the chitosan membrane. Antibodies are then deposited to specifically bind the A / G proteins. The role of the A / G proteins is to orient the antibodies to maximize antigen recognition. The antibodies are incubated for one hour, then washed and BSA is deposited to block the surface and prevent nonspecific signals. For the reference electrode, only BSA is deposited.

[0134] Example 1: Calibration of an organic electrochemical sensor A voltage of -0.5V is applied to the unfunctionalized channel C1 (i.e., the voltage is applied to the drain electrode D1 while the source electrode S1 is grounded), and a voltage of 0.15V is applied to the unfunctionalized gate G1.

[0135] Samples containing controlled concentrations of NT-pro-BNP, creatinine, potassium ions, and sodium ions are brought into contact with the channels (C1, C2, C3, C4) and the gates (G1, G2, G3, G4).

[0136] Measure these four concentrations in sequence.

[0137] The concentrations of creatinine and NT-pro-BNP were measured by applying a voltage of -0.5V to the unfunctionalized channel C1 and a voltage of -0.15V to the functionalized gates (G2, G3). The drain current I was also measured. D,2 The change in creatinine concentration was used to derive a value representing the creatinine concentration, and the drain current value I was measured. D,3 The changes in these values ​​are used to derive values ​​representing NT-pro-BNP concentrations.

[0138] The concentrations of potassium and sodium ions were measured by applying a voltage of -0.5V to the functionalized channels C2 and C3 and a voltage of -0.15V to the unfunctionalized gate G1. The current between the drain and source of each channel was measured. In other words, the current difference I between drain D2 and drain D1 was measured. D,1-2 This allows us to derive a value representing the potassium ion concentration and measure the current difference I between drain electrodes D3 and D1. D,1-3 This leads to the value representing the sodium ion concentration.

[0139] The samples were then replaced with those having higher concentrations of NT-pro-BNP, creatinine, potassium, and sodium ions. NT-pro-BNP concentrations varied from 0 pg / mL to 2500 pg / mL. Creatinine concentrations varied from 5 μM to 3000 μM. Potassium concentrations varied from 2.0 mEq / L to 5.96 mEq / L. Sodium concentrations varied from 100 mEq / L to 180 mEq / L. The spiked experiments, characterized by progressively increasing concentrations, are indicated by arrows or columns in the graph.

[0140] Figures 7A to 7D The figures show the current I as the analyte concentration gradually increases. D,2 I D,3 I D,1-2 and I D,1-3 Changes over time.

[0141] Example 2: Measurement of concentration in whole blood samples using an organic electrochemical sensor Using the aforementioned organic electrochemical sensor, the concentrations of NT-pro-BNP, creatinine, potassium ions, and sodium ions were measured in untreated whole blood samples.

[0142] To this end, the blood sample is brought into contact with the channels (C1, C2, C3, C4) and the gates (G1, G2, G3, G4).

[0143] The concentrations of creatinine and NT-pro-BNP were measured by applying a voltage of -0.5V to the unfunctionalized channel C1 and a voltage of -0.15V to the functionalized gates (G2, G3). The drain current value I was measured. D,2The change was -0.0143A, from which the concentration of creatinine was determined to be 100μM, and the leakage current value I was measured. D,3 The change was -0.0155A, from which the concentration of NT-pro-BNP was found to be 300 pg / ml.

[0144] The concentrations of potassium and sodium ions were measured by applying a voltage of -0.5V to the functionalized channels C2 and C3 and a voltage of -0.15V to the unfunctionalized gate G1. The current difference I between the drains D2 and D1 was measured. D,1-2 The current difference was 0.81 mA, from which the potassium ion concentration was calculated to be 4.6 mEq / L. The current difference between drain electrodes D3 and D1 was also measured. D,1-3 The value is -10.1 mA, from which the concentration of sodium ions is calculated to be 140 mEq / L.

Claims

1. An organic electrochemical sensor for detecting analytes in physiological samples, the sensor comprising: - At least two sets of electrodes, each set including a source electrode (S1, S2, S3, S4) and a drain electrode (D1, D2, D3, D4), wherein the source electrode and the drain electrode of each set are connected through organic conductive channels (C1, C2, C3, C4), and - At least two gates (G1, G2, G3, G4). In this configuration, at least one channel (C1) and at least one gate (G1) are unfunctionalized. Wherein, the channels (C2, C3, C4) and the gates (G2, G3, G4) satisfy at least two of the following three conditions: i. At least one channel (C2, C3, C4) or at least one gate (G2, G3, G4) includes a selective ion recognition layer, preferably a selective alkali metal ion recognition layer; ii. At least one channel (C2, C3, C4) or at least one gate (G2, G3, G4) includes a Faraday recognition layer that allows the identification of a specific analyte by reducing or oxidizing it at the channel or the gate; iii. At least one channel (C2, C3, C4) or at least one gate (G2, G3, G4) includes an affinity recognition layer that allows the recognition of biomolecules in the sample by affinity; In this process, any set of electrodes, the channels (C1, C2, C3, C4) connected to the set of electrodes, and any gate (G1, G2, G3, G4) constitute an organic electrochemical transistor.

2. The organic electrochemical sensor according to claim 1, wherein, For each organic electrochemical transistor, at least one of the channels or the gate constituting the organic electrochemical transistor is unfunctionalized.

3. The organic electrochemical sensor according to claim 1 or 2, wherein, The channels (C2, C3, C4) and gates (G2, G3, G4) satisfy the three conditions.

4. The organic electrochemical sensor according to any one of claims 1 to 3, wherein, The Faraday recognition layer is an enzyme recognition layer that allows the recognition of a specific analyte by an enzymatic reaction with the sample. Preferably, the enzyme recognition layer is configured to detect creatinine.

5. The organic electrochemical sensor according to claim 4, wherein, The enzyme recognition layer is configured to perform multi-enzyme cascade reactions.

6. The organic electrochemical sensor according to claim 5 further includes a gate (G4) configured to directly detect creatine.

7. The organic electrochemical sensor according to claim 5, wherein, The enzyme recognition layer for detecting creatine includes creatine anhydrase or creatine anhydride amide hydrolase.

8. The organic electrochemical sensor according to any one of claims 1 to 7, wherein, The affinity recognition layer includes at least one of the following: a specific antibody, an aptamer, a nanobody, or a molecularly imprinted polymer. Preferably, the affinity recognition layer is configured to detect NT-pro-BNP.

9. The organic electrochemical sensor according to any one of claims 1 to 8, wherein, At least one channel (C2, C3, C4) includes an ion recognition layer configured to detect potassium ions.

10. The organic electrochemical sensor according to any one of claims 1 to 9, wherein, At least one channel (C2, C3, C4) includes an ion recognition layer configured to detect sodium ions.

11. The organic electrochemical sensor according to any one of claims 1 to 10, wherein, At least two drain electrodes (D2, D3) are shared, such that each of the shared drain electrodes is used by two of the organic electrochemical transistors, and / or at least two source electrodes (S1, S2, S3, S4) are shared, such that each of the shared source electrodes is used by two of the organic electrochemical transistors.

12. The organic electrochemical sensor according to claim 4, comprising: A channel (C2) comprising a selective potassium ion recognition layer; A gate (G2) comprising an enzyme recognition layer configured to detect creatinine; A non-functionalized channel (C1); An unfunctionalized gate (G1).

13. The organic electrochemical sensor according to any one of claims 1 to 12, wherein, The sensor includes at least three gates (G1, G2, G3, G4).

14. The organic electrochemical sensor according to claim 13, comprising: A channel (C2) comprising a selective potassium ion recognition layer; A gate (G2) comprising a Faraday recognition layer, wherein, The Faraday recognition layer is an enzyme recognition layer that allows the recognition of a specific analyte by an enzymatic reaction with the sample, and the enzyme recognition layer is configured to detect creatinine; A gate (G3) comprising an affinity recognition layer configured to detect NT-pro-BNP; A non-functionalized channel (C1); An unfunctionalized gate (G1).

15. The organic electrochemical sensor according to claim 13, comprising: A channel (C2) comprising a selective potassium ion recognition layer; A channel (C3) comprising a selective sodium ion recognition layer; A gate (G2) comprising a Faraday recognition layer, wherein, The Faraday recognition layer is an enzyme recognition layer that allows the recognition of a specific analyte by an enzymatic reaction with the sample, and the enzyme recognition layer is configured to detect creatinine; A gate (G3) comprising an affinity recognition layer configured to detect NT-pro-BNP; A non-functionalized channel (C1); and An unfunctionalized gate (G1).

16. The organic electrochemical sensor according to claim 15, wherein, The source electrodes (S1, S2, S3, S4) include a pair of common source electrodes, each of which is used by two of the organic electrochemical transistors; wherein the drain electrodes (D1, D2, D3, D4) include a pair of common drain electrodes, each of which is used by two of the organic electrochemical transistors.

17. The organic electrochemical sensor of claim 15 or 16 further comprises a gate comprising an enzyme recognition layer configured to directly detect creatine.

18. The organic electrochemical sensor according to any one of claims 1 to 17, wherein, The organic electrochemical sensor is less than 5 cm in size. 2 .

19. An electronic device comprising the organic electrochemical sensor according to any one of claims 1 to 18.

20. A method for detecting an analyte in a physiological sample, the method comprising the following steps: a. Providing an organic electrochemical sensor according to any one of claims 1 to 18, b. The sample is disposed on the at least two channels (C1, C2, C3, C4) and the at least three gates (G1, G2, G3, G4). c. For each organic electrochemical transistor including the enzyme recognition layer or the affinity recognition layer: i. Apply voltages to the unfunctionalized channel (C1) and the unfunctionalized gate (G1), respectively; and apply voltages to the unfunctionalized channel (C1) and the functionalized gate (G2, G3, G4). ii. Then measure the output, which is the change in the drain electrode current value. d. For each organic electrochemical transistor including the selective ion recognition layer: i. Apply voltage to the unfunctionalized channel (C1), the unfunctionalized gate (G1), and the functionalized channel (C2, C3, C4) and the unfunctionalized gate (G1), respectively. ii. Then measure the output, which is the difference in current values ​​between the two drain electrodes (D1, D2, D3, D4).

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