Patent Description:
The technical field relates to organic electrochemical transistor (OECT)-based sensors. More specifically, the technical field relates to OECT-based sensors that achieve high-transconductance that use a circular and wavy channel geometry. The sensor design may be used for multiplexed OECT-based biosensors for high-throughput biomarker screening applications.

This invention was made with government support under Grant Number GM126571, awarded by the National Institutes of Health. The government has certain rights in the invention.

Biomarkers are molecules that can be detected as indicators of disease processes and body responses to therapeutic intervention. Enzyme-linked immunosorbent assay (ELISA) is currently the gold standard to measure biomarkers. ELISA uses a highly-specific antibody to capture the antigen (capturing antibody). Then the captured antigen needs to be further linked to a secondary detecting antibody for quantification. The multiple operation steps make it challenge to be widely used for point-of-care (POC) applications. A growing trend now is to develop multiplexed immunosensors for high-throughput biomarker screening in one POC device. Such a device will provide more comprehensive information to the caregivers for better disease management. Such POC screening devices are could be used for early detection and monitoring of cancer and other immune-related diseases such as cancer immunotherapy and acute respiratory distress syndrome.

Electrochemical sensors are considered competitive candidates for POC applications. To further improve the sensitivities and lower the limits of detection, electrochemical sensors must be assisted with amplification strategies, often based on enzymatic labeling on the detecting antibody. OECTs have been proved to be efficient strategy to amplify the electrochemical signal. OECT owns the highest transconductance (amplification ability) due to its unique mixed ion/electron conduction property and bulk-modulation ability of the channel. It is now clear that OECT is a superior technique for developing highly-sensitive electrochemical sensors. For example, OECT demonstrates superior signal-to-noise ratio in detecting brain signals. See <NPL>). Besides, OECT can be easily mass-produced with all solution-processable and printable techniques on cost-effective plastics. These advantages make OECT competitive for commercial applications. Currently, examples have been demonstrated with OECT for enzymatic sensors and immunosensors. However, multiplexed OECT immunosensors for high-throughput detection of biomarkers with capturing antibodies or aptamers have not been demonstrated. Such a multiplexed OECT-based immunosensing platform will revolutionize the current POC immunosensors markets regarding sensitivity, detection limit, and cost-effectiveness. <NPL>" is relevant prior art.

In one embodiment, an organic electrochemical transistor (OECT)-based biosensor device is disclosed. The OECT-based sensor comprises a substrate, a gate electrode disposed on the substrate and having capture agents disposed thereon, the capture agents are specific to an antigen (e.g., biomarker/biomolecule), and source and drain electrodes disposed on the substrate and at least partially surrounding the gate electrode, the source and drain electrodes forming an undulating pattern around the gate electrode and separated from one another by a channel. Optionally blocker or spacer molecules may be located on the gate electrode.

In one embodiment, a multiplexed OECT-based biosensor for high-throughput biomarker screening is disclosed. The sensor uses a circular and wavy (CW) channel structure to obtain high-transconductance OECT (HT-OECT). The design of multiplexed HT-OECT biosensors is accomplished by assembling multiple HT-OECT sensors together with multiple capture agent-modified gate electrodes on the same microfluidic chip (or substrate). Such multiplexed HT-OECT biosensors can be used for high-throughput biomarker screening for a better disease monitoring and treatment. Combined with portable device
characterization systems, the HT-OECT multiplexed biosensor can be used for POC applications, wearable applications on the skin, and minimally-invasive biosensing applications using microneedle and catheter as minimally invasive carriers.

In one embodiment, an organic electrochemical transistor (OECT)-based biosensor device includes a substrate and a gate electrode disposed on the substrate or a separate substrate and having capture agents disposed thereon, the capture agents specific to a biomarker. The device further includes source and drain electrodes disposed on the substrate and at least partially surrounding the gate electrode, the source and drain electrodes forming an undulating pattern around the gate electrode and separated from one another by a channel.

In one embodiment, a multiplexed organic electrochemical transistor (OECT)-based biosensing device includes a substrate and a plurality of organic electrochemical transistor (OECT) sensors disposed on the substrate, wherein the plurality of organic electrochemical transistor (OECT) sensors are specific to different antigens (e.g., biomarkers/biomolecules). Each organic electrochemical transistor (OECT) sensor includes a gate electrode disposed on the substrate (or a separate substrate) and having capture agents disposed thereon, the capture agents specific to the different antigens. Source and drain electrodes are disposed on the substrate and at least partially surrounding the gate electrode, the source and drain electrodes forming an undulating pattern that is generally described as circular and wavy (CW) around the gate electrode and separated from one another by a channel.

In another embodiment, a wearable organic electrochemical transistor (OECT)-based biosensor device includes a flexible or soft substrate comprising one or more organic electrochemical transistor (OECT) sensors disposed on the flexible or soft substrate. Each organic electrochemical transistor (OECT) sensor includes a gate electrode disposed on the flexible or soft substrate (or a separate flexible or soft substrate) and having capture agents disposed thereon, the capture agents specific to antigens (e.g., biomarkers/biomolecules). Source and drain electrodes are disposed on the substrate and at least partially surrounding the gate electrode, the source and drain electrodes forming an undulating pattern (e.g., CW configuration) around the gate electrode and separated from one another by a channel. The wearable (OECT)-based biosensor device may, in some embodiments, include microneedles that penetrate the skin or other tissue and absorb bodily fluid which is then exposed to the one or more sensors. In other embodiments, one or more inlets are provided in which sweat or other bodily fluids can enter and then flow over the one or more sensors.

In another embodiment, a minimally invasive catheter or probe device for detecting biomarkers includes an elongate catheter or shaft having one or more organic electrochemical transistor (OECT)-based biosensor(s) mounted thereon, the biosensor(s) comprising a substrate and a gate electrode disposed on the substrate and having capture agents disposed thereon, the capture agents specific to one or more biomarkers. Source and drain electrodes are disposed on the substrate and at least partially surrounding the gate electrode, the source and drain electrodes forming an undulating pattern (e.g., CW configuration) around the gate electrode and separated from one another by a channel.

To use the organic electrochemical transistor (OECT)-based biosensor device, a fluid (e.g., bodily fluid) is exposed to the sensor(s). The fluid may be input into the device or absorbed (e.g., using minimally invasive microneedles). The current response of the sensor is monitored (e.g., source-drain current (Ids)) to detect the presence of (or concentration) of one or more biomarkers/biomolecules. A Source Measuring Unit (SMU) may be used for this purpose. The sensor may be embodied as a microfluidic chip or flow cell. In other embodiments, the sensor is wearable and may be made flexible or soft. In other embodiments, the sensor is integrated in with a patch containing microneedles, a catheter or a probe device for minimally invasive applications. Example bodily fluids that can be measured include blood and urine.

<FIG> schematically illustrates an organic electrochemical transistor (OECT)-based biosensor device <NUM> according to one embodiment. The OECT-based biosensor device <NUM> includes a substrate <NUM> on which the source electrode <NUM> and drain electrode <NUM> is formed. In addition, in some embodiments, the gate electrode <NUM> of the OECT-based biosensor device <NUM> is also formed on the substrate <NUM>. In other embodiments, the gate electrode <NUM> is formed on a secondary substrate <NUM> which is located adjacent to or in close proximity with the substrate <NUM>. For example, as explained herein, the substrate <NUM> that includes the source electrode <NUM> and the drain electrode <NUM> may form the bottom surface of the OECT-based biosensor device <NUM> while the secondary substrate <NUM> that includes the gate electrode <NUM> may form the top of the OECT-based biosensor device <NUM> (or vice versa). The substrate(s) <NUM>, <NUM> may be rigid or flexible or soft. Examples include glass, plastics, elastomers, hydrogels, and other biocompatible polymers. For soft materials used as part of the substrate <NUM>, <NUM>, it is preferrable to use soft materials with low oxygen permeability such as thermoplastic polyurethane (TPU), which can contribute to a higher doping/de-doping efficiency of the channel <NUM> on the soft substrate <NUM>, <NUM>.

As seen in <FIG>, the source electrode <NUM> is formed on the substrate <NUM> and includes an undulating pattern that meanders back-and-forth in a curved and wavy pattern about the periphery of the gate electrode <NUM> (e.g., in a partial circular or arcuate path around the central gate electrode <NUM>). In <FIG>, the gate electrode <NUM> centrally located within respect to the undulating portion of the source electrode <NUM> and the drain electrode <NUM>. The gate electrode <NUM> is circular shaped as seen in <FIG> and is electrically connected to a contact pad <NUM> via a conductive trace, line, or wire. Similarly, the source electrode <NUM> is electrically connected to a contact pad <NUM> via a conductive trace, line, or wire. The drain electrode <NUM> is electrically connected to a contact pad <NUM> via a conductive trace, line, or wire. The channel <NUM> of the OECT-based biosensor device <NUM> is located between the undulating portion of the source electrode <NUM> and the drain electrode <NUM>. This is best seen in the enlarged view of <FIG>. The channel <NUM> is formed between the source electrode <NUM> and drain electrode <NUM> and may include poly(<NUM>,<NUM>-ethylenedioxythiophene) polystyrene sulfonate (PEDOT:PSS) (<FIG>). Additives such as dodecylbenzenesulphonic acid (DBSA), glycerol, ethylene glycol (EG), a fluorosurfactant (e.g., ZONYL®), (<NUM>-glycidyloxypropyl)trimethoxysilane (GOPS) may be added into PEDOT:PSS to control its properties. Thus, the channel <NUM> also at least partially surrounds the gate electrode <NUM> and has a similar undulating configuration as the source electrode <NUM> and the drain electrode <NUM> and is favorable for uniform doping/de-doping. As seen in <FIG>, by having the source electrode <NUM> and drain electrode <NUM> arranged in this curved and wavy configuration with the channel <NUM> formed between, the OECT-based biosensor device <NUM> has an increased W/L ratio for higher transconductance. The distance between the source electrode <NUM> and drain electrode <NUM> (L) is small while the distance along the width (W) is much larger (i.e., W>>L). This produces the higher W/L ratio for higher transconductance. This is in contrast with the prior art design illustrated in <FIG> which has a low W/L ratio. The distance between the source electrode <NUM> and drain electrode <NUM> (L) is substantially uniform along the length of the source electrode <NUM> and drain electrode <NUM>. Exemplary dimensions (e.g., diameter) of the gate electrode <NUM> range from <NUM> to <NUM>, although a preferred embodiment may have the gate electrode <NUM> with a diameter of several millimeters. The dimensions of the source electrode <NUM> and drain electrode <NUM> depends on the size of the gate electrode <NUM>. An exemplary dimension for the source electrode <NUM> and drain electrode <NUM> includes a length of around <NUM> and a width of around <NUM>. Exemplary dimensions of the channel <NUM> include a length of around <NUM> and a width of around <NUM>.

With reference to <FIG>, the gate electrode <NUM> is modified or functionalized with capture agents <NUM> (e.g., antibodies, aptamers, nucleic acids) for capturing specific antigen(s) <NUM>. The antigens <NUM> may include particular analytes, biomolecules, or biomarkers. In one preferred embodiment, the antigens <NUM> include biomolecules or biomarkers that are immunological biomolecules or biomarkers. As one example, the biomarkers may include interferons or interleukins. Examples include, for instance, the biomarkers IFN-λ, thrombopoietin, IL-<NUM>, IL-<NUM>, and IL-<NUM> which can be used to diagnose the severity of COVID-<NUM> patients. The biosensor device <NUM> thus operates as an immunosensor. Of course, the invention is not so limited and may include any type of antigen <NUM>. The gate electrode <NUM> may, in some embodiments, include a spacer or blocker molecule <NUM> to minimize the non-specific binding of antigen. Examples of the spacer or blocker molecule <NUM> include bovine serum albumin (BSA) and <NUM>-mecapto-<NUM>-hexanol (MCH).

With reference to <FIG>, the biosensor device <NUM> includes external circuitry <NUM> that is configured to apply a gate-source voltage and a drain-source voltage and measure the drain-source current. The source-drain voltage (Vds) is preferably a fixed DC value. The gate-source voltage (Vgs) may be constant, swept, or applied at a certain frequency (either DC or AC). DC is preferred but AC may be used. By monitoring the source-drain current (Ids), one can monitor for the presence (or absence) of antigens <NUM> (e.g., biomolecules or biomarkers). The rate of change of the source-drain current (Ids) may also be monitored to determine the concentration of antigens <NUM>. Voltages may be applied to the biosensor device <NUM> and current response measured using a Source Measure Unit (SMU) such as the B2902A Precision Source/Measure Unit (available Keysight Technologies). For example, a single SMU <NUM> may be used to apply voltages and measure current for multiple sensing elements on a multiplexed biosensor device <NUM> such as that illustrated in <FIG>.

In one embodiment, and with reference to <FIG>, a multiplexed biosensor device <NUM> is provided that incorporates multiple sensing elements 40a, 40b, 40c thereon with different sensing elements 40a, 40b, 40c tailored to different antigens <NUM>. In this embodiment, a substrate <NUM> (e.g., microfluidic chip) is provided that has a plurality of organic electrochemical transistor (OECT) sensing elements 40a, 40b, 40c disposed thereon/therein, wherein the plurality of organic electrochemical transistor (OECT) sensing elements 40a, 40b, 40c are specific to different antigens <NUM>. Each organic electrochemical transistor (OECT) sensing elements 40a, 40b, 40c includes a gate electrode <NUM> disposed on the substrate <NUM> and having capture agents <NUM> disposed thereon, the capture agents <NUM> specific to the different antigens <NUM>. A source electrode <NUM> and drain electrode <NUM> (for each sensing element 40a, 40b, 40c) are disposed on the substrate <NUM> and at least partially surround the gate electrode <NUM> as described herein, the source and drain electrodes <NUM>, <NUM> forming an undulating pattern around the respective gate electrode <NUM> and separated from one another by a channel <NUM>.

A multi-layer structure may be formed as illustrated in <FIG> to create a complete microfluidic chip device <NUM> that acts as a flow cell device in which a sample flows through the microfluidic chip device <NUM>. The microfluidic chip device <NUM> includes an intermediate layer <NUM> that forms or defines devices microfluidic chambers <NUM> and channels <NUM>. The chambers <NUM> surround the gate electrodes <NUM> of the sensing elements 40a, 40b, 40c and channels <NUM> allow for fluid to flow across each of the sensing elements 40a, 40b, 40c to expose the fluid to the respective gate electrodes <NUM>. The microfluidic chip device <NUM> includes a top or capping layer <NUM> that includes an inlet <NUM> and outlet <NUM>. Tubing <NUM> may be coupled to the inlet <NUM> and outlet <NUM> and is used to deliver a fluid sample to the microfluidic chip device <NUM> as well as evacuate fluid from the microfluidic chip device <NUM>. Fluid can be delivered to the microfluidic chip device <NUM> manually or using a pump device. Fluid may be continuously pumped through the microfluidic chip device <NUM> or, alternatively, fluid may be flowed into the microfluidic chip device <NUM> and stopped to expose the fluid to the sensing elements 40a, 40b, 40c for a period of time to allow binding/hybridization and signal generation. In some embodiments, the sensing elements 40a, 40b, 40c may be used repeatedly to detect or measure antigens <NUM> in a sample. For instance, antigens <NUM> that are bound to the capture agents <NUM> (such as antibody Ab of <FIG>) may be liberated back into the surrounding fluid using a chemical agent or the like to strip adhered antigens <NUM> from the gate electrode(s) <NUM>. In some embodiments, the top or capping layer <NUM> serves as the secondary substrate <NUM> and holds the gate electrode(s) <NUM>. In this embodiment, the gate electrode(s) <NUM> is/are located on the secondary substrate <NUM> (e.g., on the side of the secondary substrate <NUM> that faces the source/drain electrodes <NUM>, <NUM>) while the source electrode(s) <NUM> and drain electrode(s) <NUM> are located on the "main" substrate <NUM>. The gate electrode <NUM> for each sensing element 40a, 40b, 40c is positioned adjacent to the respective source electrode(s) <NUM> and drain electrode(s) <NUM> as explained herein.

In another embodiment, and with reference to <FIG>, a wearable organic electrochemical transistor (OECT)-based biosensor device <NUM> is disclosed that can be worn on the skin of the subject (or other tissue). The biosensor device <NUM> includes a flexible substrate <NUM> comprising one or more organic electrochemical transistor (OECT) sensors <NUM> disposed on the flexible substrate <NUM>, the flexible substrate <NUM> having one or more organic electrochemical transistor (OECT) sensors <NUM>. Each organic electrochemical transistor (OECT) sensor <NUM> is similar to the construction used in the biosensor device <NUM> described previously. Namely, each sensor <NUM> includes a gate electrode <NUM> disposed on the flexible substrate <NUM> and having capture agents <NUM> disposed thereon, the capture agents <NUM> specific to a particular antigen <NUM>. Each sensor <NUM> may sense the same antigen <NUM> or, alternatively, different sensors <NUM> sense different antigens <NUM>. Source and drain electrodes <NUM>, <NUM> are disposed on the flexible substrate <NUM> and at least partially surrounding the gate electrode <NUM>, the source and drain electrodes <NUM>, <NUM> forming an undulating pattern around the gate electrode <NUM> and separated from one another by a channel <NUM>.

In this embodiment, the flexible substrate <NUM> is made from a biocompatible polymer such as a hydrogel material that allows for the passage of fluid and chemical species through the material of the flexible substrate <NUM> and in contact with the sensors <NUM>. This may include, for example, materials such as poly (methyl methacrylate) (PMMA), polylactic acid (PLA), poly (lactic-co-glycolic acid) (PLGA), polyglycolic acid (PGA), poly (carbonate), cyclic-olefin copolymer, poly (vinylpyrrolidone) (PVP), poly (vinyl alcohol) (PVA), polystyrene (PS) poly (methyl vinyl ether-co-maleic anhydride), and gelatin methacryloyl (GelMA). The biosensor device <NUM> includes a plurality of microneedles <NUM> that extend generally orthogonally from the flexible substrate <NUM>. The microneedles <NUM> may be conical, pyramidal, or similar shape that terminate at a tip. The microneedles <NUM> may have a length of less than about <NUM> and more preferably several hundred micrometers and are able to penetrate the tissue <NUM> of a subject. Thus, the biosensor device <NUM> may be worn like a patch, bandage or the like as seen in <FIG>. The microneedles <NUM> that penetrate the subject's skin or other tissue <NUM> absorb or transfer bodily fluids (e.g., interstitial fluids) to one or more sensors <NUM> in the biosensor device <NUM>. The microneedles <NUM> thus act as functional inlets to the biosensor device <NUM> and allow flow and antigen(s) to enter the flexible substrate <NUM> and contact the sensors <NUM>. For example, as one example, the wearable biosensor device <NUM> may include sensors <NUM> used to measure glucose levels in the subject via interstitial fluids measurements that are accessed using microneedles <NUM>.

The wearable biosensor device <NUM> may include electrical contacts <NUM> that are used to connect with external circuitry <NUM> to measure the response of the sensors <NUM>. For example, the wearable biosensor device <NUM> may be periodically interrogated with external circuitry <NUM> that can be temporarily connected to the wearable biosensor device <NUM> via the electrical contacts <NUM>. Alternatively, external circuitry <NUM> may be permanently connected to the wearable biosensor device <NUM>.

In another embodiment, a minimally invasive catheter or probe device <NUM> incorporates one or more (OECT)-based sensors <NUM> for detecting antigens <NUM>. The device includes <NUM> an elongate catheter or shaft <NUM> having one or more organic electrochemical transistor (OECT)-based sensor(s) <NUM> mounted thereon, the sensor(s) <NUM> including a substrate <NUM> (which may be a surface of the catheter or shaft <NUM>) and a gate electrode <NUM> disposed on the substrate <NUM> and having capture agents <NUM> disposed thereon, the capture agents <NUM> specific to one or more antigens <NUM>. Source and drain electrodes <NUM>, <NUM> are disposed on the substrate <NUM> and at least partially surround the gate electrode <NUM>, the source and drain electrodes <NUM>, <NUM> forming an undulating pattern around the gate electrode <NUM> and separated from one another by a channel <NUM> as described previously. Electrically traces, lines, or wires <NUM> may extend proximally form the source electrodes <NUM>, drain electrodes <NUM>, and gate electrodes <NUM> through the catheter or shaft <NUM> which are then coupled to external circuitry <NUM> to apply the voltages to the sensor(s) <NUM> and monitor response. The catheter or probe device <NUM> may be used intravascularly in the cardiovascular system. The catheter or probe device <NUM> may also be used to traverse the urethra to detect and/or measure antigens <NUM> in urine. In addition, the catheter or probe device <NUM> could be incorporated into a probe device <NUM> that is used as a brain probe.

The fluid that may come into contact with the biosensor device <NUM>, <NUM>, <NUM>, <NUM> may be a physiological fluid such as blood, urine, interstitial fluid, or the like. The fluid may enter the device <NUM>, <NUM>, <NUM>, <NUM> via one or more inlets <NUM> or microneedles <NUM> in other embodiments. Alternatively, the sensing surface of the sensor device <NUM> may be exposed directly to the biological fluid. The electrical response of the device <NUM>, <NUM>, <NUM>, <NUM> is then monitored or analyzed as the fluid is run through the sensor device <NUM>, <NUM> or otherwise exposed to the device <NUM>, <NUM>. This includes applying a gate-source voltage and a drain-source voltage and measuring the drain-source current with the circuitry <NUM>. The substrate <NUM> may include electrical contacts or pads <NUM>, <NUM>, <NUM> as explained herein to interface with respective gate electrode <NUM>, source electrode <NUM>, and drain electrode <NUM>. For example, leads or wires can connect the external circuitry <NUM> to the sensor device <NUM>, <NUM>, <NUM>, <NUM>. Alternatively, such electronics could be integrated on-chip/on-substrate or adjacent thereto in some embodiments.

The transconductance of an OECT represents its amplification ability. Higher transconductance contributes to higher sensitivity and lower detection limits. The transconductance can be increased by increasing the W/L ratio and thickness of the channel <NUM>. The W/L ratio can be increased by changing the geometry of the device. Interdigitated electrodes (IDE) have been demonstrated to increase the W/L ratio in a limited area. To increase to the W/L, at the same time considering the uniformity modulation of the gate electrode <NUM>, the biosensor device of <FIG> was created. The source electrode <NUM> and drain electrode <NUM> surround the circular gate electrode <NUM>. Compared to conventional IDE, this design is favorable for a more uniform doping/doping of the channel <NUM> at a given gate potential. The undulating or wavy structure of the source electrode <NUM> and drain electrode <NUM> is also favorable to increase the mechanical flexibility and stretchability of the device <NUM>. Compared to the commercial electrochemical sensing electrodes, the W/L ratio can be increased by several orders of magnitudes, which indicates a much higher amplification ability of the biosensor device <NUM>, thus enables a higher sensitivity and a lower detection limit.

First, the gate electrode <NUM>, the source electrode <NUM>, and the drain electrode <NUM> of HT-OECT biosensor device <NUM> were fabricated on a glass substrate <NUM> (but can be plastics and elastomers). Afterward, CW PEDOT:PSS channel <NUM> is patterned between the source electrode <NUM> and drain electrode <NUM> with photolithographic technique. The Au gate electrode <NUM> is then modified with capturing antibodies <NUM> with high affinity for the targeted biomarker (antigen <NUM>, <FIG>). As a demonstration, the use of the HT-OECT biosensor device <NUM> was shown for the detection of codeine, a pain reliever small molecular drug that is sometimes used by immunotherapy patients. The codeine-antibody <NUM> on gate electrode <NUM> was modified. The thiol group in the antibody <NUM> improved the bonding with the Au electrode <NUM>. The gate electrode <NUM> is then further treated with spacers or blockers <NUM> such as <NUM>-mecapto-<NUM>-hexanol (MCH) and bovine serum albumin (BSA) to minimize the non-specific binding of the antigen codeine to its surface.

A PDMS microfluidic chip device <NUM> such as that illustrated in <FIG> includes a microfluidic channel <NUM> is then fabricated to guide the flow to the HT-OECT sensing element (e.g., a flow cell or microfluidic chip is formed). Finally, a top or capping layer <NUM> is used to seal the microfluidic channel <NUM>. The codeine-antibody modified gate electrode <NUM> may also be fabricated on the capping or top layer <NUM>. In this case, the assembly of HT-OECT biosensor device <NUM> is completed by aligning and sandwiching the top or capping layer <NUM> (w/ gate electrode <NUM>) and the substrate <NUM> (w/ source and drain electrodes <NUM>, <NUM>) through the PDMS microfluidic channels <NUM>.

To detect the codeine with HT-OECT biosensor device <NUM>. The gate voltage (Vgs) may be constant, swept, or applied at a certain frequency, while the source-drain voltage (Vds) is fixed constant. <FIG> illustrates a graph showing impedance changes in the HT-OECT biosensor device <NUM> (without a capture agent <NUM>) for the detection of codeine. The gate electrode <NUM> modified with spacer/ blocker <NUM> (but without antibody) shows no impedance change to the concentration change of the antigen codeine. <FIG> illustrates a graph showing impedance changes in the HT-OECT biosensor device <NUM> (with a capture agent <NUM>) for the detection of codeine. The gate electrodes <NUM> modified with both spacer/blocker <NUM> and antibody <NUM> shows clear impedance change to the concentration change of codeine. When there is no antibody-antigen (codeine) binding, the source-drain current (Ids) maintains a stable value (<FIG>). When the antigen <NUM> (i.e., codeine) is present in the fluid, it binds with the antibody <NUM>, which increases the gate/electrolyte capacitance and finally decreases the Ids. The Ids-decrease profile is dependent on the concentration of the antigen codeine. A steeper decrease in Ids with time indicates a higher codeine concentration in the fluid. As shown in <FIG>, increasing the codeine concentration in the fluid causes a steeper decrease in the Ids, demonstrating the potential of the HT-OECT biosensor device <NUM> in detecting the antigen codeine.

The demonstrated HT-OECT biosensor device <NUM> may be scaled up for multiplexed HT-OECT immunosensors. The design is illustrated in <FIG>. In this design, a number of gate electrodes <NUM> are introduced on the same microfluidic chip device <NUM> (e.g., chip) and modified with capturing antibodies <NUM> for the detection of different antigens <NUM>. The number of gate electrodes <NUM> depends on the number of antigens <NUM> required to monitor the disease. For example, mild or severe COVID-<NUM> patients can be diagnosed by measuring the following five biomarker antigens <NUM> IFN-λ, thrombopoietin, IL-<NUM>, IL-<NUM>, and IL-<NUM>. In this case, corresponding five multiplexed HT-OECT-based sensing elements <NUM> can be fabricated on the same substrate <NUM> for high-throughput biomarker screening. The detection of more biomarkers can be realized by increasing the number of OECT sensing elements <NUM> integrated on the microfluidic chip device <NUM>. Because the antibody is highly specific to a certain biomarker or antigen <NUM>, the presence of other antigens <NUM> or biomarkers in the sample will have little influence on other HT-OECT-based sensing elements <NUM>.

The design of multiplexed HT-OECT biosensor devices <NUM> can be used for POC tests, wearable biosensing, and minimally-invasive biosensing by using microneedles <NUM> (<FIG>), catheter and/or probe (<FIG>) as carriers for the sensors <NUM>, <NUM>. In particular, the multiplexed HT-OECT can be developed as ion sensors, enzymatic sensors, electrolyte sensors, metabolites or metal ion sensors on microneedles and catheters, which have not been reported so far.

First, the design of multiplexed HT-OECT biosensor device <NUM> can be used for the POC tests. For example, with reference to the embodiment of <FIG>, the collected bodily fluid can be pumped or input into the inlet <NUM> of the microfluidic chip device <NUM>. The flow is enabled either by external pressure or capillary force. Once antigens <NUM> (e.g., biomarkers) in the flow pass the corresponding antibody-modified gate electrodes <NUM>, they are captured. The binding process increases the gate/electrolyte capacitance. The change in the capacitance leads to the change in the effective Vgs and finally results in a more pronounced change in the Ids of the HT-OECT which is detected or measured by external circuitry <NUM>. Further integration with a portable data readout system allows the simultaneous detection with the multiplexed HT-OECT biosensor device <NUM>.

Second, the multiplexed HT-OECT biosensor device <NUM> can be used for wearable biosensing applications. The multiplexed HT-OECT biosensor devices <NUM> can be first fabricated on ultrathin plastic or elastomer. The wavy structure of the source/drain electrodes <NUM>, <NUM>, used to increase the transconductance, also contributes to a higher stretchability on the elastomer. The flexible and miniature multiplexed HT-OECT biosensor devices <NUM> can be laminated on the skin for direct sweat analysis-the sweat flows into a microchannel through an inlet (which may be a dedicated inlet <NUM> or a microneedle <NUM>) under external pressure and a capillary force (like a POC test on the skin) to flow over the gate electrode <NUM>. Fluid can thus enter the HT-OECT biosensor device <NUM> from the external environment and come into contact with the gate electrode <NUM>. Biomarkers in the sweat are captured by corresponding gate electrodes <NUM>. The detected signal is then amplified by corresponding HT-OECT by causing a more pronounced change in the Ids. The data can be collected by mobile phones via cable or wireless technology such as Bluetooth or the like.

The multiplexed HT-OECT biosensor devices <NUM> or sensors <NUM>, <NUM> sensors described herein can be used for minimally-invasive biosensing applications such as that illustrated in <FIG>. To enable minimally-invasive detection with the multiplexed HT-OECT sensors, they may be integrated with microneedles <NUM> (<FIG>), or catheters/probes <NUM> (<FIG>) (e.g., brain probe). Microneedles <NUM> penetrate the skin to access the interstitial fluid (ISF) and capillary blood, with less or no pain. Bodily fluid in ISF or capillary blood can contact microneedles <NUM> or flow into the lumen of a hollow microneedle <NUM>. When an antigen <NUM> such as a biomarker in the liquid pass the antibody-modified gate electrode <NUM> of the HT-OECT, they are captured, which, in turn, leads to a more pronounced change in the Ids which can be measured and/or detected using external circuitry <NUM>.

To integrated HT-OECT with medical catheters and brain probes, the multiplexed HT-OECT sensors <NUM>, <NUM> can be fabricated on ultrathin plastic substrates <NUM> (such as parylene). The fabricated sensor devices <NUM>, <NUM> can be wrapped on a medical catheter or laminated on a brain probe with biocompatible adhesive glues such as that illustrated in <FIG>. The integration of multiplexed HT-OECT biosensors <NUM> with a medical catheter or probe <NUM> allows precise biomarker screening directly in blood vessels and ureter. Integration of multiplexed HT-OECT biosensors <NUM> with brain probes <NUM> allows directly biomarker-screening in brain cerebrospinal fluid, which is essential for a better understanding and treating brain cancer diseases.

The HT-OECT sensing elements or sensors can be fabricated on rigid, flexible, stretchable, and soft substrates <NUM>. The CW configuration of the channel <NUM> that surrounds the circular gate electrode <NUM> with the source and drain electrodes <NUM>, <NUM>, allows a more uniform modulation of the Vgs to the Ids. With the CW HT-OECT, the sensor design can be used in single and multiplexed sensors on a chip. The multiplexed HT-OECT sensors on a chip can be used for high-throughput biomarker screening applications. They can be used for POC tests at home, wearable tests on the skin (in sweat), and minimally-invasive tests with microneedles, catheters, and brain probes.

The PEDOT:PSS aqueous suspension (Clevios PH1000) was purchased from Heraeus Electronic Materials GmbH (Leverkusen, Germany). <NUM>+% pure glycerol, dodecylbenzenesulfonic acid (DBSA), and glycidoxypropyltrimethoxysilane (GOPS) were bought from Sigma-Aldrich. The gold, silver, titanium (Ti), and palladium were provided by UCLA nano-lab. The photoresist SPR <NUM>, MF 26A developer, and acetone for photolithography and patterning of electrodes <NUM>, <NUM>, <NUM> were provided by Integrated Systems Nanofabrication Cleanroom (ISNC) at California NanoSystems Institution (CNSI). PDMS (SYLGARD™ <NUM> Silicone Elastomer Kit) was purchased from Dow Inc. The antibody <NUM> for codeine <NUM> capturing was purchased from Base Pair Biotechnologies, Inc. Codeine was obtained from Spectrum Chemical Mfg. Corp, under a DEA license and with UCLA EHS Controlled Substances Use Authorization.

The Au electrodes <NUM>, <NUM>, <NUM> were deposited by e-beam evaporation using a CHA Solution Electron Beam Evaporator. The substrates <NUM> were first cleaned by acetone, IPA, and DI water. Then photoresist SPR <NUM> was spin-coated with <NUM> rpm on substrates <NUM> using Headway spinner with PWM32 controller. Photolithography was performed to create patterns for the gate electrode <NUM>, source electrode <NUM>, and drain electrode <NUM>. The exposed part was washed away with MF-26A developer. Then <NUM>-nm-thick Ti and <NUM>-nm-thick Au were deposited. The remaining photoresist was stripped off via acetone. The PDMS microfluidic channel (e.g., microfluidic chamber <NUM> and channel <NUM> as illustrated in intermediate PDMS layer <NUM> in <FIG>) was fabricated in a <NUM>:<NUM> ratio with a laser-ablated PMMA mold (Universal Laser VLS <NUM>). UV exposure was performed to enable bonding between PDMS and substrate <NUM> and top or cover <NUM>.

The sequence of the antibody (aptamer) capture agent <NUM> for the antigen <NUM> codeine capturing is GGG ACA GGG CTA GCA GTA GGA TTG GGT GAG GGG ATG TGC TGT GGA GGC AAA GCT TCC G [SEQ ID NO: <NUM>]. The powdered antibody <NUM> was first used to make an original solution in resuspension buffer (<NUM>). Before using, the powder was de-frozen at room temperature (RT) for <NUM>. The original solution was diluted with a folding buffer (<NUM>). The diluted solution was then heated up <NUM> for <NUM> and cooled down to RT. The, the antibody solution was mixed with the reducing buffer (<NUM>:<NUM>) and incubated for another <NUM>. The reduced antibody solution was then diluted with PBS (w/ <NUM> MgCl<NUM>) to desired concentrations. Before modification, the Au gate electrode <NUM> was first cleaned with H<NUM>SO<NUM> (<NUM>), then cleaned with cyclic voltammetry (-<NUM>, <NUM> V) until the curves stabilize. Afterward, the electrodes <NUM> were rinsed with ethanol and dried with N<NUM> gun. The antibody buffer solution was then incubated on electrodes for <NUM> hrs. The mercaptoethanol (MCH) blocker solution (<NUM>) with blocker <NUM> (MCH) was prepared by diluting the solution in PBS (w/ <NUM> MgCl<NUM>). The blocker solution was then incubated on the Au electrode <NUM> for <NUM> hour. Afterward, the electrodes <NUM> were rinsed with PBS solution and dried with N<NUM> gun.

Claim 1:
A multiplexed organic electrochemical transistor (OECT)-based biosensing device comprising:
a substrate;
one or more organic electrochemical transistor (OECT) sensors disposed on the substrate, wherein the one or more organic electrochemical transistor (OECT) sensors are specific to different biomarkers, each organic electrochemical transistor (OECT) sensor comprising:
a gate electrode disposed on the substrate or a separate substrate and having capture agents disposed thereon, the capture agents specific to one of the different biomarkers; and
source and drain electrodes disposed on the substrate and at least partially surrounding the gate electrode,
characterized in that the source and drain electrodes form an undulating pattern around the gate electrode with a channel formed between them.