Direct functionalized electrochemical transistors and convective driven ultra-fast detection of biomarkers using transistors

Abstract

Devices and methods for analyte detection using AC electrokinetics / electrohydrodynamics forces in combination with an OECT-based immunosensor are disclosed. An analyte binding agent, such as a nanoantibody functionalized organic electrochemical transistor (OECT), is combined with the micro-agitation effect of an alternating current electric heat flow (ACET) for ultra-fast detection of an analyte at a single molecule to nanomole level. The ACET flow caused by the biased AC electric field can rapidly convect the analyte onto a concentric gate electrode within one minute, and the analyte is captured via a recognition unit that binds to the analyte binding agent while sweeping non-specifically bound analyte from the surface.

Description

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS

[0002] This application claims the benefit of and priority to U.S. Provisional Application No. 63 / 280,887 filed on November 18, 2021, U.S. Provisional Application No. 63 / 283,447 filed on November 27, 2021, and U.S. Provisional Application No. 63 / 318,188 filed on March 9, 2022, which are hereby incorporated by reference in their entireties. Technical Field

[0003] The present invention is generally in the field of organic electrochemical transistor (OECT) based immunosensor devices and methods useful for detecting target analytes in a sample. Background Art

[0004] Reliable biomolecular diagnostics are important tools for early detection of disease, especially for preventing outbreaks of infectious diseases such as HIV, Ebola, and the recent widespread coronavirus (SARS-CoV-2), which has killed a million people worldwide at the time of writing. Early stage detection is particularly useful for identifying and isolating asymptomatic infected patients before the disease spreads. Therefore, it is crucial to develop tests that can provide rapid, easy-to-use, and accurate results.

[0005] Test methods such as RRT-PCT have inherent limitations, such as labor-intensive sample preparation in a laboratory setting, which increases test turnaround time and is not suitable for small sample volumes. In addition, sample transportation and complex sample preparation steps before testing may also reduce clinical sensitivity and lead to false-negative results.

[0006] Organic electrochemical transistors (OECTs) have attracted attention as a promising alternative biosensing technology that can surpass other state-of-the-art electrical biosensor methods such as field effect transistors (FETs). OECTs allow biosensing applications in aqueous environments with low voltage operation (<1 V) and high amplification (Lin and Yan, 2012). OECTs detect target analytes by monitoring the capacitance or faradaic changes occurring in the device interface. To date, the most common practical biosensing signal is to provide an enzymatic redox reaction promoted by a specific analyte. Analyte binding to the enzyme induces a redox reaction, which then causes a potential drop across the device interface, which is then amplified as a change in the current flowing from the transistor source terminal to the drain terminal (Bernard et al., 2008; Ohayon et al., 2020). An important advantage of OECT biosensors is that their selectivity can be tuned by incorporating suitable biorecognition elements without the need for fluorescent, radioactive or other labels (Wustoni et al., 2019; Wustoni et al., 2020).

[0007] Diagnostic tools using the enzymatic (mainly paralogous cleavage) activity of CRISPR-Cas proteins are programmable for different targets, some of which have been integrated with portable detection methods such as lateral flow strips. However, they still require time-consuming sample pretreatment steps, including target amplification and / or RNA extraction, to achieve the required attomole sensitivity. Therefore, the time from obtaining the sample to obtaining the final result is much longer than 15min. Although these tests can reduce the cost and complexity of molecular testing, they seem to have very limited advantages compared to, for example, established RT-LAMP or more simplified RT-PCR solutions. There is a demand for RNA diagnostic tools that achieve rapid response and high sensitivity in original untreated samples.

[0008] OECT-based immunosensors have been developed in which p-type accumulation mode materials (p(g0T2-g6T2)) were used as channel materials, exhibiting single protein molecule sensitivity and extended dynamic range. The VHH-72 nanobody-OECT biosensor detected specific proteins from untreated human samples under ambient conditions by manual pipetting for 30 s every 3 min followed by incubation for 10 min. This biosensor involves cumbersome pipetting steps to accelerate the transport of proteins to the sensor, resulting in relatively short incubation times, which is not possible for diffusion-based transport. Therefore, highly automated particle transport methods for rapid screening and field-use applications are crucial.

[0009] It is an object of the present invention to provide OECT-based biosensors which achieve reduced incubation times and / or increased sensitivity.

[0010] Another object of the present invention is to provide a method for fabricating OECT-based immunosensors for use in diagnostic assays.

[0011] Another object of the present invention is to provide improved methods for detecting analytes in samples using OECT-based biosensors. Summary of the invention

[0012] Provided herein are OECT-based biosensors and methods of manufacturing and using the same. The biosensor comprises an OECT and a biorecognition layer. The OECT is preferably an electrolyte gate-controlled transistor, i.e., a three-terminal electronic device comprising a source electrode, a drain electrode, a channel, and a gate electrode. In some preferred embodiments, the biorecognition layer is integrated on the gate electrode of the OECT.

[0013] In some forms, the biorecognition layer comprises a bio-self-assembled monolayer (SAM) (herein Bio-SAM) of biomolecules formed by a specific bio-autocatalytic coupling strategy, as disclosed herein, and preferably excludes the use of a layer of SAM formed using organic molecules, i.e., Chem-SAM, which can be formed by self-assembly of organic molecules (such as thiol-containing organic molecules disclosed in PCT / IB2021 / 055981). Bio-SAM includes biorecognition elements, proteins that are not complete antibodies, fusion proteins that do not include complete antibodies, or CRISPR proteins (e.g., ribonucleoprotein complex Cas:gRNA). The biorecognition element includes a binding partner for a target analyte, preferably a pathogen or a component thereof. In a preferred embodiment, the biorecognition element can be bound to the SARS-CoV-2 receptor binding domain (RBD) or spike protein (S1), and the biorecognition element is preferably a nanobody (e.g., VHH72) or a related antibody fragment, or a CRISPR protein (such as a ribonucleoprotein complex Cas:gRNA).

[0014] The organization of the biorecognition elements on the OECT surface can be represented by the following general formula:

[0015] L1-AP1:AP2-L2-B

[0016] Formula I

[0017] Wherein L1 is the first linker, AP1 is the first peptide binding partner possessing a specific biological recognition sequence; AP2 is the second peptide binding partner that recognizes this sequence on AP1; AP1 and AP2 are binding partners, preferably, covalent binding partners; however, they can be members of an affinity pair; L2 is the second linker, and B is the biological recognition element of the biosensor. AP2-L2-B can be produced as a single fusion protein without chemical modification. The OECT surface is modified with L1-AP1 via L1. Preferably, L1 is a peptide sequence comprising a cysteine ​​residue at its N or C terminus, which provides a free SH group for directly coupling L1 to the surface of the gate electrode. Preferably, the OECT surface is not modified with a simple organic alkanethiols or their derivatives such as 1,6-hexanedithiol (HDT) between the OECT and L1, i.e., L1 is directly coupled to the OECT (referred to herein as directly functionalized OECT) and provides a specific (bioderived) recognition sequence AP1 for AP2, rather than a non-specific reactive group. In other words, the electrode having the biorecognition layer formed thereon does not contain free organic molecules, i.e., organic molecules that self-assemble on the electrode surface but are not attached to L1-AP1 and therefore not attached to the biorecognition element. For example, the gate electrode modified with the biorecognition layer does not contain free cysteine ​​that is not connected to the first linker L1 (such as a peptide sequence). The binding or coupling of AP1 and AP2 produces a biological self-assembled monolayer, herein Bio-SAM. Therefore, the biorecognition layer includes the linker L1 and the Bio-SAM. In a specific embodiment, AP1 is a spyTag and AP2 is a spyCatcher.

[0018] The biorecognition layer is preferably integrated on the gate electrode of the OECT.

[0019] The channel of the OECT can be formed directly from a conductive polymer, or by incorporating a conductive polymer on the surface of a conductive substrate (such as by spin coating a conductive polymer on the surface of a metal electrode). Any suitable conductive polymer can be used for the channel of the OECT. In some embodiments, the conductive polymer used for the channel of the OECT is a p-type polymer. In some embodiments, the conductive polymer used for the channel of the OECT is an n-type polymer. In some embodiments, the conductive polymer used for the channel of the OECT is poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonate) (PEDOT:PSS) or any other mixed (ionic and electronic) semiconductor. In some embodiments, the conductive polymer used for the channel of the OECT is p(g0T2-g6T2). In some embodiments, the conductive polymer used for the channel of the OECT is p(g3C2T2-T). In some embodiments, the conductive polymer used for the channel of the OECT is p(C6NDI-T). Depending on the channel material, the OECT can be operated in depletion mode or accumulation mode.

[0020] Also disclosed is a method for making an OECT device containing a binding partner for any target analyte (e.g., an antigen from any pathogen). The method comprises functionalizing a gate electrode of an OECT with a biorecognition layer containing a binding partner for the target analyte as follows: (i) contacting at least a portion of the surface of the gate electrode with a first solution containing a peptide sequence and a first peptide binding partner, the peptide sequence preferably containing a cysteine ​​residue at its N- or C-terminus, and (ii) contacting the first binding partner-modified surface with one or more solutions containing a linker or linker element and a binding partner for the analyte. The method for functionalizing a gate electrode of an OECT excludes the step of contacting at least a portion of the surface of the gate electrode with a solution containing a thiol-containing organic molecule (such as cysteine) prior to step (i).

[0021] Referring to Formula I (L1-AP1:AP2-L2-B), at least a portion of the surface of the gate electrode is contacted with a solution containing: a peptide sequence (L1), which preferably contains a free SH group most conveniently provided by a cysteine ​​residue at its N or C terminus; and a first peptide binding partner (AP1) for forming a first binding partner layer on the surface of the electrode, and this first binding partner layer is contacted with a composition containing AP2-L2-B under conditions that result in conjugation of AP1 and AP2 to form a bio-self-assembled layer, and is referred to herein as Bio-SAM. In a preferred embodiment, AP1 and AP2 are spyTag and spyCatcher, respectively. The autocatalytic covalent binding between the spyTag / spyCatcher pair orients the analyte binding partner on the surface of the OECT device. In a particularly preferred embodiment, the method comprises contacting the OECT surface with a first solution containing a spyTag (one partner in the spyTag / spyCatcher pair), followed by contacting with a second solution containing a spyCatcher (the second partner in the spyTag / spyCatcher pair, which in turn is linked to a binding partner for the target analyte). In a preferred embodiment, the target analyte is SARS-CoV-2 RBD and its binding partner is a specific nanobody.

[0022] Also disclosed are methods for identifying target analytes using directly functionalized OECTs. One embodiment includes contacting a sample with a directly functionalized OECT-based biosensor disclosed herein that includes a binding partner for the target analyte. Another embodiment includes using AC electrokinetic / electrohydrodynamic forces, as described herein.

[0023] Devices and methods for analyte detection using AC electrokinetic / electrohydrodynamic forces in combination with OECT-based immunosensors are disclosed. Generally, the gate electrode of an OECT-based immunosensor is surrounded by a conductive element that is used to apply the AC electrokinetic / electrohydrodynamic forces in the device.

[0024] One embodiment provides a directly functionalized OECT as described above, while another embodiment provides an indirectly nanobody functionalized OECT (such as those described in PCT / IB2021 / 055981), either embodiment is combined with a conductive element that can provide a micro-stirring effect of alternating current electrothermal flow (ACET) for ultra-fast detection of pathogens such as severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) in complex body fluids (such as saliva, plasma, serum and blood) at single molecules to nanomolar levels. The ACET flow induced by the biased AC electric field applied to the conductive element can rapidly convect the analyte to the concentric gate electrodes within one minute, and capture antigens from the pathogen such as SARS-CoV-2 spike protein (S1) or receptor binding protein domain (RBD) via the recognition unit (nanobody), while sweeping non-specific antigens from the surface. BRIEF DESCRIPTION OF THE DRAWINGS

[0025] Figure 1A A schematic representation of a SpyDirect nanobody functionalized OECT sensor is shown. The gate electrode is directly biofunctionalized with a spyTag-linked cysteine ​​peptide and a spyCatcher-linked nanobody. BSA is used as a blocking agent. The sensor is incubated with saliva. Figure 1B A schematic diagram demonstrating the operation of an OECT sensor is shown. Either a bare sensor gate or a hatched sensor gate is mounted on top of the channel for signal acquisition. Figure 1C The output signals of the OECT sensor before and after exposure to COVID-19 positive saliva are shown. NR = |(g mD -g m0 )| / g m0

[0026] Figure 2A and 2B The SpyCatcher-nanobody biofunctionalized QCMD of the gate electrode is shown ( Figure 2A ) and AFM( Figure 2B ) characterization.

[0027] Figures 3A to 3G The electrochemical characterization and XPS spectrum of the biofunctionalized Au gate electrode are shown: Cyclic voltammogram ( Figure 3A ), Bode plot (solid line and dash-dot line correspond to the magnitude and phase of impedance, respectively) ( Figure 3B), Nyquist plots of gold electrodes before and after subsequent functionalization with SpyTag-cysteine ​​peptide and nanobody-SpyCatcher ( Figure 3C ; The inset is the equivalent circuit model used to fit the impedance spectrum), the calculated charge transfer resistance (Rct) and double layer capacitance (Cdl) changes of the electrode ( Figure 3D ; measurements were performed in 10 mM [Fe(CN)6]3- / 4- in 10 mM PBS at pH 7.4), high-resolution XPS spectra of Au 4f, C 1s and N1s during each modification step ( Figures 3E to 3G ). Nanobody-SpyCatcher buffer contains BSA.

[0028] Figures 4A to 4E Shown is the detection of SARS-CoV-2 using the SpyDirect SARS-CoV-1 nanobody sensor. Figure 4A The transfer properties of accumulation-mode OECTs after sequential incubation of SARS-CoV-2 spike at different concentrations are shown. Figure 4B The transconductance of an OECT sensor with a nanobody-functionalized gate electrode after incubation is shown as a function of gate voltage. Figure 4C The normalized response (NR) to SARS-CoV-2 spike and non-target (lysozyme) in saliva is shown. Error bars represent the standard deviation calculated from at least 3 gate electrodes. Figure 4D The schematic diagram of the OECT interface is shown. The ionic circuit used to simulate the OECT consists of capacitors C corresponding to the gate and channel, respectively. G and C CH and the resistor R corresponding to the electrolyte E composition. Figure 4E The compaction mechanism is shown. The gate capacitance decreases after protein binding (C G →C′ G,eff ) schematic diagram.

[0029] Figures 5A to 5E Shown is random testing of the SARS-CoV-2 spike protein in saliva and untreated wastewater. Figure 5A Shown is the random detection of SARS-COV-2 spike in saliva using the SpyDirect SARS-CoV-1 Nanobody Gate while using the GFP Nanobody Gate as a control. Figure 5B Detection of SARS-CoV-2 spike in untreated wastewater is shown. Figure 5C and 5D shows a HDT SAM based gate ( Figure 5C ) and SpyDirect Gate ( Figure 5D) Long-term stability test against SARS-CoV-2S1 (1.2 pM) in saliva. The gates were stored in 10 mM PBS, pH 7.4 at 4 °C for 3, 7, and 14 days before use. Figure 5E The sensor response of SpyDirect gates prepared from dried cysteine ​​peptides is shown. These cysteine ​​peptide-modified gates were stored under ambient conditions for 7 days before nanobody immobilization.

[0030] Figures 6A to 6F Detection of SARS-CoV-2 pseudotyped lentivirus in human saliva, universal transport medium (UTM), and untreated wastewater is shown. Fig. 6A Shown are the background responses of the SpyDirect sensor and the HDT SAM sensor to human saliva, UTM, and untreated wastewater. Figure 6B Shown are the responses of the SpyDirect sensor to SARS-CoV-2 pseudotyped lentivirus and non-targets in human saliva. Figures 6C to 6F The use of SpyDirect-based nanobody gates for UTM ( Figure 6C ) or wastewater ( Fig. 6E Sensor responses of SARS-CoV-2 pseudotyped lentivirus in ) and UTM ( using HDT SAM-based nanobody gates Fig.6D ) or wastewater ( Fig. 6F ) Sensor response of SARS-CoV-2 pseudotyped lentivirus in . GFP nanobody gate was used as control. Error bars represent standard deviation calculated from at least 3 gate electrodes.

[0031] Figures 7A to 7H Electrochemical and chemical stability testing of the SARS-CoV-1 nanobody gate is shown. Fig. 7A and 7B shows a SpyDirect-based gate ( Fig. 7A ) and HDT SAM-based gates ( Figure 7B )’s charge transfer resistance (Rct). Figures 7C to 7H shows a SpyDirect-based gate ( Figures 7C to 7E ) and HDT SAM-based gates ( Figures 7F to 7H) XPS spectra of C1s, O1s, and N1s before and after 7 days of storage. SARS-CoV-1 nanobody functionalized gates were stored in 10 mM PBS, pH 7.4 at 4 °C and data were acquired at 0, 3, and 7 days (n=3). OECT planar gates with well-defined areas (500 μm × 500 μm) were functionalized with SARS-CoV-1 nanobodies using the SpyDirect method or HDTSAM as a linker, and these SARS-CoV-1 nanobody gates were then used to monitor their long-term electrochemical stability. 10 mM [Fe(CN)6] 3- / 4- Impedance measurements were performed using the Randles circuit model to fit the impedance spectrum.

[0032] Figures 8A to 8B The stability of the gate functionalization is shown. The charge transfer resistance of the gate coated with Nanobodies was measured after storage in PBS for 0, 3 and 7 days. ( Fig. 8A )Gates fabricated using spyDirect(Cys-peptide) SAM exhibit stable low charge transfer resistance over time. ( Figure 8B ) Gates fabricated using HDT-SAM exhibit a dramatic increase in charge transfer resistance over time, indicating surface degradation due to, for example, oxidation processes.

[0033] Fig. 8A A circular gate electrode is shown surrounded by a gold layer to apply alternating current heat flow (ACET). Scale bar: 100 μm. Figure 8B is a schematic diagram showing the architecture of the biorecognition unit on the gate electrode. The SpyTag peptide is chemically coupled to a HDT monolayer assembled on a gold electrode. The Nanobody-SpyCatcher fusion protein is attached to this chemical layer by autocatalytic formation of a covalent SpyCatcher-SpyTag bond. A complex of three Ty1 Nanobody structures with the trimeric SARS-CoV-2 spike protein is shown (partially shown in gray, based on PDB entry 6ZXN). Figure 8C Schematic diagram showing the operating steps of the ACET-enhanced nanobody-OECT biosensor. 10 μL of sample solution is introduced on top of the concentric gate electrode. The function generator provides AC potential and induces ACET to concentrate the target molecule (e.g., spike protein) and the immobilized nanobody on the surface. This process takes 2 min. The electrode is then rinsed with phosphate buffered saline (PBS), flipped and mounted on top of the channel for signal acquisition.

[0034] Figures 9A to 9F Characterization of Ty1-nanobody functionalized concentric gate electrodes is shown. Fig.9AShown are Nyquist plots of the Au electrode before and after subsequent functionalization with HDT, SpyTag peptide, and SpyCatcher / Nanobody. Fig. 9B The Randles equivalent circuit model for extracting solution resistance (Rsol), double electric layer capacitance (CEDL), charge transfer resistance (Rct), and Weber impedance (W) is shown. Fig. 9C Calculated values ​​of the electrical components after each functionalization step are shown. Fig.9D Real-time monitoring of Ty1-nanobody immobilization on Au surface using QCM-D is shown (top: frequency change, bottom: mass of each layer on the crystal). A two-step functionalization scheme consisting of SpyTag peptide coupling followed by SpyCatcher / nanobody immobilization was performed on a gold QCM-D sensor comprising an HDT layer. PBS was used to remove unbound species from the surface. Fig.9E and 9F shows the high-resolution C1s XPS spectra of the gold surface after immobilization of different functional layers ( Fig.9E ) and N1s XPS spectrum ( Fig.9F ).

[0035] Figures 10A to 10I The biosensor performance of p-type and n-type OECTs is shown. Figures 10A to 10D The use of Ty1-nanobody electrode ( Fig. 10A and 10C ) and GFP-nanoantibody electrodes ( Fig. 10B and 10D ) recorded p(g3C2T2-T) channel ( Fig. 10A and 10B ) and p(C6NDI-T)( Fig. 10C and 10D )Channel transfer curve. Under ACET, the gate electrode was incubated with solutions containing different concentrations of SARS-CoV-2S1. ( Fig. 10A )and( Fig. 10C ) indicates that the protein concentration of the solution increased from 1×10 -18 m increases to 1×10 -9 m. Figures 10E to 10G The normalized responses (NR) of p(g3C2T2-T)OECTs determined at VG = -0.1 V and VD = -0.1 V are shown, respectively. Fig.10E ) and the normalized responses of p(C6NDI-T) OECTs determined at VG = 0.5 V and VG = 0.9 V (VD = 0.1 V) ( Fig.10F and 10G ). Error bars represent standard deviations calculated from at least three gate electrodes. Fig. 10H and10I Shown are the NRs reported at VG = 0.5 V and VG = 0.9 V (VD = 0.1 V), respectively, for a p(C6NDI-T) OECT operating with the same gate electrode but without ACET. Fig.10J The voltage drop at the gate / electrolyte interface (V G 压降 ) is a schematic diagram of the equivalent circuit model. Figure 10K Representative Bode plots are shown for an Au gate electrode and the channel material biased to its doped state at V = |0.5| V vs. Ag / AgCl. Fig.10L Shows V G 压降 The gate capacitance (Cg) decreases with increasing concentration of target molecules bound on its surface.Analytical models are considered for four types of devices based on different gate to channel capacitance ratios: 1) Cg=0.1Cch, 2) Cg=Cch, 3) Cg=10Cch and 4) Cg=100Cch.

[0036] Figures 11A to 11G The performance of ACET-assisted n-type OECT sensors in saliva and a benchmark analysis of their performance are shown. Fig.11A and 11B The responses of the p(C6NDI-T) channel gated by the Ty1-nanobody functionalized electrode and the GFP-nanobody functionalized electrode to different concentrations of SARS-CoV-2S1 in saliva are shown, which are respectively at VG = 0.5V ( Fig.11A ) and VG=0.9V( Fig. 11B Saliva was diluted with lysis buffer at a ratio of 1:4. Fig. 11C and 11D The response of n-type OECTs to randomly selected saliva samples containing different amounts of SARS-CoV-2S1 is shown. Each data point represents the response of a single gate electrode. The data in these two figures were acquired using two sets of electrodes prepared simultaneously. The small circles represent negative control measurements made with GFP-nanobody functionalized gate electrodes, while the large circles represent records from Ty1-nanobody functionalized gate electrodes. Data recorded at VG = 0.9 V and VD = 0.1 V are presented. Fig.11E The NR difference between Ty1-nanobody-gated OECTs and GFP-nanobody-gated OECTs is shown. Fig.11F and 11G The power consumption versus operating voltage of the OECT-based immunosensor (indicated by stars) is shown in Figure 3 compared to other transistor-based immunosensors (see Table 3 for a list). Fig.11F ) as well as the limit of detection (LOD), dynamic range, and incubation time ( Fig.11G).

[0037] Fig. 12A and 12B An exemplary reusable OECT for measuring a signal is shown. Fig. 12A ) and an exemplary reader device ( Fig. 12B ) schematic diagram.

[0038] Fig.13A and 13B An exemplary point-of-care ("POC") OECT is shown. Fig.13A ) and exemplary operating steps using POC OECT ( Fig. 13B ) schematic diagram.

[0039] Fig.14 An exemplary high-throughput device is shown, where 96 gate electrodes are used for simultaneous measurements. The gate holder is transferred between different plates containing, for example, patient samples, wash buffers, electrolytes, with shaking to accelerate the (otherwise severely diffusion-limited) target capture by the nanobody surface for measurement. The holder can be placed on top of a custom measurement plate containing the OECT channel, where electrical contacts can be established with all gates from the top. DETAILED DESCRIPTION

[0040] Devices for sensitive and rapid analyte detection using OECT sensors are provided. The devices can be used to rapidly detect infections caused by pathogens such as SARS-CoV-2 using directly biofunctionalized OECT sensors (where the OECT sensors are functionalized to include binding partners for the target antigen) by detecting the presence of antigens (analytes) specific for the pathogen.

[0041] The disclosed directly functionalized OECT sensor is based on the unexpected discovery that an improved and simplified immobilization strategy using a fully protein-based Bio-SAM that excludes the underlying Chem-SAM (which is made of organic molecules such as alkanethiols and their derivatives) improves analyte detection using OECT sensors. Using SpyTag and SpyCatcher as examples herein, a cysteine-terminated peptide with SpyTag (and without any non-biochemical modification) is directly coupled to an OECT sensor surface (such as the gate electrode surface of an OECT), and is then used to immobilize anti-SARS-CoV nanobodies on the OECT sensor without involving any organic chemicals, such as HDT, that would self-assemble into a monolayer (Chem-SAM) at the OECT surface and then attach to the peptide with SpyTag. Using a direct biofunctionalization approach, the electrode surface containing a binding partner for the target analyte formed thereon does not contain free organic molecules, i.e., organic molecules that self-assemble on the electrode surface but are not attached to the peptide with SpyTag.

[0042] Compared to the Chem-SAM-based nanobody OECT sensor, the cysteine ​​peptide-linked nanobody-OECT sensor achieves a higher packing density of the coupled nanobody and thus a higher binding capacity (such as at least 1.5 times higher or at least 2 times higher) for the target analyte. For example, the packing density of the peptide is as high as 82×10 12 SpyDirect peptides / cm 2 (In contrast, using Chem SAM, 32×10 12 Maleimide peptide / cm 2 ), and the packing density of nanobodies is as high as 7.8×10 12 Pieces / cm 2 (In contrast, using Chem SAM, 6.5 × 10 12 Pieces / cm 2 They exhibited reduced background noise levels, lower limits of detection (LODs), and wide detection windows (e.g., LODs as low as 6×10 -22M, with a detection window of the order of 10), and are insensitive to contaminants such as dithiothreitol (DTT). They also exhibit improved long-term stability. For example, the mean Rct value of the SpyDirect gate showed negligible change after 3 days of storage in PBS, while showing a 12.5% ​​decrease after 7 days of storage. In contrast, the HDTSAM-based SARS-CoV gate showed a dramatic increase in mean Rct value after storage (2600% after 3 days of storage and 6400% after 7 days of storage, respectively). In addition, manufacturing becomes easier and less expensive for three reasons: (1) the three-step immobilization is reduced to two steps, (2) the remaining two steps are performed in aqueous solution, thus avoiding the use of organic solvents, and (3) the peptide used for immobilization is now composed only of natural amino acids without any chemical modification (although versions of such peptides with non-natural linkers or modifications can be easily imagined). The sensor using cys-peptide coupling was again tested at 6×10 -22 The nominal LOD of M selectively detects SARS-CoV-2 spike protein (in contrast to sensors using Chem-SAM, which exhibit a detection limit of 2.8*10 -16 M or 1.8x 10 -20 The sensor fabricated using this strategy was also validated using SARS-CoV-2 pseudotyped lentivirus spiked into untreated human saliva and reliably detected approximately 20 copies of the virus particles in 5 μL of sample solution.

[0043] The devices can be used with minute volumes of solution (as little as 5 μl) from readily available samples, including saliva (and also blood from a finger prick); they are readily available and affordable, allowing large-scale application; they are easy to use by minimally trained users or even patients themselves (see, e.g., respectively, Figures 13A to 13B and Fig.14 A to 14B); can be operated using an inexpensive, portable and handheld readout unit (e.g., connected to a smartphone); can be designed to be detected by a biological detection unit (see, e.g., Fig.14 The modular exchange of high-throughput devices (shown in Figure 1) uses the same platform to detect different targets. The device can have multiple channels in the area below the gate electrode, meaning that one sensing gate electrode can be used to obtain multiple readouts simultaneously (see, for example, Fig. 12A These channels can be made of different organic materials.

[0044] Conventional biosensors rely on passive diffusion-dominated transport of target species on the sensor surface. The sample containing the analyte to be detected is incubated with the sensor for a certain period of time to interact with the binding sites on the sensor surface. This incubation time (sometimes up to several hours) is an important step that limits the speed of the sensor.

[0045] Provided is an ACET-enhanced OECT-based immunosensor for rapid and reliable detection of biomarkers, SARS-CoV-2 is used as an example herein. ACET can be induced by applying an AC potential to a conductive layer placed close to a gate electrode, which causes a target molecule (such as a spike protein) to accumulate on an immobilized binding partner (such as a nanobody) for the target molecule. The AC potential can be in the range of 1Vpp to 8Vpp or 2Vpp to 6Vpp, such as about 6Vpp (pp refers to peak to peak). ACET-enhanced OECT-based immunosensors can use directly functionalized OECTs or indirectly functionalized OECTs. The disclosed device is based at least on the following findings (through numerical simulations and experimental studies): ACET-induced mixing can significantly reduce the time for immune complex formation (≤2min from sample to result) and can achieve higher specificity and lower background compared to the same OECT-based immunosensor without ACET, due to the electrothermal flow-induced removal of nonspecific species from the sensor surface. ACET-enhanced OECT-based immunosensors also exhibit low energy consumption (i.e., nW levels, such as about 100nW). For example, the ACET-enhanced OECT-based immunosensor can be operated using only 100 nW of power and about 2 min of incubation with 10 μL or 5 μL of sample to detect target analytes in complex media. Two types of (semi)conductive polymers in the channel are illustrated: p-type (p(g0T2-g6T2)) and n-type (p(C6NDI-T)). In some embodiments, n-type OECTs can outperform p-type OECTs in terms of sensitivity and lower power consumption. The developed sensor is largely reusable, easy to manufacture, and highly modular. Its performance was validated using clinical untreated saliva samples from patients with COVID-19 and demonstrated comparable sensitivity to RT-PCR methods.

[0046] I. Definitions

[0047] As used herein, "affinity interaction" refers to the combination of non-covalent interactions between a ligand and its binding partner that form a complex.

[0048] As used herein, an "affinity tag" is a peptide sequence that is appended to a protein so that the protein can be purified from a crude biological source using affinity techniques.

[0049] "Covalent bond" refers to a bond or organic moiety that covalently links a molecule (eg, a fusion protein) to a non-cellular surface.

[0050] As used herein, the term "polypeptide" refers to amino acid chains of any length, regardless of modification (eg, phosphorylation or glycosylation).

[0051] As used herein, a "vector" is a replicon, such as a plasmid, phage or cosmid, into which another DNA segment can be inserted so as to bring about the replication of the inserted segment. The vector described herein may be an expression vector.

[0052] As used herein, an "expression vector" is a vector that contains one or more expression control sequences.

[0053] As used herein, an "expression control sequence" is a DNA sequence that controls and regulates the transcription and / or translation of another DNA sequence.

[0054] As used herein, "operably linked" means incorporated into a genetic construct so that the expression control sequence effectively controls the expression of the coding sequence of interest.

[0055] As used herein, "conservative" amino acid substitutions are those in which the substituted amino acid has similar structural or chemical properties.

[0056] As used herein, "non-conservative" amino acid substitutions are those in which the charge, hydrophobicity, or bulk of the substituted amino acid is significantly altered.

[0057] As used herein, the term "host cell" refers to prokaryotic and eukaryotic cells into which a recombinant expression vector can be introduced.

[0058] As used herein, "transformed" and "transfected" encompass the introduction of a nucleic acid (eg, a vector) into a cell by a variety of techniques known in the art.

[0059] II. Compositions and Devices

[0060] A. Direct functionalization of OECTs

[0061] The disclosed devices include directly functionalized OECTs engineered to include a biorecognition layer, which includes a biorecognition element. The biorecognition element is a component that can specifically interact with its cognate target. The OECT is an electrolyte-gated transistor, i.e., a three-terminal electronic device including a source electrode, a drain electrode, a channel, and a gate electrode. In some preferred embodiments, the biorecognition layer is integrated on the gate electrode of the OECT.

[0062] 1. Biometric layer

[0063] The biorecognition layer is designed to provide a stable complex, preferably an immune complex, between the biorecognition element in the biorecognition layer and the target analyte / antigen. The biorecognition element is preferably not a multidomain antibody. The biorecognition layer is preferably integrated on the surface of the gate electrode of the OECT.

[0064] The biorecognition layer comprises a self-assembled monolayer (herein Bio-SAM) formed by a specific bioautocatalytic coupling strategy, as disclosed herein. Bio-SAM comprises a biorecognition element, such as a nanobody, a fusion protein or a Cas protein, preferably a nanobody.

[0065] The organization of the biorecognition elements on the OECT surface can be represented by the following general formula:

[0066] L1-AP1:AP2-L2-B

[0067] Formula I

[0068] Wherein L1 is the first linker, AP1 is the first peptide binding partner having a specific biological recognition sequence; AP2 is the second peptide partner that recognizes this sequence, that is, AP1 and AP2 are binding partners, preferably covalent binding partners; however, they can be members of an affinity pair; L2 is the second linker, and B is a biological recognition element for a biosensor. AP2-L2-B can be produced as a single fusion protein without chemical modification. The OECT surface is modified with AP1 via L1, producing a layer of OECTs modified with the first peptide binding partner, and providing a specific (bioderived) recognition sequence AP1 for AP2, while recognizing non-specific reactive groups. The combination of AP1 and AP2 produces a biological self-assembled monolayer, herein Bio-SAM. Therefore, the biological recognition layer includes L1 and Bio-SAM.

[0069] L1 and L2 are preferably peptide linker sequences of at least 2 amino acids in length. Preferably, the peptide or polypeptide domain is a flexible peptide or polypeptide. "Flexible linker" herein refers to a peptide or polypeptide containing two or more amino acid residues bound by one or more peptide bonds, which provide an increased degree of rotational freedom for the two polypeptides connected thereto compared to what the two connected polypeptides would have in the absence of a flexible linker. Exemplary flexible peptides / polypeptides include, but are not limited to, the amino acid sequences Gly-Ser, Gly-Ser-Gly-Ser (SEQ ID NO: 9), Ala-Ser, Gly-Gly-Gly-Ser (SEQ ID NO: 10), (Gly4-Ser) 3 (SEQ ID NO: 11) and (Gly4-Ser) 4 (SEQ ID NO: 12), GSGSGSGS (SEQ ID NO: 13) and SGSG (SEQ ID NO: 14). Additional flexible peptide / polypeptide sequences are well known in the art. In a preferred embodiment, L1 is a flexible peptide as disclosed herein, which is modified to include a cysteine ​​residue at its N or C terminus, for example CGGSGSGSG (SEQ ID NO: 22) or GSGC (SEQ ID NO: 23) are preferred sequences for L1. The SpyTag peptide already contains a disordered sequence at its C terminus. If the spyTag sequence is added to the N terminus of peptide 23 (GSGC), the entire flexible linker region can be about the same length as in the case of SEQ 22.

[0070] Nanobodies can be naturally derived from immunization of animals or synthetically derived (ie, sybodies) or a combination thereof. Sybodies are disclosed in, for example, Walter et al., doi: https: / / doi.org / 10.1101 / 2020.04.16.045419. In some preferred embodiments, the organic molecules forming the SAM include thiols. In some preferred embodiments, the linker is a biomolecular linker formed by bioconjugation of two peptides such as SpyTag / SpyCatcher bioconjugation (Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012)).

[0071] In a particularly preferred embodiment, the biorecognition layer does not include a SAM formed by an alkanethiol or a derivative thereof (e.g., 1,6-hexanedithiol) between the OECT and L1, i.e., L1 is directly coupled to the OECT (i.e., directly functionalized OECT). The electrode surface having the biorecognition layer formed thereon does not contain free organic molecules, i.e., organic molecules that self-assemble on the electrode surface but are not attached to L1-AP1 and therefore not attached to the biorecognition element. For example, the gate electrode modified with the biorecognition layer does not contain free cysteines that are not connected to L1-AP1, such as peptide sequences with SpyTag. In addition, the biorecognition layer preferably includes SpyTag / SpyCatcher bioconjugation as a linker, which enables controllable orientation of nanobody functionalization, maximizing the small sensor area (0.64 mm 2 ) and enhance the sensitivity of the OECT sensor. 2 At least 50-120×10 12 SpyTag peptides (such as at least 50, at least 60, at least 70, at least 80, at least and up to 90, at least 95, or about 97×10 12 SpyTag peptides) and per cm 2 At least 10×10 12 Nanobody-SpyCatcher molecules (such as at least 10, at least 12, at least 14 and at most 15×10 12 Nanobody-SpyCatcher molecules) are coupled to each cm of the surface of the gate electrode of the OECT. 2 superior.

[0072] i. OECTs modified with the first binding partner

[0073] The first peptide binding partner is typically coupled to the gate electrode of the OECT by reacting the first binding partner connected to the linker with the surface of the gate electrode of the OECT under conditions that result in the coupling of the linker to the surface of the OECT, followed by bioconjugation of the first peptide binding partner to the second peptide binding partner by contacting the OECT modified with the first building partner with the second binding partner required to form a SAM (i.e., Bio-SAM). The peptide-modified OECTs as disclosed herein do not include free organic molecules, i.e., organic molecules that self-assemble on the electrode surface but are not attached to the first linker and the first peptide binding partner (L1-AP1). For example, the gate electrode of the OECT modified with a peptide does not contain free cysteine ​​that is not attached to the first peptide binding partner.

[0074] Exemplary thiols that are excluded from coupling to the surface of the gate electrode of the OECT include, but are not limited to, alkylmonothiols (e.g., methylthiol, ethylthiol, 2-propylthiol, butylthiol, pentylthiol, tert-butylthiol, 1-hexylthiol, 1-octylthiol, 1-nonylthiol, 1-decylthiol, 1-undecanethiol, 1-dodecylthiol, 1-tridecylthiol, 1-tetradecylthiol, 1-pentadecanethiol, 1-hexadecanethiol, 1-octadecanethiol, 1-nonadecanethiol, 1-decylthiol, 1-undecylthiol, 1-dodecylthiol, 1-tridecylthiol, 1-tetradecylthiol, 1-pentadecanethiol, 1-hexadecanethiol, 1-octadecanethiol, 1-nonadecanethiol and 1-eicosanethiol, etc.), alkanedithiols (e.g., 1,2-ethanedithiol, 1,4-butanedithiol, 1,3-butanedithiol, 1,5-pentanethiol, 1,6-hexanedithiol (HDT), 1,7-octanedithiol, 1,8-nonanedithiol, 1,3-propanedithiol (PDT), etc.), 3-mercaptopropionic acid (MPA), thiophenol, dimercaptosuccinic acid, glutathione, 2-mercaptoethanol, lipoic acid, and 1,4-benzenedimethanethiol. Additional exemplary thiols that can form a SAM on the surface of the gate electrode of an OECT and be excluded from the gate electrode are disclosed in Love et al., Chem. Rev., 105: 1103-1169 (2005).

[0075] In a particularly preferred embodiment, the first peptide binding partner is coupled to the gate electrode surface of the OECT via a linker, preferably a peptide linker comprising a cysteine ​​residue at preferably the C- or N-terminus of the peptide linker.

[0076] Reference Formula I:

[0077] L1-AP1:AP2-L2-B

[0078] Formula I

[0079] The first peptide binding partner modified surface is formed by L1-AP1, wherein L1 is the first linker and is preferably coupled to the surface of the OECT gate electrode via a terminal cysteine, AP1 is the first peptide binding partner; AP2 is the second peptide binding partner; AP1 and AP2 are binding partners, preferably, covalent binding partners; however, they can be members of an affinity pair; L2 is the second linker, and B is a biological recognition element, such as a nanobody, a fusion protein or a Cas protein.

[0080] ii. Bio-SAM

[0081] The biorecognition layer integrated on the gate electrode of the OECT includes a SAM, which is a biological self-assembled monolayer formed due to the covalent binding of two binding partners or the affinity interaction of the affinity pair, thereby forming a Bio-SAM, which includes a biorecognition element, such as a protein that is not a complete antibody or a fragment thereof, a fusion protein that does not contain a complete antibody or a fragment thereof, or a CRISPR protein (e.g., a ribonucleoprotein complex Cas: gRNA). In a preferred embodiment, the biorecognition element is a fusion protein as described herein. In another preferred embodiment, the biorecognition element is a CRISPR protein as described herein, such as a ribonucleoprotein complex Cas: gRNA.

[0082] Reference Formula I:

[0083] L1-AP1:AP2-L2-B

[0084] Formula I

[0085] Bio-SAM is formed by the interaction of a first peptide binding partner AP1 with a second peptide binding partner AP2, thereby forming a SAM, ie Bio-SAM, which includes a biorecognition element B exposed to interact with its binding partner in a sample contacted therewith.

[0086] In particularly preferred embodiments, AP1 and AP2 comprise a spyTag / spyCatcher pair, wherein in some preferred embodiments, AP1 is preferably a spyTag, such as AHIVMVDAYKPTK (SEQ ID NO: 6), i.e., spyTag, and in preferred embodiments, the spyTag consists of: SEQ ID NO: 6; VPTIVMVDAYKRYK (SEQ ID NO: 7), i.e., spytag002 (Keeble et al., Angew. Chem. Int. Ed. Engl. 2017, 56, 16521-16525) or RG VPHIVMVDAYK RYK (SEQ ID NO: 8), i.e., spyTag 0003 (Keeble et al.: Proc Natl Acad Sci U SA. 2019 Dec 26; 116(52): 26523-26533).

[0087] The preferred SpyTag / SpyCatcher bioconjugation system is disclosed in Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012). It is based on a modified domain from a surface protein of Streptococcus pyogenes (SpyCatcher) that recognizes a homologous 13 amino acid peptide (SpyTag). After recognition, the two form a covalent isopeptide bond between the side chain of lysine in SpyCatcher and the side chain of aspartic acid in SpyTag. SpyTag / SpyCatcher bioconjugation is a robust method for conjugating recombinant proteins, in which the peptide SpyTag can spontaneously react with the protein SpyCatcher in a facile manner and with high specificity (Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012); Keeble et al., Angew. Chem. Int. Ed. Engl. 2017, 56, 16521-16525. The peptide SpyTag Tag (13 amino acids) spontaneously reacts with the protein SpyCatcher (12.3 kDa) to form an intermolecular isopeptide bond between the pair. The SpyTag peptide is preferably not chemically functionalized with a maleimide functional group, and the SpyCatcher is linked to a biorecognition element (e.g., a nanobody) such that upon bioconjugation between the SpyTag and the SpyCatcher, the biorecognition element is integrated on the gate electrode surface in a consistent orientation. In an optional embodiment, the SpyTag may include a linker.

[0088] Other binding partners can be used as AP1:AP2, for example, snoopTag peptide:SnoopCatcher (Veggiani et al., PNAS 2016 113 (5) 1202-1207), MoonTag:MoonCatcher (homolog of SpyCatcher); snapTag labeling system; sortase reaction to connect two shorter peptides; coating the surface with biotinylated peptides or biotin-modified hydrocarbons and then coupling streptavidin nanobody fusion protein can also be used. SnoopTag is a peptide covalently bound to the SnoopCatcher protein (KLGDIEFIKVNK) (SEQ ID NO: 18). The second generation SnoopTagJr has been developed for binding to SnoopCatcher or DogTag (mediated by SnoopLigase) (KLGSIEFIKVNK) (SEQ ID NO: 19). DogTag is a peptide covalently bound to SnoopTagJr (DIPATYEFTDGKHYITNEPIPPK) (SEQ ID NO: 20) mediated by SnoopLigase; SdyTag is a peptide covalently bound to the SdyCatcher protein (DPIVMIDNDKPIT) (SEQ ID NO: 21). SdyTag / SdyCatcher has kinetically dependent cross-reactivity with SpyTag / SpyCatcher. These systems are known in the art (reviewed in Hatlem et al., Int J Mol Sci. 2019 May; 20(9): 2129). The NAP tag protein labeling system and the CLIP tag protein labeling system enable specific covalent attachment of almost any molecule to a target protein. There are two steps to use this system: clone and express the target protein as Fusions (such as nanobodies- Fusions) and labeling of the fusions with a SNAP tag substrate of choice. The SNAP tag is a small protein based on human O6-alkylguanine-DNA-alkyltransferase (hAGT), a DNA repair protein.

[0089] SNAP-tag substrates are dyes, fluorophores, biotin, or beads conjugated to guanine or chloropyrimidine leaving groups via a benzyl linker. In the labeling reaction, the substituted benzyl group of the substrate is covalently attached to the SNAP tag. TM It is a modified version of the SNAP tag that is engineered to react with benzylcytosine rather than benzylguanine derivatives. When used in conjunction with the SNAP tag, the CLIP tag enables orthogonal and complementary labeling of two proteins simultaneously in the same cell.

[0090] a. Non-antibody protein biorecognition elements

[0091] Preferred biorecognition elements (B) for incorporation into the disclosed device are not complete antibodies, but more compact recognition domains, such as nanobodies or sybodies. As used herein, non-antibody proteins refer to proteins that are not complete antibodies, which are multi-domain proteins. In contrast, nanobodies, also referred to as single-domain antibodies, are antibody fragments consisting of a single monomer variable antibody domain. Like full antibodies, they can selectively bind to specific antigens. In the case of molecular weight of only 12 to 15 kDa, single-domain antibodies are much smaller than ordinary antibodies (150 to 160 kDa) consisting of two heavy protein chains and two light chains, and are also smaller than Fab fragments (about 50 kDa, one light chain and half a heavy chain) or single-chain variable fragments (scFv, about 25 kDa fusions of two variable domains, one variable domain from a light chain, and the other from a heavy chain).

[0092] Camels naturally produce a class of antibodies consisting only of heavy chains, in which the target recognition module consists of a single variable domain (VHH or Nb). (De Meyer et al., 2014; Kubala et al., 2010). Nanobodies offer unique properties, simpler structural conformations and high stability under a range of different conditions (De Meyer et al., 2014).

[0093] In a preferred embodiment, the device comprises a nanobody that binds to an antigen from SAR-CoV-2 such as a spike protein. Specific nanobodies are available for the SARS-CoV-2 spike protein (S protein) or more specifically for the receptor binding domain (RBD) within the S1 subunit of the protein (Wrapp et al., Science 2020, 367(6483), 1260-1263). Sybodies are disclosed in, for example, Walter et al., doi: https: / / doi.org / 10.1101 / 2020.04.16.045419. Examples include SARS1 / 2 nanobodies (VHH72), as shown below:

[0094] QVQLQESGGGLVQAGGSLRLSCAASGRTFSEYAMGWFRQAPGKEREFVATISWSG GSTYYTDSVKGRFTISRDNAKNTVYLQMNSLKPDDTAVYYCAAAGLGTVVSEWDYD YDYWGQGTQVTVSSGS(SEQ ID NO:17)

[0095] b. Fusion protein

[0096] In some embodiments, the biorecognition element is expressed as a fusion protein containing a first polypeptide domain, a linker domain, and a purification tag. Biorecognition elements (such as nanobodies, sybodies, etc.) can be easily expressed in various forms by fusing with other proteins, peptides, or effector domains, thereby adjusting their uses. Utilizing this feature, the disclosed biorecognition element is expressed as a fusion protein that contains a linker and preferably a protein purification tag; and can be represented by the following general formula II:

[0097] AP1:AP2-L2-BC-PT

[0098] Formula II,

[0099] wherein AP1:AP2 and L2 are as defined above for Formula I, and PT is a purification tag (an amino acid sequence that allows the polypeptide to be captured on an affinity matrix) and C is an optional cleavage site, e.g., a 3C cleavage site. In one embodiment, the biorecognition element B is first oriented in the fusion protein sequence (i.e., the N-terminus). However, the biorecognition element B can be located at the C-terminus of the fusion protein. For example, a tag such as c-myc, hemagglutinin, polyhistidine, or Flag TM (Kodak) can be used to aid in peptide purification. Such tags can be inserted anywhere within a polypeptide, including the carboxyl or amino termini. Polyhistidine affinity tags, also known as His tags or His6, are usually composed of six consecutive histidine residues, but can vary in length from two to ten histidine residues; glutathione S-transferase (GST); maltose binding protein (MBP), calmodulin binding peptide (CBP); intein-chitin binding domain (intein-CBD), streptavidin tags, etc. are other known tags.

[0100] A preferred fusion protein is

[0101] MTGQVQLQESGGGLVQAGGSLRLSCAASGRTFSEYAMGWFRQAPGKEREFVATISWS

[0102] GGSTYYTDSVKGRFTISRDNAKNTVYLQMNSLKPDDTAVYYCAAAGLGTVVSEWDY

[0103] DYDYWGQGTQVTVSSGSGSGSGSGSVDTLSGLSSEQGQSGDMTIEEDSATHIKFSKRD

[0104] EDGKELAGATMELRDSSGKTISTWISDGQVKDFYLYPGKYTFVETAAPDGYEVATAITFTVNEQGQVTVNGKATKGDAHISGLEVLFQGPTGHHHHHHHH (SEQ ID NO: 15), wherein the underlined residues constitute the linker (8 amino acids long) between the SAR1 / 2-nanobody and the SpyCatcher-3C cleavage site-His8. His8 is shown in bold, and the HRV (human rhinovirus) 3C protease cleavage site is located within the underlined sequence. The HRV 3C protease is a recombinant 3C protease derived from human rhinovirus type 14 expressed in E. coli. The enzyme has the same activity as the native protein and cleaves a specific amino acid sequence (LEVLFQ↓GP) (SEQ ID NO: 16).

[0105] A schematic diagram of the SpyDirect nanobody-functionalized OECT sensor is shown in Figure 1A (Direct biofunctionalization of gate electrodes using spyTag-linked cysteine ​​peptides and spyCatcher-fused nanobodies). BSA was used as a blocking agent. The sensor was incubated with a saliva buffer mixture. ( Figure 1B ) A bare sensor gate or an incubated sensor gate is mounted on top of the channel for signal acquisition. As illustrated herein, the synthetic SpyTag peptide is directly coupled to the gold surface via a linker without the need for a 1,6-hexanedithiol (HDT) monolayer to form a Chem-SAM, as disclosed in, for example, PCT / IB2021 / 055981. The Nanobody-SpyCatcher fusion protein then attaches itself to the spyTag by autocatalytic formation of a covalent SpyCatcher-SpyTag bond, forming a Bio-SAM. The Nanobody domains define sensor specificity and are interchangeable (the examples below illustrate device design and functionality for GFP, SARS-CoV, and MERS-CoV).

[0106] In some embodiments, the peptide binding partner is prepared by chemical synthesis. The peptide binding partner can be chemically synthesized by methods known in the art such as by using solid phase peptide synthesis, solution phase synthesis, chemical connection (see, for example, Chandrudu et al., Molecules, 18 (4): 4373-4388 (2013)). In some embodiments, SpyTag can be chemically synthesized. For example, SpyTag comprising a joint and optionally comprising a cleavage site and / or a purification tag is chemically synthesized by a known method as described above, and can be represented by the following general formula III.

[0107] Alternatively, the peptide binding partner can be expressed as a fusion protein containing a linker domain and a purification tag. Peptide binding partners (such as SpyTag) can be easily expressed in various forms by fusion with other proteins, peptides or effector domains to adjust their use. In these embodiments, the disclosed SpyTag can be expressed as a fusion protein comprising a linker and preferably a protein purification tag; and can be represented by the following general formula III:

[0108] L1-AP1-C-PT

[0109] wherein AP1 and L1 are as defined above for Formula I, and C and PT are as defined above for Formula II. AP1 is preferably C-terminally cysteine-linked.

[0110] c. CRISPR protein

[0111] Clustered regularly interspaced short palindromic repeats (CRISPR) and CRISPR-associated (Cas) protein systems have been used for gene editing, and nuclease dead mutants (dCas9) have been used for gene regulation across different species. Many categories and types of Cas effectors have been reported in the literature. For example, Cas effectors that selectively bind or cleave DNA, RNA, or both can be selected. In addition, many (but not all) Cas proteins have two cleavage activities, one in cis (for target molecules), and the other in trans (also known as the paralogous activity of cleaving "bystander" non-specific nucleic acids after a specific target has been identified). Due to its programmability (i.e., custom design of guide RNA spacer sequences) and multiple features (i.e., by using Cas orthologs targeting different nucleic acid molecules or cleaving different sequences), Cas-based systems are very attractive for molecular diagnosis.

[0112] In some embodiments, the CRISPR enzyme is mutated relative to the corresponding wild-type enzyme so that the mutated CRISPR enzyme lacks the ability to cleave one or both strands of a target polynucleotide containing a target sequence. Cas proteins without nucleolytic activity (dead Cas proteins; dCas) are known. Inactivation by mutation of two nuclease domains will produce catalytic dead Cas. The CRISPR enzyme can be Cas9, dCas9 (dead Cas9), dCas12, Cas12a, Cas12b, dCas13 or Cpf1, etc. Catalytically inactive or "dead" Cas9 (dCAS9) (or any Cas enzyme) is a mutant form of the protein that is unable to cut, but still tightly binds to a specific DNA sequence specified by a guide RNA. An example is Streptococcus pyogenes Cas9 (SpCas9), which can also be used in its deactivated form (SpdCas9), i.e., deactivated SpdCas9. In a preferred embodiment, the Cas effector is from Staphylococcus aureus (SauCas9), Leptotrichia wadeii (LwaCas13a), Leptotrichia buccalis (LbuCas13a) or Eubacterium siraeum (EsCas13d).

[0113] iii. Sealing agent

[0114] Optionally, the biorecognition layer comprises a blocking agent. Exemplary blocking agents include, but are not limited to, bovine serum albumin (BSA), ethanolamine (ETA), and casein. In some preferred embodiments, the blocking agent contained in the biorecognition layer is BSA. While not being bound by theory, the blocking agent is provided to capture potential contaminating proteins in a sample containing the analyte, which is contacted with the biorecognition layer to bind the analyte to its binding partner (i.e., biorecognition element) present in the biorecognition layer.

[0115] 2. Indirect Functionalization of OECTs

[0116] Alternatively, an indirect method can be used to functionalize the OECT, as described in PCT / IB2021 / 055981. For example, the biorecognition layer comprises two self-assembled monolayers (SAMs, the first of which is formed by organic molecules, which are chemically modified as disclosed in PCT / IB2021 / 055981, i.e., Chem-SAM, and the second of which is Bio-SAM, as disclosed herein). The Bio-SAM comprises a biorecognition element, such as a nanobody, a fusion protein, or a Cas protein as described above, preferably a nanobody.

[0117] The organization of the indirectly functionalized biorecognition layer on the OECT surface can be represented by the following general formula:

[0118] N-L1-AP1-AP2-L2-B

[0119] Formula I'

[0120] Wherein N is one or more organic molecules capable of self-assembly to form a first SAM, L1 is an optional first linker, AP1 is a first peptide binding partner; AP2 is a second peptide partner, L2 is a second linker, and B is a biorecognition element. The first SAM formed by N is chemically modified with AP1 to produce a Chem-SAM. AP1, AP2, L2 and B can be any of those described above for direct functionalization. Therefore, the biorecognition layer includes Chem-SAM and Bio-SAM. The first SAM on the gate electrode of the OECT is typically formed by reacting a plurality of organic molecules on the surface of the gate electrode of the OECT under conditions that result in self-assembly, after which it chemically reacts with L1-AP1 to form a Chem-SAM.

[0121] In contrast, the method for direct functionalization of the gate electrode of OECT excludes the step of contacting the gate electrode with a solution containing organic molecules capable of self-assembly (such as cysteine) before contacting the gate electrode with L1-AP1, and thereby eliminates the surface chemical interaction between the organic molecules and L1-AP1.

[0122] 3. OECT components

[0123] The disclosed device includes the above described biorecognition layer and its integration with a high gain ion to electron transducer device OECT. OECT is a three terminal transistor, which includes a source electrode, a drain electrode, a channel and a gate electrode. The channel electronically connects the source electrode and the drain electrode. The gate electrode is integrated with / functionalized with the biorecognition layer. The source electrode and the drain electrode are spaced apart and electronically connected through the corresponding channel.

[0124] In some embodiments, the source electrode, drain electrode, and channel can be patterned on a supporting substrate such as a glass substrate, a silicon substrate, or a plastic substrate (such as a polyimide substrate or a textile). For example, the source electrode, drain electrode, and channel are patterned on a glass substrate. The gate electrode is placed separately from the source electrode, drain electrode, and channel. The gate electrode is removable, that is, it is not physically connected to the supporting substrate on which the source electrode, drain electrode, and channel are patterned. This configuration allows the gate electrode to be easily handled when used. For example, the user can incubate the gate electrode with a blank solution or a biological sample and perform subsequent rinsing, away from the rest of the OECT component, and then place the gate electrode in position for measurement (see, e.g. Figure 1B and 8C ).

[0125] During measurement, the channel and optionally the source and drain electrodes are in contact with an electrolyte solution in which the gate electrode is also immersed. Optionally, the OECT includes a reservoir for containing the electrolyte solution. The reservoir (i.e., PDMS hole) is placed on top of the channel to contain the electrolyte solution so that the channel and the gate electrode are in contact with the electrolyte solution. Optionally, the reservoir is configured so that the electrolyte solution contained therein also contacts the source and drain electrodes, but is insulated with an insulator such as parylene or SU-8.

[0126] In some embodiments, the device includes more than one OECT in the form of an array. When the device contains an array of OECTs, each OECT may contain a source electrode, a drain electrode, and a corresponding channel, and all OECTs in the array use a common gate electrode. For example, if the OECT array includes multiple OECTs, such as 2, 3, 4, 5, or 6 OECTs, each of which contains a source electrode, a drain electrode, and a corresponding channel; however, all OECTs in the array use a common gate electrode integrated with the biorecognition layer. The source electrode, the drain electrode, and the corresponding channel of each of the OECTs are patterned on a supporting substrate. Optionally, when the device contains an array of OECTs, each OECT may contain a source electrode, a drain electrode, a corresponding channel, and a gate electrode, so that each gate electrode can be engineered to have the same or different biorecognition elements for simultaneously detecting one or more analytes. The device can be incorporated into a microfluidic configuration. An example of a device comprising an array of OECTs is shown in Figure 1B , 12A and 14 in.

[0127] The channel is typically made of an ion-permeable organic electronic material through which holes or electrons flow from the source electrode to the drain electrode. OECTs rely on ions injected from an electrolyte solution into an ion-permeable organic electronic material, thereby changing its doping state and thus its conductivity. The operation of an OECT is controlled by a voltage applied to a gate electrode (gate voltage, VG) and a voltage applied to a drain electrode (drain voltage, VD), which are referenced relative to the source electrode. The drain voltage induces a current (drain current, ID) that is proportional to the number of holes or electrons moving in the channel, and thus detects the doping state of the organic electronic material. The gate voltage controls the injection of ions into the channel, and thus controls the doping state of the organic electronic material, resulting in a change in ID. The change in ID can be expressed as a normalized response (also referred to herein as "NR"), which can be calculated based on Equation 1:

[0128] NR = (|ID-I0|) / I0 Equation 1

[0129] where I0 is the leakage current obtained after the OECT is incubated with a blank solution (ie, a solution without the presence of target analyte) under the same measurement conditions as the biological sample.

[0130] Due to the barrier-free penetration of electrolyte ions into the bulk of the organic channel, doping changes in OECTs occur over the entire volume of the channel, resulting in large modulations of the carrier density therein. The device translates a small ion flux in the electrolyte into a large electrical readout from the channel. Thus, the transduction event is combined with amplification and gives OECTs high gain at low voltages (<1 V).

[0131] i. Source electrode and drain electrode

[0132] The source electrode and the drain electrode are made of materials capable of conducting current. The electrode material can be organic or inorganic in nature, as long as it can conduct electrons and inject electronic charge into the channel material. The electrode can be a polymer electrode, a metal electrode, a carbon-based material, a metal oxide or a modified electrode. In some embodiments, the source electrode and the drain electrode are made of electrochemically inert materials (such as gold, platinum, chromium or carbon in a conductive form, or a combination thereof). In some embodiments, the source electrode and the drain electrode are gold electrodes. In some embodiments, the source electrode and the drain electrode are stacks of metal layers, such as a gold layer coated on a chromium layer. Each of the source electrode and the drain electrode can be insulated using an insulator such as polyparaxylene. In some embodiments, SU-8 (an epoxy-based photoresist (also used as an insulator) for patterning electronic devices with photolithography) is used to insulate the source electrode and the drain electrode.

[0133] In certain embodiments, the electrode is made of a metal conductor. Suitable metal conductors include, but are not limited to, gold, chromium, platinum, iron, nickel, copper, silver, stainless steel, mercury, tungsten and other metals suitable for electrode construction. The metal conductor can be a metal alloy made of a combination of metals described above, such as gold / chromium. In certain embodiments, the electrode can be formed by different layers of metal, such as a gold layer coated on a chromium layer. In addition, the conductive substrate as a metal conductor can be constructed by nanomaterials made of gold, cobalt, diamond and other suitable metals.

[0134] In other embodiments, the electrode is made of a carbon-based material. Exemplary carbon-based materials are conductive polymers (in the form of films or fibers), carbon cloth, carbon paper, carbon screen printed electrodes, carbon paper, carbon black, carbon powder, carbon fibers, single-walled carbon nanotubes, double-walled carbon nanotubes, multi-walled carbon nanotubes, carbon nanotube arrays, diamond-coated conductors, glassy carbon, and mesoporous carbon. In addition, other exemplary carbon-based materials are graphene, graphite, uncompressed graphite worms, delaminated purified flake graphite, high-performance graphite and carbon powder, highly ordered pyrolytic graphite, pyrolytic graphite, and polycrystalline graphite.

[0135] The electrodes may be doped semiconductors. Suitable semiconductors are made of silicon and germanium, which may be doped with other elements (i.e., impurities are intentionally introduced into an intrinsic semiconductor to modulate its electrical and structural properties). The semiconductor may be doped with phosphorus, boron, gallium, arsenic, indium, or antimony, or a combination thereof.

[0136] Other electrode materials can be metal oxides, metal sulfides, main group compounds and modified materials. Exemplary materials of this type are nanoporous titanium oxide, tin oxide coated glass, glass, cerium oxide particles, molybdenum sulfide, boron nitride nanotubes, aerogels modified with conductive materials such as gold, sols modified with conductive materials such as carbon, ruthenium carbon aerogels, and mesoporous silica modified with conductive materials such as gold.

[0137] The source electrode and the drain electrode can be any suitable shape, such as a cuboid, a cube, a circle, and a cylinder. In some preferred embodiments, the electrode is a cuboid gold electrode or a cuboid gold-coated chromium electrode. In some embodiments, each of the source electrode and the drain electrode has a first size (i.e., width), a second size (i.e., length), and a third size (i.e., thickness). For example, each of the source electrode and the drain electrode has a width in the range of 100 μm to 1 mm (e.g., 0.8 mm), a length in the range of 100 μm to 1 mm (e.g., 0.8 mm), and a thickness of about 100 nm.

[0138] ii. Channel

[0139] Generally speaking, the channel is configured to establish an electrochemical connection between a pair of source and drain electrodes so that holes or electrons flow from the source electrode to the drain electrode.

[0140] The channel generally contains an ion-permeable organic electronic material, such as a conductive polymer disclosed in, for example, Rivnay et al., Nature Reviews, 3: 17086 (2018) and Sun et al., J. Mater. Chem. C, 6: 11778-11784 (2018). The channel of an OECT can be formed directly from a conductive polymer or by incorporating a conductive polymer on the surface of a conductive substrate. For example, a conductive polymer can be spin-coated, drop-cast, inkjet-printed, or screen-printed on the surface of a conductive substrate such as a gold substrate or a gold-coated chromium substrate. Any suitable conductive polymer can be used for the channel of an OECT. In some embodiments, the conductive polymer used for the channel of an OECT is a p-type polymer. In some embodiments, the conductive polymer used for the channel of an OECT is an n-type polymer. Depending on the channel material, the OECT can be operated in a depletion mode or an accumulation mode.

[0141] An exemplary conductive polymer for the channel is poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonate) (PEDOT:PSS). The conductive PEDOT is p-type doped (oxidized), which causes mobile holes to jump from one chain to another, resulting in a hole current once a drain voltage is applied. These holes are compensated by the sulfonate anions of the PSS. A channel made of PEDOT:PSS can charge a depletion mode OECT. For example, in the absence of a gate voltage, a hole current flows in the channel. Once a positive gate voltage is applied, cations from the electrolyte are injected into the channel, and the anions are compensated, resulting in a decrease in the number of holes in the channel. This results in a decrease in the drain current. Alternatively, the channel can be made of materials that work in accumulation mode OECTs, such as semiconductors based on: polythiophenes with sulfonic acid groups attached to the main chain as hexyl chains (PTHS) (Inal et al., Adv. Mater., 26: 7450-7455 (2014)), ethylene glycol units attached to bithiophene (Moser et al., Adv. Mater., 32: 2002748 (2020)) or g7-NDI-Br and NDI-T2 copolymers (P-90) (Giovannitti et al., Chem. Mater., 30: 2945-2953 (2018)). In accumulation mode p-type OECTs, the application of a negative gate voltage leads to the injection of anions into the channel and to the corresponding accumulation of holes, resulting in an increase in the drain current ID. It is also possible to use n-type semiconductors and establish accumulation mode n-type OECTs. In an accumulation-mode n-type OECT, application of a positive gate voltage results in the injection of cations into the channel and causes a corresponding accumulation of electrons, leading to an increase in the drain current ID.

[0142] Other suitable conducting polymers for the channel include, but are not limited to, conductors based on PEDOT with sulfonic acid side groups (PEDOT-S), PEDOT doped with toluenesulfonate (PEDOT:TOS), PEDOTOH:ClO4, PEDOT-co-PEDOTOH:ClO4 (Schmode et al., Chem. Mater., 31(14):5286-5295 (2019)), poly( 2-((3,3′-bis(2-(2-(2-methoxyethoxy)ethoxy)ethoxy)-[2,2′bithiophene]-5-yl)thieno[3,2b]thiophene), p(g2TTT), poly(ethoxy)ethyl 2-(2-(2-methoxyethoxy)ethoxy)acetate)-naphthalene 1,4,5,8-tetracarboxylic acid-imide-co-3,3′bis(2-(2-(2-methoxyethoxy)ethoxy)ethoxy)-(bithiophene)) p(gNDI g2T), P3HT (poly (3-hexylthiophene-2,5-diyl)), BBL (polybenzimidazole-benzoisoquinoline), PTHS-TMA+-co-P3HT (poly [(6-thiophene-3-yl) hexane-1-sulfonic acid-co-3-(hexylthiophene)]), poly (N,N'-bis (7-ethylene glycol)-naphthalene-1,4,5,8-bis (dicarboximide)-co-2,2'-bithiophene-co-N,N '-Bis(2-octyldodecyl)-naphthalene-1,4,5,8-bis(dicarboximide), a polymer (P-90) based on naphthalene-1,4,5,8-tetracarboxylic acid-diimide-bithiophene (NDI-T2) with a 90% ethylene glycol chain percentage, and glycolated diketopyrrolopyrrole (DPP) or glycolated thiophene, bithiophene such as p(g0T2-g6T2) (Moser et al., Advanced Materials, 32: 2002748 (2020)).

[0143] Examples of suitable polymers for the channel in depletion mode OECTs are PEDOT polymers or copolymers thereof, such as polymers based on: PEDOT with sulfonic acid side groups (PEDOT-S), PEDOT doped with toluenesulfonate (PEDOT:TOS), PEDOTOH:ClO4, PEDOT-co-PEDOTOH:ClO4 and PEDOT:PSS.

[0144] Examples of suitable polymers for the channel in accumulation mode OECTs are poly(2-((3,3′-bis(2-(2-(2-methoxyethoxy)ethoxy)ethoxy)-[2,2′bithiophene]-5-yl)thieno[3,2b]thiophene), p(g2TT), poly(ethoxy)ethyl 2-(2-(2-methoxyethoxy)ethoxy)acetate)-naphthalene 1,4,5,8-tetracarboxylic acid-imide-co-3,3′bis(2-(2-(2-methoxyethoxy)ethoxy)ethoxy)-(bithiophene)), p(gNDI g2T), P3HT (poly (3-hexylthiophene-2,5-diyl)), BBL (polybenzimidazole-benzoisoquinoline), PTHS-TMA+-co-P3HT (poly [(6-thiophene-3-yl) hexane-1-sulfonic acid-co-3-(hexylthiophene)]), poly (N,N'-bis (7-ethylene glycol)-naphthalene-1,4,5,8-bis (dicarboximide)-co-2,2'-biphenyl) Thiophene-co-N,N'-bis(2-octyldodecyl)-naphthalene-1,4,5,8-bis(dicarboximide), a polymer (P-90) with a 90% ethylene glycol chain percentage based on naphthalene-1,4,5,8-tetracarboxylic acid-diimide-bithiophene (NDI-T2), and glycolated diketopyrrolopyrrole (DPP) or glycolated thiophene, bithiophene such as p(g0T2-g6T2).

[0145] In some embodiments, the conductive polymer for the channel of the OECT is poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonate) (PEDOT:PSS) or any other mixed (ionic and electronic) semiconductor. In some embodiments, the conductive polymer for the channel of the OECT is p(g0T2-g6T2). In some embodiments, the conductive polymer for the channel of the OECT is p(g3C2T2-T). In some embodiments, the conductive polymer for the channel of the OECT is p(C6NDI-T).

[0146] In general, the channel has a first dimension (i.e., width), a second dimension (i.e., length), and a third dimension (i.e., thickness). The length is the distance between the source electrode and the drain electrode. The width is the remaining size of the rectangle. The values ​​of the length and width of the channel can be between about 5 μm and about 5 mm, for example, between 5 μm and 1000 μm, between 5 μm and 100 μm, between 5 μm and 100 μm, or between 10 μm and 100 μm, for example, about 10 μm or about 100 μm. The width of the channel can be between 10 μm and 1000 μm, between 20 μm and 500 μm, or between 20 μm and 200 μm, for example, about 100 μm. The thickness of the channel may have a value between about 10 nm and about 200 nm, such as between 10 nm and 150 nm, between 20 nm and 100 nm, between 50 nm and 150 nm or between 50 nm and 100 nm, such as about 85 nm.

[0147] The first size and the second size may be the same or different. In some embodiments, the width of the channel is the same as the length of the channel. For example, the channel has a width of about 100 μm and a length of about 100 μm. In some embodiments, the width of the channel is less than the length of the channel. In some preferred embodiments, the width of the channel is greater than the length of the channel. For example, the channel has a width of about 100 μm and a length of about 10 μm. In some embodiments, an OECT with a small channel (e.g., a width of about 100 μm and a length of about 10 μm) is more sensitive (i.e., NR is higher) than an OECT with the same but larger channel (e.g., a width of about 100 μm and a length of about 100 μm) under the same measurement conditions (e.g., VG, VD, incubation time, temperature, and pressure).

[0148] In general, the channel remains unmodified, i.e., it is not chemically modified to anchor the binding partner. Analyte recognition occurs at a functionalized gate electrode that is not in physical contact with the channel. This configuration enables the device to have an extended shelf life compared to OECTs containing a channel modified with a binding partner. For example, the OECTs disclosed herein can maintain their leakage current (i.e., a change of less than about 5% in leakage current compared to a fresh sensor) after being stored in air for at least 6 months under ambient conditions (i.e., about 25° C. at 1 atm) or in a nitrogen or vacuum environment for at least 1 year. For example, the OECTs disclosed herein can maintain their NR determined using a standard solution containing the same amount of the same analyte (i.e., a change of less than about 5% in NR compared to a fresh sensor) after being stored in air for 6 months under ambient conditions (i.e., about 25° C. at 1 atm) or in a nitrogen or vacuum environment for 1 year.

[0149] iii. Gate electrode

[0150] The gate electrode is configured to control the injection of ions into the channel and is typically placed in the measurement solution along with the source electrode, drain electrode, and channel, but is not in physical contact with these components.

[0151] The gate electrode can be any suitable shape, such as a cuboid, a cube, a circle, and a cylinder. In some embodiments, the gate electrode is a circular electrode having a diameter in the range of 0.5 mm to 10 mm, 1 mm to 10 mm, 0.5 mm to 5 mm, or 1 mm to 5 mm, such as 2.8 mm or 5 mm. In some embodiments, the gate electrode is a circular electrode with a diameter of about 5 mm. In some embodiments, the gate electrode is a square electrode having a length in the range of 0.1 mm to 10 mm, 0.1 mm to 8 mm, 0.1 mm to 5 mm, 0.5 mm to 10 mm, or 0.5 mm to 5 mm, such as about 0.8 mm or about 4.8 mm. In some embodiments, the gate electrode is a square electrode with a length of about 0.8 mm.

[0152] In some embodiments, the OECT contains a gate electrode having any shape and size as described above and a channel having any shape and size as described above. In some embodiments, the OECT contains a square gate electrode having any length as described above and a rectangular channel having any size (i.e., width and length) as described above. In some embodiments, the OECT contains a circular gate electrode having any diameter as described above and a rectangular channel having any size (i.e., width, length, and thickness) as described above. For example, the OECT contains a square gate electrode having a length of about 0.8 mm and a rectangular channel having a width of about 100 μm and a length of about 10 μm. For example, the OECT contains a circular gate electrode having a diameter of about 5 mm and a rectangular channel having a width of about 100 μm, a length of about 10 μm, and a thickness of about 85 nm.

[0153] In general, the gate electrode can be made of a substrate coated with any of the conductive materials described above for the source and drain electrodes (e.g., Pt, Au, Cr, Au-coated Cr, etc.), or a conductive polymer such as poly(fluorine), polyphenylene, polypyrene, polyazulene, polynaphthalene, polypyrrole, polycarbazole, polyindole, polyazacycloheptane, polyaniline, polythiophene, poly(3,4-ethylenedioxythiophene), poly(p-phenylene sulfide), poly(acetylene), or poly(p-phenylene vinylene). In some embodiments, the gate electrode contains two or more conductive materials. For example, the gate electrode contains two conductive materials, wherein the first conductive material can be the material disclosed above for the source and drain electrodes, and the second conductive material can be a conductive polymer. In some embodiments, the gate electrode contains two conductive materials, wherein the first conductive material is a first type of metal conductor, and the second conductive material is a second type of metal conductor. The two conductive materials can be coated on the substrate simultaneously or successively. An exemplary gate electrode is formed from a Kapton (polyimide) substrate that is subsequently sputter coated with Cr and Au. Another exemplary gate electrode is formed from a glass substrate that is subsequently sputter coated with Cr and Au.

[0154] In general, the gate electrode is engineered to include a biorecognition layer. For example, the gate electrode is sequentially modified with a first peptide binding partner via a first linker and with a biorecognition element (such as a nanobody) via a second peptide binding partner to form a biorecognition layer on the gate electrode (i.e., direct functionalization). For example, the gate electrode is sequentially modified with a thiol-containing organic molecule, with a first peptide binding partner via a first linker, and with a biorecognition element (such as a nanobody) via a second peptide binding partner to form a biorecognition layer on the gate electrode (i.e., indirect functionalization).

[0155] iv. Electrolyte solution

[0156] When used, the electrolyte solution is in electrical contact with the channel and the gate electrode, and optionally in electrical contact with the source electrode and the drain electrode (the latter two electrodes are insulated with an insulator). The electrolyte solution is a solution containing ions or molecules that lose or gain electrons and is conductive and mostly water. The electrolyte solution includes but is not limited to buffers, such as water, phosphate buffered solution, phosphate buffered saline, salt water, MES buffer, Bis-Tris buffer, ADA, ACES, PIPES, MOPSO, Bis-Tris propane, BES, MOPS, TES, HEPES, DIPSO, MOBS, TAPSO, Trizma, HEPPSO, POPSO, TEA, EPPS, Tricine, Gly-gly, Bicine, HEPBS, TAPS, AMPD, TABS, AMPSO, CHES, CAPSO, AMP, CAPS, CABS. The electrolyte solution may have a pH between about 4 and about 8.5, between about 4.5 and about 8.5, between about 5 and about 8.5, between about 5.5 and about 8.5, between about 6 and about 8, or between about 6.5 and about 7.5, preferably about 7.4.

[0157] v. Storage

[0158] The device may include a reservoir for containing an electrolyte solution. The reservoir may be incorporated into the OECT by any suitable method. For example, the reservoir is glued or molded on top of the channel. The reservoir may also be a microfluidic channel. When more than one channel is included in the device as an array, the reservoir is incorporated into the device so that the electrolyte solution contained therein is in contact with all channels and gate electrodes when in use.

[0159] The reservoir is generally defined by a sidewall and a bottom surface, and contains an opening configured to allow an electrolyte solution to enter the reservoir. The reservoir can have any suitable shape, such as a cylindrical hole, a cubic shape, or a cuboid shape. The bottom surface is formed by at least a portion of the channel, and optionally, at least a portion of the sidewall is formed by a source electrode and a drain electrode, respectively. For example, the reservoir is a cylindrical hole and has a bottom surface formed by at least a portion of the channel. The sidewall is perpendicular to the surface of the channel, and at least a portion of the sidewall is in contact with the source electrode and the drain electrode. The cylindrical reservoir contains an opening to allow an electrolyte solution to enter the reservoir. Figure 1B , 8C Exemplary OECTs containing rectangular, trapezoidal, or cylindrical reservoirs to contain electrolyte solutions are shown in and 12A.

[0160] For example, the reservoir is a rectangular parallelepiped defined by a bottom surface and four side walls. The bottom surface is formed by at least a portion of the channel, and each of the first side wall and the second side wall is parallel to each other and vertically placed on top of each of the source electrode and the drain electrode. The third side wall and the fourth side wall are parallel to each other and perpendicular to the first side wall and the second side wall, so that the reservoir contains an opening to allow the electrolyte solution to enter the reservoir.

[0161] For example, the reservoir is a rectangular parallelepiped and has a bottom surface formed by the channel, a first side wall formed by a portion of the source electrode, a second side wall formed by a portion of the drain electrode, and a third side wall and a fourth side wall parallel to each other and perpendicular to the first side wall and the second side wall, so that the reservoir contains an opening to allow the electrolyte solution to enter the reservoir. When the electrolyte solution is in the reservoir, at least a portion of the source electrode and the drain electrode, the channel, and the gate electrode are in contact with the electrolyte solution.

[0162] The reservoir may be made of any suitable inert material such as plastic, glass, or a polymeric material such as polydimethylsiloxane (PDMS).

[0163] B. Convection-driven ultrafast detection transistor

[0164] One embodiment provides a directly nanobody-functionalized organic electrochemical transistor (directly functionalized OECT), while another embodiment provides an indirectly nanobody-functionalized organic electrochemical transistor (indirectly functionalized OECT), both of which combine the microstirring effect of alternating current thermal flow (ACET) for ultrafast detection of pathogens such as severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) at single-molecule to nanomolar levels in complex body fluids.

[0165] ACET-enhanced OECT-based immunosensors can use directly functionalized OECTs or indirectly functionalized OECTs, such as any of those described above, and a conductive layer placed in close proximity (i.e., less than 1 μm shortest edge-to-edge distance, measured from the edge of the gate electrode to the edge of the conductive layer) to the gate electrode of the OECT to induce ACET. In general, the conductive layer is placed along the edge of the gate electrode and surrounds at least a portion of the gate electrode. Optionally, the ACET-enhanced OECT-based immunosensor includes more than one conductive layer, wherein each conductive layer is placed along the edge of the gate electrode, and the conductive layers together surround at least a portion of the gate electrode. The conductive layer can have any suitable shape as long as it can be placed along the edge of the gate electrode and surround at least a portion of the gate electrode. For example, as Fig. 8AAs shown, the gate electrode 10 is circular in shape, and the arched conductive layer 20 is placed along the edge 12 of the gate electrode and concentrically surrounds at least a portion of the circle 10. Optionally, the conductive layer contains one or more extensions (see, for example, Fig. 8A 22a and 22b in the embodiment of the present invention) to facilitate the application of an AC potential from a power source to the conductive layer. When used, ACET can be induced by optionally applying an AC potential to the conductive layer via one or more extensions. The induced ACET generates an electrohydrodynamic force on the fluid, which stirs the solute molecules through the induced flow (also referred to herein as "micro-stirring effect"), allowing the target molecule (such as a spike protein) to accumulate at the gate electrode on an immobilized binding partner (such as a nanobody) for the target molecule.

[0166] The conductive layer may be formed using any conductive material suitable for forming an electrode, such as any of those materials described above for forming the source electrode, drain electrode, channel and gate electrode. For example, the conductive layer is formed of gold or gold-coated chromium.

[0167] The disclosed ACET-enhanced OECT-based sensors are based at least on the following findings (through numerical simulations and experimental studies): ACET-induced mixing can significantly reduce the time for immune complex formation (≤2 min from sample to result) and can achieve higher specificity and lower background compared to the same OECT-based immunosensor without ACET, due to the electrothermal flow-induced removal of nonspecific species from the sensor surface. The ACET-enhanced OECT-based immunosensor also exhibits low energy consumption (i.e., nW level, such as about 100 nW). For example, the ACET-enhanced OECT-based sensor can be operated using only 100 nW power and about 2 min of incubation with 10 μL or 5 μL sample to detect target analytes in complex media.

[0168] ACET flow induced by the biased AC electric field can rapidly convect analytes to the concentric gate electrodes within 2 minutes (such as within one minute) and capture the SARS-CoV-2 spike protein (S1) or receptor binding domain (RBD) via the recognition unit (nanoantibody), while sweeping non-specifically bound analytes from the surface.

[0169] The sensor includes a solution-processable conjugated polymer in the transistor channel and a variety of recognition units on the gate electrode, for example, a high-density and orientation-controlled bioconjugation of a nanobody-SpyCatcher fusion protein on the gate electrode, which is directly or indirectly functionalized on the gate electrode as described above. The synergistic effect of ACET and OECT provides results after exposure to 10 μL or 5 μL of sample for 2 min or 1 min, maintaining high specificity and single-molecule sensitivity for biological samples (such as human saliva) while consuming low power, such as ≤100 W.

[0170] II. Preparation Method

[0171] The disclosed biosensor integrates OECT technology, which is engineered to include a biorecognition layer by functionalizing the sensing electrode (i.e., gate electrode) with a binding partner (preferably, a covalent binding partner or affinity pair (e.g., SpyTag / SpyCatcher conjugation system)) and a biorecognition element. Optionally, the biorecognition layer is also exposed to a blocking agent, which can be introduced simultaneously with the biorecognition element or subsequently to the biorecognition element.

[0172] Methods for preparing OECTs or arrays of two or more OECTs are known. An exemplary method for preparing an array of OECTs is described in the Examples. The methods described herein can be used to immobilize biorecognition elements, such as nanobody units, in a uniform orientation; this makes the biorecognition layer more precise and compact.

[0173] Generally speaking, a method for integrating a biorecognition layer on an electrode, such as a gate electrode of an OECT, comprises: (i) contacting at least a portion of the surface of the electrode with a first solution containing a peptide sequence preferably comprising a cysteine ​​residue at its N or C terminus, a peptide linker and a first peptide binding partner to produce a first peptide binding partner modified OECT surface; and (ii) contacting the first peptide binding partner modified surface with a second solution containing a second peptide (which is a binding partner of the first peptide), a biorecognition element for a target analyte and optionally containing a blocking agent, wherein the first peptide is conjugated to the second peptide to form a monolayer, referred to herein as a Bio-SAM. The method for functionalizing the gate electrode of an OECT excludes the step of contacting at least a portion of the surface of the gate electrode with a solution containing a thiol-containing organic molecule, such as cysteine, to form a Chem SAM prior to step (i). That is, the method for electrode functionalization disclosed herein does not form a Chem SAM, and then chemically attaches the first linker and the first peptide binding partner to the Chem SAM, which would produce free organic molecules on the gold surface (i.e., organic molecules in the Chem SAM that are not attached to the first peptide binding partner and therefore not attached to the biorecognition element).

[0174] Modification of the OECT electrode by Bio-SAM results in a modified OECT gate electrode surface, while the recognition element is exposed to interact with (and bind to) the analyte from the sample. (See Figure 1A and 1B ). The second solution preferably comprises a blocking agent molecule, such as BSA, with or without a mild detergent, e.g. 20 (polyethylene glycol sorbitan monolaurate).

[0175] Referring to Formula I, (i) the surface of the electrode is contacted with a composition comprising AP1 and L1 under conditions that result in chemical coupling of AP1 to the surface of the OECT via L1, thereby forming an AP1-L1-OECT electrode surface, and (ii) the AP1-L1-OECT electrode surface is contacted with a composition comprising AP2-L2-B- under conditions that result in conjugation of AP1 and AP2 to form a biological self-assembled monolayer, herein Bio-SAM. ( Figure 1A ). The second solution in step (ii) preferably comprises a blocking agent molecule, such as BSA, with or without a mild detergent, e.g. 20 (polyethylene glycol sorbitan monolaurate). When spyTag and Spycatcher are used as binding partners, spyTag is AP1 and spyCatcher is Ap2, i.e. the surface of the OECT is first contacted with a solution containing spyTag. This is in contrast to the procedure described in Oloketuyi et al., Biosenbsores and Bioelectronics, 154: 112052 (2020), which relies on random chemical immobilization of the spyCatcher on the surface described herein. The coupling of COOH from the spyCatcher to the activated Cys on the surface in Oloketuyi et al., Biosenbsores and Bioelectronics, 154: 112052 (2020) is random because it can occur through any acidic residue (Asp or Glu) of the protein or through more than one of the acidic residues. Therefore, there is no control over the orientation of the spyCatcher. This then has three consequences: (1) control over orientation is lost at that point; (2) a large portion of the spyCatcher will not be able to react with the spyTag due to steric or chemical blocking or unfolding, resulting in lower density coupling; and (3) flexibility is lost because the spyCatcher adheres directly to the surface (this is in contrast to the disclosed method, in which the spyCatcher dangles on a flexible linker above the OECT surface). The disclosed method overcomes these limitations, in which the spyTag is first attached to a gold surface, after which the spyTag-modified surface is contacted with a solution containing the spyCatcher. In contrast to the "spyCatcher first" approach, the disclosed method retains complete orientation control throughout the immobilization stack. For example, due to the lack of chemical modification or adsorption, it retains complete spyCatcher integrity and allows the spyCatcher to dangle on a flexible linker above the OECT surface. This maximizes the ability of the entire system to reorient itself as needed for close packing and binding of analytes.

[0176] A. Modification with the first peptide

[0177] Generally speaking, the first peptide (of the binding pair) is dissolved in a buffer as described above (such as PBS) to form a first incubation composition. The first peptide is a modified peptide comprising a short linker. Preferred linkers include CGGSGSGSG (SEQ ID NO: 22) and GSGC (SEQ ID NO: 23). Exemplary modified first peptides (of the binding pair) include: CGGSGSGSGAHIVMVDAYKPTK (SEQ ID NO: 24) and AHIVMVDAYKPTKGSGC (SEQ ID NO: 25). The electrode surface is immersed in the first incubation composition containing the first peptide for a period of time sufficient to couple the first peptide to the OECT surface, so that an OECT surface modified by the first peptide is produced. Generally, the incubation period sufficient to couple the first peptide to the SAM on the surface of the electrode is in the range of 10 minutes to 24 hours, 10 minutes to 20 hours, 10 minutes to 15 hours, 10 minutes to 12 hours, 10 minutes to 10 hours, 10 minutes to 5 hours, 10 minutes to 2 hours, 10 minutes to 1.5 hours or 10 minutes to 1 hour, such as about 1 hour.

[0178] The first peptide can be coupled and conjugated to the second peptide as described below. An exemplary first peptide is the SpyTag peptide, which is unmodified or preferably modified with a short linker (SEQ ID NO: 1) (disclosed in Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012)), preferably not modified with a maleimide functional group).

[0179] Generally, the concentration of the first peptide in the second incubation solution is in the range of 0.01 mg / mL to 20 mg / mL, 0.01 mg / mL to 15 mg / mL, 0.01 mg / mL to 10 mg / mL, 0.05 mg / mL to 20 mg / mL, 0.05 mg / mL to 15 mg / mL, 0.05 mg / mL to 10 mg / mL, 0.01 mg / mL to 5 mg / mL, 0.01 mg / mL to 1 mg / mL, 0.05 mg / mL to 5 mg / mL, or 0.05 mg / mL to 1 mg / mL, such as about 0.1 mg / mL.

[0180] Optionally, after incubation, the first peptide-SAM modified electrode surface is rinsed with a second rinse composition such as a buffer as described above.

[0181] B. Methods for producing fusion proteins containing biological recognition elements

[0182] The fusion protein of formula II can be obtained by, for example, chemical synthesis, and more preferably by recombinant production in host cells. In order to recombinantly produce the fusion protein of formula II, a nucleic acid containing a nucleotide sequence encoding a polypeptide can be used to transform, transduce or transfect bacteria or eukaryotic host cells (e.g., insects, yeast or mammalian cells). Generally speaking, a nucleic acid construct comprises a regulatory sequence operably connected to the nucleotide sequence of the fusion protein of formula II encoding. Regulatory sequences (also referred to herein as expression control sequences) generally do not encode gene products, but affect the expression of nucleic acid sequences operably connected thereto. The nucleotide sequence encoding the fusion protein is usually inserted into a recombinant vector, which can be any vector that can conveniently carry out a recombinant DNA procedure, and the selection of the vector will generally depend on the host cell into which the vector is introduced. Therefore, the vector can be an autonomously replicating vector, i.e., a vector that exists as an extrachromosomal entity and is independent of chromosome replication, such as a plasmid. Alternatively, the vector can be a vector that is integrated into the host cell genome when introduced into the host cell and replicated together with the chromosome into which it has been integrated. The vector is preferably an expression vector, in which the DNA sequence encoding the fusion protein is operably connected to the additional fragment required for the transcription of the DNA. Generally speaking, expression vectors are derived from plasmids or viral DNA, or may contain elements of both. The term "operably linked" indicates that the fragments are arranged so that they work together for their intended purpose, for example, transcription is initiated in a promoter and continues through a DNA sequence that is decoded for the fusion protein. The expression vector for expressing the fusion protein will include a promoter that can direct the transcription of a cloned gene or cDNA. The promoter can be any DNA sequence that exhibits transcriptional activity in the selected host cell, and can be derived from a gene that encodes a protein that is homologous or heterologous to the host cell. The expression vector for expressing the fusion protein will include a promoter that can direct the transcription of a cloned gene or cDNA. The promoter can be any DNA sequence that exhibits transcriptional activity in the selected host cell, and can be derived from a gene that encodes a protein that is homologous or heterologous to the host cell. Examples of suitable promoters for directing transcription of DNA in mammalian cells are the SV40 promoter (Subramani et al., Mol. Cell. Biol. 1 (1981), 854-864), the MT-1 (metallothionein gene) promoter (Palmiter et al., Science 222 (1983), 809-814), the CMV promoter (Boshart et al., Cell 41:521-530, 1985), or the adenovirus 2 major late promoter (Kaufman and Sharp, Mol. Cell. Biol, 2:1304-1319, 1982).

[0183] Useful prokaryotic and eukaryotic systems for expression and production of polypeptides are well known in the art and include, for example, E. coli strains such as BL-21, and cultured mammalian cells such as CHO cells.

[0184] In eukaryotic host cells, many viral-based expression systems can be used to express the fusion protein of II. Virus-based expression systems are well known in the art and include, but are not limited to, baculovirus vectors, SV40 vectors, retroviral vectors, or vaccinia-based viral vectors.

[0185] The expressed tagged protein or fusion protein produced by the cell can be recovered from the culture medium by conventional procedures, which include separating the host cells from the culture medium by centrifugation or filtration, releasing the fusion protein by mechanical cell disruption such as ultrasonic treatment or pressure, and precipitating the protein aqueous component of the supernatant or filtrate by salt such as ammonium sulfate. After ultrasonic treatment, the NaCl of suitable concentration can be added to further reduce the ability of the host cell pollutants to be combined with the cation exchange matrix. After the cation exchange chromatography, the fusion protein can be eluted according to the salt gradient and the eluent fraction containing the fusion protein can be collected. In some preferred forms, the fusion protein is captured from the lysate by its His tag. Therefore, IMAC (immobilized metal affinity chromatography) is used, and then after concentrating the protein-containing fraction, they are subjected to size exclusion chromatography (SEC) to carry out final purification. In a particularly preferred embodiment for nano antibody purification, nano antibodies are purified from the periplasmic space, wherein the host cell is a bacterium, such as Escherichia coli. This would include (1) centrifugation, (2) osmotic shock to release proteins from the cell wall compartment, (3) IMAC (immobilized metal ion affinity chromatography), (4) SEC (size exclusion chromatography).

[0186] C. Modification with biorecognition elements

[0187] Generally speaking, the second peptide-biorecognition element conjugate and the blocking agent are dissolved in a blocking buffer solution to form a second incubation solution, the blocking buffer solution comprising a buffer such as HEPES. Preferably, the blocking agent contained in the second incubation solution is BSA.

[0188] Generally, the concentration of the blocking agent in the second incubation solution is in the range of 0.01% w / v to 10% w / v, 0.01% w / v to 5% w / v, 0.01% w / v to 1% w / v, 0.05% w / v to 10% w / v, 0.05% w / v to 5% w / v or 0.05% w / v to 1% w / v, such as about 0.1% w / v.

[0189] The electrode surface modified with the first peptide is immersed in a second incubation solution containing a second peptide-biorecognition element conjugate and a blocking agent. Generally speaking, the second peptide of the peptide-biorecognition element conjugate can be conjugated to the first peptide to form a biomolecular linker, and thereby attach the biorecognition element to the electrode surface. An exemplary second peptide is a SpyCatcher peptide as disclosed in: Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012). For example, a SpyTag / SpyCatcher linker is formed to control the directional immobilization of a biorecognition element (e.g., a nanobody) on an electrode surface. After the first peptide binding partner-second peptide binding partner linker is formed, the biorecognition element (e.g., a nanobody) is functionalized on the gate electrode surface of the OECT in a configuration perpendicular to the electrode surface and perpendicular to the channel surface, thereby allowing binding to an analyte in the sample when the sample contacts the electrode functionalized with the biorecognition element. The biorecognition element can vary depending on the target analyte. For example, the biological recognition element is a nanobody, such as an anti-GFP nanobody or an anti-SARS-1-RBD nanobody or an anti-MERS RBD nanobody (see Table 1).

[0190] Table 1. Protein sequences

[0191]

[0192]

[0193] The conjugation between SpyTag and SpyCatcher is a robust method for conjugating recombinant proteins, in which the peptide SpyTag can react spontaneously with the protein SpyCatcher in a simple manner and with high specificity (Zakeri et al., Proc. Natl. Acad. Sci., 109: E690-E697 (2012); Keeble et al., 2017). An additional advantage of the disclosed technology is that there is no need to construct a maleimide functional group in the SpyTag peptide. The SpyTag peptide can be chemically synthesized by any known method described above. Nanobodies fused to the SpyCatcher protein can be expressed as a fusion protein in any suitable protein expression system, preferably with a tag such as a His tag to aid in its purification.

[0194] The electrode surface modified with the first peptide is incubated with the second incubation solution for a period of time sufficient to form a linker (i.e., a first peptide / second peptide conjugate), resulting in an electrode surface modified with a biorecognition element-linker-SAM (with a biorecognition layer integrated on the electrode surface).

[0195] Generally, the incubation period sufficient to form the second peptide / first peptide conjugate on the surface of the electrode is in the range of 5 minutes to 24 hours, 10 minutes to 20 hours, 10 minutes to 15 hours, 10 minutes to 12 hours, 10 minutes to 10 hours, 10 minutes to 5 hours, 10 minutes to 2 hours, 10 minutes to 1.5 hours or 10 minutes to 1 hour, such as about 1 hour.

[0196] In general, the concentration of the second peptide-recognition element conjugate in the second incubation solution is in the range of 1 μM to 100 μM, 1 μM to 100 μM, 1 μM to 90 μM, 1 μM to 80 μM, 10 μM to 100 μM, 20 μM to 90 μM, 10 μM to 80 μM, 5 μM to 100 μM, 5 μM to 90 μM, 5 μM to 80 μM, 20 μM to 100 μM, 20 μM to 90 μM or 20 μM to 80 μM, such as about 50 μM.

[0197] Optionally, after incubation, the biorecognition layer-integrated electrode surface is rinsed with a third rinse solvent, such as a buffer as described above.

[0198] D. Exemplary Methods for Electrode Functionalization

[0199] Exemplary methods for preparing biorecognition layer-integrated electrode surfaces are described in the Examples.

[0200] The preferred method uses p-type depletion (PEDOT:PSS) and accumulation (such as p(g0T2-g6T2)) mode materials in the OECT channel, while the gate electrode is gold. The gate electrode (preferably gold) surface is functionalized with a biorecognition layer (i.e., a biorecognition element, such as a nanobody). The gate electrode surface is electrochemically cleaned and then exposed to the first and second incubation compositions as described above.

[0201] The methods disclosed above produce OECT devices containing a biological recognition element whose binding partner is any target analyte, such as an antigen from any pathogen.

[0202] III. How to use

[0203] OECT devices containing Cas proteins and gRNAs whose binding partners are any target RNA (e.g., RNA from any pathogen) can be used to detect target pathogens in samples. OECT devices containing Cas proteins and gRNAs whose binding partners are any target RNA (e.g., RNA from any pathogen) can be used to detect target pathogens in samples.

[0204] A. Direct functionalization of OECTs

[0205] The disclosed device is used to measure the presence, absence or concentration of any analyte in a sample, wherein the analyte is a binding partner of a biorecognition element functionalized onto the device. The biorecognition element functionalized OECT sensor has the following features: improved sensitivity (aM or zM vs. fM) compared to prior art devices, and improved selectivity, providing accurate results according to gold standard tests (correlated with DNA or RNA detection results of conventional PCR methods where possible and applicable); rapid detection time and high detection limit (the lower limit of detection (LOD) exemplified below for SARS-CoV-2S1 in saliva is 6×10 -22 M). In addition, the peptide-based biorecognition element (which omits the Chem-SAM) gate exhibits improved chemical and electrochemical stability compared to the HDT SAM-based gate

[0206] OECT sensors rely on a channel made of or containing ion-permeable organic electronic materials, through which holes or electrons flow from the source electrode to the drain electrode. When ions are injected into the channel from an electrolyte solution, its doping state changes, and therefore its conductivity changes. The operation of the OECT is controlled by a voltage applied to the gate electrode (gate voltage, VG) and a voltage applied to the drain electrode (drain voltage, VD), which are referenced relative to the source electrode. The drain voltage induces a current (drain current, ID) proportional to the number of holes or electrons moving in the channel, and therefore detects the doping state of the organic electronic material. The gate voltage controls the injection of ions into the channel, and therefore controls the doping state of the organic electronic material, resulting in a change in ID. The change in ID can be expressed as a normalized response (also referred to herein as "NR"), which can be calculated based on Equation 1:

[0207] NR=(|ID-I0|) / I0

[0208] where I0 is the leakage current obtained after the OECT is incubated with a blank solution (ie, a solution without the presence of target analyte) under the same measurement conditions as the biological sample.

[0209] Taking the biorecognition element as an example, the nanobody selectively captures its target. In this paper, GFP protein and SAR-CoV-2 protein are used as examples. This binding event changes the capacitance of the gate electrode and induces a potential across the gate / electrolyte interface, inhibiting the gate control of the OECT. Therefore, the OECT signal varies depending on the concentration of the analyte in the sample, as low as the femtomolar and attomole range. This strategy of using nanoantibody units opens up new avenues for the design of electronic biosensors and can be further expanded to different sensing platforms and various target proteins.

[0210] Generally speaking, methods of using the disclosed device to test the presence, absence and / or concentration of an analyte in a biological sample include: (i) incubating a gate electrode (functionalized with a biorecognition element as described above) with a blank solution, e.g., a buffer such as PBS, a universal transport medium (UTM), or a viral transport medium (VTM), (ii) placing the gate electrode on top of a channel, wherein the channel is in contact with an electrolyte solution, (iii) applying VG and VD; (iv) measuring a first ID (also referred to herein as a background ID); (v) incubating the gate electrode with the biological sample for a period of time sufficient to allow binding between the analyte and the biorecognition element; (vi) rinsing the gate electrode with a wash buffer; and (vii) measuring a second ID (also referred to herein as a signal ID), wherein the difference between the second ID and the first ID indicates the absence, presence, or concentration of the analyte in the biological sample. The second ID may be greater than the first ID, the same or substantially the same as the first ID, or less than the first ID. Optionally, steps (v) to (vii) are repeated one or more times. In some embodiments, steps (i) to (iv) are optional, and the context ID is provided to the user in other ways (such as by including calibration or standard data in an operator manual).

[0211] Optionally, the method comprises the following steps: before any one of the steps (i) to (vii) described above, such as before step (v) or before step (vi) and after step (v), an electrolyte solution is added to the reservoir. The electrolyte solution added to the reservoir can be any electrolyte solution described above, such as PBS having a pH of about 7.4. The blank solution can be a buffer solution that does not contain the target analyte and is used to establish a baseline leakage current, i.e., I0. In general, a drop of the blank solution is applied to the gate electrode and incubated for about 10 minutes. Generally speaking, the time period for incubating the gate electrode with the blank solution is the same as the incubation time period sufficient to allow the binding between the analyte and the biorecognition element. Optionally, after incubating with the blank solution for about 10 to 15 seconds, the gate electrode is rinsed with a rinse buffer. The rinse buffer can be any electrolyte solution as described above, such as PBS.

[0212] In general, VG and VD are applied to the gate electrode and the drain electrode simultaneously, respectively. In some embodiments, VG is applied to the gate electrode by scanning from a first gate voltage to a second gate voltage with a gate voltage step size, and VD is applied to the drain electrode by scanning from a first drain voltage to a second drain voltage with a drain voltage step size. For example, VG is applied to the gate electrode by scanning from -1V to 1V, from -0.8V to 0.8V, or from -0.6V to 0.6V with a gate voltage step size of 0.05V, 0.1V, 0.2V, or 0.5V, and VD is applied to the drain electrode by scanning from 0V to 1V, from 0 to 0.6V, from 0.5V to -1V, from 0V to -1V, from 0V to -0.8V, or from 0V to -0.6V with a drain voltage step size of 0.05V, 0.1V, 0.2V, or 0.5V. In some embodiments, VG is applied to the gate electrode by scanning from -0.6 V to 0.6 V, from 0.2 V to -0.4 V, or from 0.2 V to -0.6 V with a gate voltage step of 0.1 V or 0.05 V, and VD is applied to the drain electrode by scanning from 0 V to -0.6 V, from 0 V to -0.4 V with a drain voltage step of 0.1 V or 0.05 V. In some embodiments, VG is applied to the gate electrode by scanning from 0 V to 0.9 V or from 0 V to 0.6 V with a gate voltage step of 0.1 V or 0.05 V, and VD is applied to the drain electrode by scanning from 0 V to 0.6 V with a drain voltage step of 0.1 V or 0.05 V. When both VG and VD are applied by scanning, a series of IDs can be collected.

[0213] In some embodiments, a fixed VG is applied to the gate electrode and a fixed VD is applied to the drain electrode. The fixed VG and the fixed VD may be the same or different. For example, a fixed VG in the range of -1V to 1V, -0.6V to 0.6V or -0.1V to 0.1V, such as -0.6V, -0.1V or 0.6V, is applied to the gate electrode, and a channel VD in the range of -1V to 1V, -0.6V to 0.6V, 0V to -0.6V or -0.1V to 0.1V, such as -0.6V, -0.1V or 0.6V, is applied to the drain electrode. In some embodiments, a VG of -0.1V is applied to the gate electrode and a VD of -0.1V is applied to the drain electrode. In some embodiments, a VG of 0.1V is applied to the gate electrode and a VD of 0.1V is applied to the drain electrode. In some embodiments, a VG of -0.1V is applied to the gate electrode and a VD of 0.1V is applied to the drain electrode. In some embodiments, 0.1V of VG is applied to the gate electrode and -0.1V of VD is applied to the drain electrode. In some embodiments, -0.6V of VG is applied to the gate electrode and -0.6V of VD is applied to the drain electrode. In some embodiments, 0.6V of VG is applied to the gate electrode and 0.6V of VD is applied to the drain electrode. In some embodiments, 0.6V of VG is applied to the gate electrode and -0.6V of VD is applied to the drain electrode. In some embodiments, -0.6V of VG is applied to the gate electrode and 0.6V of VD is applied to the drain electrode. In some embodiments, 0.5V of VG is applied to the gate electrode and 0.1V of VD is applied to the drain electrode. In some embodiments, 0.9V of VG is applied to the gate electrode and 0.1V of VD is applied to the drain electrode. When a fixed VG and a fixed VD are applied, a single ID can be collected.

[0214] The biological sample is generally in liquid form and is applied to the gate electrode as a droplet and incubated for a period of time sufficient to allow binding between the analyte and the biorecognition element (e.g., nanobody) immobilized on the gate electrode of the OECT. Generally speaking, the volume of the biological sample sufficient for incubation is small. For example, the volume of the biological sample sufficient for incubation is less than 20 μL, less than 10 μL, or preferably ≤5 μL, and more preferably about 5 μL.

[0215] The biological sample can be a body fluid such as whole blood, plasma, serum, saliva, nasal swab, mucus, sputum (processed or unprocessed), bronchoalveolar lavage (BAL), bronchial washings (BW), cerebrospinal fluid (CSF), and urine.

[0216] In some embodiments, the biological sample is not a body fluid, but a liquid obtained from a solid sample such as a tissue (e.g., biopsy material), stool, rectal swab, nasopharyngeal swab, and throat swab. When the biological sample is not a body fluid, the exemplary method described above can be modified to include an initial step of processing sample t to obtain a sample in liquid form, and then subjecting the sample to steps (i) to (v) described above for methods of using the disclosed device to test the presence, absence, and / or concentration of an analyte in a biological sample. Processing methods for converting a sample (not a fluid) into a liquid form are known. For example, when the sample is a nasopharyngeal swab, the sample is processed by placing the proximal portion of the swab in a buffer to produce a biological sample in liquid form (Lopez et al., Pediatr. Res., 86(5): 651-654 (2019)). Optionally, when sample processing is required, the method further comprises a desalting step before sample processing, wherein a buffer is passed through a desalting column to remove any redox reagent, such as dithiothreitol (DTT), to reduce interference signals caused by non-specific interactions of the redox reagent with the gate electrode surface. Optionally, the desalting step is performed after sample processing and before step (v), wherein the biological sample obtained in liquid form is passed through a desalting column to remove any redox reagent introduced during the sample processing step.

[0217] In some embodiments, the exemplary method described above includes the following steps: before other steps, such as before any one of steps (i) to (v), in particular before step (v), a protease inhibitor is added to the biological sample. The protease inhibitor can prevent or reduce damage to the protein-based recognition layer of the sensor caused by protease activity in the biological or environmental sample.

[0218] The biological sample is hatched with a gate electrode for a period of time sufficient to allow the combination between the analyte in the biological sample and the biorecognition element (e.g., nano antibody) modified on the gate electrode surface. For example, the incubation period is up to 60 minutes, up to 50 minutes, up to 40 minutes, up to 30 minutes, up to 20 minutes, or up to 10 minutes, such as 5, 6, 7, 8, 9, or 10 minutes. For example, the biological sample is hatched with a gate electrode for about 10 minutes to allow the combination between the analyte in the biological sample and the nano antibody modified on the gate electrode surface. Incubation can be performed statically, under shaking, or with up and down pipetting, and in some embodiments, for about 3 to 5 minutes. In some preferred embodiments, incubation is performed with pipetting to promote the combination between the analyte in the biological sample and the biorecognition element (e.g., nano antibody) modified on the gate electrode surface, and thus reduce the time required for incubation. When pipetting is performed during incubation, the biological sample is mixed up and down with a pipette for a period of time in the range of 5 seconds to 1 minute, 5 seconds to 50 seconds, 5 seconds to 40 seconds, 10 seconds to 1 minute, 10 seconds to 50 seconds, 10 seconds to 40 seconds, or 20 seconds to 1 minute, such as about 30 seconds, and the pipetting is repeated at least once or at least twice. For example, during a 10-minute incubation, the biological sample is mixed up and down with a pipette for about 30 seconds, and the pipetting is performed a total of 3 times (i.e., 3×30 second pipettings, or 30 seconds pipetting once every 3 minutes during a 10-minute incubation period).

[0219] After this incubation step (v), the gate electrode is rinsed with a rinse buffer to remove any unbound analyte, for example, the gate electrode can be rinsed for up to 1 min, 30 seconds, 15 seconds, etc., and in some embodiments, the gate electrode can be rinsed by soaking in a rinse buffer (about 15 seconds) and taking it out, and repeating this twice. Then, before the step (vii) of measuring the second ID, the rinsed gate electrode is re-contacted with the electrolyte solution. The rinse buffer can be any electrolyte solution as described above, such as PBS.

[0220] Optionally, the method described above comprises the following step: after step (vii), exposing the gate electrode surface to a glycine solution having a pH of less than 5, less than 4 or less than 3, such as about 2. The acidic glycine solution can destroy the binding between the analyte and the recognition element (e.g., analyte / nanoantibody binding), and thus regenerate the OECT sensor for further sample measurement. For example, after the regeneration step, the ID generated from the OECT is comparable to I0 (i.e., the difference between the ID obtained after regeneration and I0 is less than about 10%).

[0221] Exemplary methods can be used to prepare OECT devices containing biorecognition elements (such as nanobodies), the binding partner of which is any target analyte, such as an antigen from SAR-CoV-2, such as a spike protein. Specific nanobodies for the SARS-CoV-2 spike protein (S protein) or more specifically for the receptor binding domain (RBD) within the S protein are available. Two nanobodies that initially targeted SARS-1RBD have recently been shown to recognize SARS-CoV-2RBD with high affinity. There are other nanobodies designed by human antibody framework engineering, which are specifically developed for SARS-CoV-2 (Wu et al., Fully human single-domain antibodies against SARS-CoV-2.bioRxiv2020). The human ACE2 receptor protein to which the virus binds with high affinity can be used as an alternative recognition module. The subject may be symptomatic or asymptomatic. In some embodiments, the subject has been exposed to the virus, however, the subject may not have known exposure to the virus.

[0222] B. Convection-driven ultrafast detection of biomarkers In some embodiments, the disclosed electrochemical transistor (OECT) is combined with the micro-stirring effect of alternating current thermal flow (ACET) for ultrafast detection of severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) at single molecule to nanomolar levels in complex body fluids.

[0223] By generating directional convective microflows that transport analytes to functionalized sites, an effective method for reducing incubation time while improving sensitivity can be achieved. Alternating current (AC) electrokinetic / electrohydrodynamic forces are used as an easy method for electrode / electrolyte systems to accelerate the transport of biological species to the sensor surface to achieve enhanced immune complex formation. Especially for biologically relevant high-conductivity media, AC electrothermal (ACET) flow becomes the dominant phenomenon among other AC electrokinetic / electrohydrodynamic forces and induces directional long-range convective vortices that can deliver target proteins to the electrode surface. Then, it transports nonspecific proteins to the electrode surface in a tangential direction. This article discloses a method for analyte detection via a combination of AC electrokinetic / electrohydrodynamic forces and an OECT-based immunosensor.

[0224] The method may follow the steps and conditions for using a directly functionalized OECT sensor as described above, comprising: (i) optionally incubating a gate electrode (functionalized with a biorecognition element as described above) with a blank solution, e.g., a buffer such as PBS, a universal transport medium (UTM), or a viral transport medium (VTM), (ii) optionally placing a gate electrode on top of a channel, wherein the channel is in contact with an electrolyte solution, (iii) optionally applying VG and VD; (iv) optionally measuring a first ID; (v) incubating the gate electrode with a biological sample for a period of time sufficient to allow binding between the analyte and the biorecognition element; (vi) rinsing the gate electrode with a wash buffer; and (vii) measuring a second ID, wherein the difference between the second ID and the first ID indicates the absence, presence, or concentration of the analyte in the biological sample. The second ID may be greater than the first ID, the same or substantially the same as the first ID, or less than the first ID. Optionally, steps (v) to (vii) are repeated one or more times. Optionally, the method comprises the step of adding an electrolyte solution to the reservoir before step (i) or before step (ii) and after step (i).

[0225] In addition to these steps, the method for analyte detection using AC electrokinetic / electrohydrodynamic forces further comprises the step of applying an AC potential to the conductive layer, which step is optionally performed before and / or during step (v) and optionally before and / or during step (i). The AC potential applied to the conductive layer (1) is sufficient to produce a micromixing effect (e.g., up to at least 15 μm s -1 The average water velocity of the conductive layer is (1) to reduce the average water velocity of the conductive layer, thereby rapidly transporting the target molecule to the sensor surface and allowing the target molecule (such as the spike protein) to accumulate on the immobilized binding partner (such as the nanobody) at the gate electrode and (2) maintaining the temperature of the electrolyte solution (the change in solution temperature is only a few degrees Celsius, such as less than about 8 degrees Celsius, less than 6 degrees Celsius, or less than 4 degrees Celsius). Generally speaking, the AC potential applied to the conductive layer can be in the range of 1 Vpp to 8 Vpp or 2 Vpp to 6 Vpp, such as about 6 Vpp (pp refers to peak to peak).

[0226] Alternating current electrothermal flow (ACET) is used as a tool to reduce the detection time of transistor-based biosensors. ACET flow causes a micro-stirring effect to rapidly detect single-molecule to nanomolar levels of biomarkers in complex body fluids. ACET flow is induced by an AC electric field applied to the gate electrode, and the flow moves the analyte to the concentric gate electrodes within one minute. The analyte is then captured by the recognition unit on the gate electrode, while non-target species are moved away from the detection area.

[0227] OECTs designed and operated as disclosed herein provide results after less than 5 minutes (such as 2 minutes or 1 minute) of sample exposure (without manual pipetting), which would otherwise take about 1 hour (from sample incubation to result) in the case of passive diffusion, and at least 10 to 15 minutes in the case of manual pipetting. The increase in detection speed does not reduce the detection quality; ACET-integrated OECTs maintain high specificity and single-molecule sensitivity in buffers and biological samples (such as saliva). Compared with the same OECT-based immunosensor without ACET, the ACET-enhanced OECT-based sensor can also achieve higher specificity and lower background due to the electrothermal flow-induced removal of non-specific species from the sensor surface. In addition, the ACET-enhanced OECT-based immunosensor exhibits low energy consumption (i.e., nW level, such as about 100nW). For example, the ACET-enhanced OECT-based sensor can be operated using only 100nW power and about 2 minutes of incubation with 10μL or 5μL sample to detect target analytes in complex media with single-molecule sensitivity.

[0228] Particularly preferred embodiments are exemplified below.

[0229] Examples

[0230] Example 1. Directly functionalized OECTs

[0231] Materials and methods

[0232] Material

[0233] Sodium chloride, Tween-20, glycerol, HEPES, bovine serum albumin (BSA), 1,6-hexanedithiol (HDT) and PBS (pH 7.4) were purchased from Sigma-Aldrich and used as received. All aqueous solutions were prepared with ultrapure water (Millipore Milli-Q). p(g0T2-g6T2) was synthesized as previously reported. Protein purification materials: agar, LB broth, 2xYT broth, kanamycin, glucose, isopropyl β-D-1-thiogalactopyranoside (IPTG), BugBuster (Novagen), cOmplete protease inhibitor cocktail (Sigma), Benzonase (Novagen), hen egg white lysozyme (Fluka), tris(2-carboxyethyl)phosphine (TCEP), tris(hydroxymethyl)aminomethane hydrochloride (Tris-HCl), imidazole, glycerol, dithiothreitol (DTT), ethylenediaminetetraacetic acid (EDTA), D-dethiobiotin, 10K Amicon ultra spin concentrator (Milipore). Purification columns and materials were purchased from GE. Health care: HisTrap-HP 5ml, StrepTrap-HP 5ml, Superdex75 16 / 600. Cysteine-terminated spyTag peptide and MCA-spyTag peptide were synthesized by GenScript Biotech (Singapore). SARS-CoV-2S1 (40591-V08B1) was purchased from Sino Biological. Universal Transport Medium Kit (UTM, proprietary composition) was obtained from NobleBiosciences, Inc. SARS-CoV-2 spike pseudotyped lentivirus (catalog SP101-100) and non-typed lentivirus negative control (catalog SP401-025) were purchased from GeneCopoeia Inc. Untreated wastewater was collected from the KAUST wastewater treatment plant. Saliva was collected from volunteers. Protocols and procedures involving human saliva were approved by KAUST IBEC (approval numbers 18IBEC11 and 20IBEC25) and the National Committee of BioEthics, Saudi Arabia (registration number HAP-02-J-042).

[0234] OECT and gate electrode fabrication

[0235] As previously reported, OECTs were fabricated photolithographically using a parylene C (PaC) lift-off method. A p(g0T2-g6T2) thin film was spin-coated (800 rpm, 45 s) from a chloroform solution (5 g / L) on the substrate to produce a film thickness of about 70 nm in the channel. The second PaC layer was then removed by tape.

[0236] The gate electrode was made using a flexible substrate polyimide (175 μm thickness). A 10 nm thick Cr layer was sputtered on the substrate as an adhesion promoter and a 180 nm thick gold layer. The gold-coated substrate was then cut into sections with defined areas (0.64 or 4 mm 2 ) square geometry. All electrodes were cleaned by ultrasonic treatment in isopropanol, deionized water, and dried in a vacuum oven overnight. The gate electrodes were electrochemically cleaned using cyclic voltammetry (CV) in 10 mM sulfuric acid. -1 20 CVs in the potential range between -0.2 V and 1.2 V were applied at a scan rate of 1.5 Å.

[0237] Biofunctionalization of SpyDirect Gate Electrodes

[0238] First, 0.1 mg / mL cysteine-terminated peptides with spyTag linkers were dissolved in water and applied to the gate electrode for 1 h. The electrode was rinsed thoroughly with water. Secondly, 50 μM green fluorescent protein (GFP) or SARS-CoV-1 nanobody (with spyCatcher) was dissolved in binding buffer (20 mM HEPES, 150 mM NaCl, 0.02% w / v NaN3, 0.05% v / v Tween-20, 0.1% w / v BSA at pH 7.4) and incubated with the peptide-linked electrode for 1 h. Subsequently, the nanobody-functionalized gate electrode was rinsed with binding buffer.

[0239] Electrochemical measurements

[0240] All electrochemical measurements were performed using a potentiostat (Autolab PGstat128N, MetroOhm) in a conventional three-electrode assembly. Platinum wire and Ag / AgCl electrodes were used as counter and reference electrodes, respectively. A gold electrode was used as the working electrode. The electrochemical properties of the gold electrode before and after surface modification with different reagents were studied by electrochemical impedance spectroscopy (EIS) and CV. In the presence of 10 mM [Fe(CN)6] 3- / 4- The measurement was performed in 5 mL of 10 mM PBS solution (pH 7.4). For CV measurement, the potential window of gold was generally determined between -0.2 and 0.6 V, and the scan rate was kept at 10 mV s -1 The impedance spectra were recorded in the frequency range of 0.1 to 100000 Hz at a DC voltage of 0 V and open circuit potential and an AC modulation of 10 mV.

[0241] X-ray photoelectron spectroscopy (XPS) characterization.

[0242] XPS measurements were performed using AMICUS / ESCA (1468.6 eV). The source was operated at 10 kV with a current of 10 mA, generating a power of 100 W. The vacuum level of the analysis chamber was maintained at 10 -7 The acquired spectra were calibrated to the C1s reference at 284.8 eV. The XPS spectra were deconvoluted using Gaussian and Lorentzian methods and background subtracted by the Tougaard method.

[0243] Dissipative Quartz Crystal Microbalance (QCMD) Monitoring

[0244] QCMD measurements were performed using a Q-sense analyzer (QE401, Biolin Scientific) following HDT SAM biofunctionalization or SpyDirect biofunctionalization. Piezoelectric active gold sensors (0.7854 cm 2 ). All solutions were injected into the chamber at a flow rate of 100 μL / min, controlled by a peristaltic pump. After ensuring that the sensor was completely covered by the solution, the pump was stopped for static incubation for a specific period of time. All QCMD data presented in this paper were recorded at the 7th overtone and analyzed using the same method detailed in previous work.

[0245] Atomic force microscopy (AFM)

[0246] AFM scans were obtained with a Veeco Dimension 3100 scanning probe system. Topography scans in electrolyte were performed using a Bruker SCANASYST-FLUID module equipped with a Scanasyst fluidic probe commercialized by Bruker (nominal resonant frequency: 150 kHz, spring constant: 0.7 N m -1 ). During scanning, both the sample and the probe were immersed in 10 mM PBS, pH 7.4 at room temperature. Gwyddion software was used for statistical data and post-processing.

[0247] protein

[0248] Nanobody-spyCatcher fusion protein is designed based on previous studies. The SARS-CoV-2 spike protein was thawed on ice and centrifuged at 14.000rpm for 45min at 4°C to remove potential aggregates (although no aggregation was observed). The dilution series was prepared using SARS-CoV-2S1 from Sino Biologicals as is. An equal amount of dilution of SinoBiologicals storage buffer itself was used as a negative control. Before use, the laboratory-produced protein was desalted into a storage buffer (20mM HEPES, 150mM NaCl, 0.05% v / v Tween-20, 0.02% w / v NaN3 at pH 7.4) without DTT. Protein concentration was assessed by spectrophotometry (Nanodrop, Thermofisher). Protein dilutions were prepared in standard sensor binding buffer (20 mM HEPES, 150 mM NaCl, 0.05% v / v Tween-20, 0.02% w / v NaN3, 0.1% w / v BSA at pH 7.4). For measurements in saliva samples, cOmplete protease inhibitor cocktail (Sigma) with EDTA was added at 4 times the concentration recommended by the manufacturer (yielding twice the concentration in the final 1:1 mixture with saliva), and 0.5% w / v BSA was included (buffer 3.2). For measurements in standard binding buffer, a 4-fold dilution series was prepared in 96-well microplates over 23 steps starting from 320 nM. For measurements in saliva, a target protein dilution series was prepared in an appropriate buffer (standard or saliva binding buffer) starting from 640 nM so that the final concentration was the same after 1:1 mixing with saliva.

[0249] SARS-CoV-2 pseudotyped lentivirus

[0250] According to the supplier, the concentration of SARS-CoV-2 pseudotyped lentivirus was 8.28 × 10 10 copies / ml (determined by RT-qPCR), while the concentration of the negative control was 1.39×10 11 copies / ml (determined by RT-qPCR). The stock samples were thawed on ice and used as is to obtain 1×10 10A dilution series was prepared in binding buffer (20 mM HEPES, 150 mM NaCl, 0.05% v / v Tween-20, 0.02% w / v NaN3 at pH 7.4) with a starting concentration of 10 copies / ml, enriched for saliva measurements to contain 0.5% w / v BSA and enriched to 4 times the manufacturer's recommended concentration of cOmplete protease inhibitor cocktail with EDTA (Roche). UTM and wastewater were diluted.

[0251] OECT sensor characterization and operation

[0252] The transistors were electrically characterized using a Keithley source meter, which was used to apply drain and gate voltages, while the source electrode acted as a common ground in both circuits. All measurements were performed under ambient conditions. A PDMS well was glued on top of the channel and filled with 200 μL of 40 mM phosphate buffer (PB) as an electrolyte. The OECT channel was stabilized with a reference Ag / AgCl gate by repeating the output curve (IV). The gate voltage (V G ) changes between 0.2 and -0.4V (step size 0.05V) when the leakage current (I D ) and drain voltage (V D ) to perform steady-state measurements of the OECT. Sweeping V from 0 to -0.4 V D For each measurement, 5 IVs were obtained and the fifth IV was used to calculate the NR. All Nanobody gates were kept in PB (40 mM, pH 7.4) for at least 10 min to stabilize the sensor before sensing. A blank baseline was obtained before sensing and the readout signal obtained was used as the baseline (g m0 ). The gate electrode functionalized with the nanobody was incubated with 5 μL of the sample solution at room temperature for 10 min (pipetted for 30 s every 3 min). The normalized response (NR) was used to determine the calibration curve according to the following equation:

[0253] NR=|(g mD -g m0 )| / g m0 (4).

[0254] result

[0255] Sensor Overview

[0256] The sensor contains an OECT channel and a gate electrode functionalized with SpyDirect nanobodies ( Figures 1A to 1C The p-type conjugated polymer p(g0T2-g6T2) (whose structure is shown below) was spin-casted on the OECT channel to make the OECT function in the accumulation mode.

[0257]

[0258] In the absence of gate voltage, the OECT is initially in its OFF state. Once a negative bias is applied, anions from the electrolyte are injected into the film and compensate the holes, resulting in an ON state. The high operational stability of the OECT device was demonstrated with a standard Ag / AgCl reference gate electrode. The output characteristics of the OECT were recorded using 10 mM PBS as the electrolyte. At low V D Down, I D The increase is significant, and then at higher V D The saturation state occurs at 0.1 V, which is consistent with accumulation mode OECT operation. As observed from the forward and reverse voltage sweeps (data not shown), the devices exhibit minimal hysteresis with almost identical behavior. The p(g0T2-g6T2) transistor has a low OFF current of approximately 10 μA and a maximum g at saturation state. m The accumulation mode OECT has a low power requirement (V G =-0.05V, V D =75 μW at -0.1 V) (data not shown). The device was switched "ON" and "OFF" for 10 s each and the I D The device retained 98% of its initial current, demonstrating that the device can operate at V G =V D =-0.5V electrical stimulation, the device is highly stable in aqueous media. In addition, the small gate voltage applied to keep the device in its ON state reduces the risk of material instability to meet long-term use requirements, and excellent stability can be achieved under device IV operating conditions. Atomic force microscopy (AFM) was used to confirm the changes in surface roughness and feature height during biofunctionalization (data not shown). Before immobilizing any biomolecules, the root mean square (RMS) roughness of the Au electrode was 4.5nm and the average height of the Au particles was 14.7nm (data not shown). After incubation with a nanobody solution containing BSA, large particles were observed on the gate electrode. . The RMS roughness of the gate electrode modified with nanobody / BSA increased to 6.0nm. And the feature height of these added biomolecule layers increased by 8.4nm (from 14.7 to 23.1nm), confirming the immobilization of nanobodies and BSA.

[0259] The sensor response was normalized according to the transfer characteristics. A blank measurement was obtained before sensing and the obtained readout signal was used as the baseline (I0). The gate electrode functionalized with the nanobody was incubated with 5 μL of the protein target solution for 10 min and then washed thoroughly with the binding buffer to remove unbound protein. The same gate was immersed in the electrolyte on top of the OECT channel to obtain a second transfer curve at this given concentration (I D As shown in these transfer curves, the OECT will input (V G ) is converted to output (I D ) changes significantly. The efficiency of the energy conversion is calculated as the first derivative of the transfer curve and is defined as the transconductance For each sensor, the normalized response (NR) was determined as the normalized change in OECT modulation.

[0260] Nanobody gate design HDT and SpyDirect

[0261] Previous designs of nanobody gates were built on the basis of HDT SAMs, as described in PCT / IB2021 / 055981, in which maleimide-modified SpyTag peptides were chemically coupled to HDT SAMs to form a combined chemical SAM (chem-SAM) on gold, and the nanobody-SpyCatcher fusion protein then attached itself to this chem-SAM by autocatalytic formation of covalent SpyCatcher-SpyTag bonds. The design of SpyDirect biofunctionalization described herein allows sensor surface functionalization to be achieved in two steps: the 17-amino acid spyTag peptide (AHIVMVDAYKPTKGSGC) is directly anchored to the gold gate electrode by forming an Au-S bond via the thiol side chain of the C-terminal cysteine, and then the spyCatcher-nanobody fusion protein ( Figure 1C ).

[0262] To compare the packing density of biomolecules on the surface, both the HDT SAM immobilization step and the SpyDirect immobilization step were monitored by a dissipative quartz crystal microbalance (QCMD) on a piezoelectrically active gold sensor ( Figure 2A , Table 2). The gold QCMD sensors were subjected to either HDT SAM-based functionalization or SpyDirect-based functionalization. During the biofunctionalization step, the cumulative mass gains of (1) peptide-spyTag, (2) nanobody-spyCatcher, and (3) BSA over time were introduced into the system, followed by a wash step. However, the SpyDirect method had a significant difference in the mass gain of the anchor peptide (241 ± 9 vs. 94 ± 32 ng / cm 2) and nanobodies (376±18 and 313±26 ng / cm 2 ) respectively. Based on the molecular weights of maleimide peptide, SpyDirect peptide and SARS-CoV-1 nanobody-SpyCatcher protein of 1.76, 1.78 and 29.05 kDa, the SpyDirect method has the following advantages: a higher packing density of peptides (82×10 12 SpyDirect peptides with 32×10 12 Maleimide peptide / cm 2 ) and the higher packing density of nanobodies (7.8×10 12 With 6.5×10 12 / cm 2 The formation of this unusually high-density biorecognition layer on the surface of the SpyDirect-based nanobody only allowed the absorption of less BSA (43 ± 26 vs. 90 ± 42 ng / cm 2 ), but showed excellent anti-biofouling performance by exposure to non-target proteins, namely green fluorescent protein (GFP) at a concentration of 200 nM. Negligible mass (less than 1 ng / cm 2 However, for the real target protein (SARS-CoV-2 spike), the SpyDirect-based SARS-CoV-1 nanobody surface had a greater binding capacity than the HDT SAM-based surface (659 vs. 290 ng / cm 2 ). Thus, the SpyDirect approach doubled the binding capacity compared to the HDT SAM-based biofunctionalization used in previous studies.

[0263] Table 2. QCMD summary of biofunctionalization and target protein detection.

[0264]

[0265] Atomic force microscopy

[0266] Atomic force microscopy (AFM) confirmed the changes in surface roughness and feature height of the biofunctionalized gate in its wet state (10 mM PBS, pH 7.4) ( Figure 2B ). Compared with bare gold, the gold gate modified with nanobody / BSA increased its root mean square (RMS) roughness by 1.5nm and the average feature height by 8.4nm.

[0267] Biofunctionalization of SpyDirect Gates

[0268] The biofunctionalization of the gate electrode was assessed biophysically. First, CV and EIS were used to monitor the surface modification process by observing the changes in the gate electrode after each assembly step. As shown in CV ( Figure 3A ), the bare Au electrode shows [Fe(CN)6] 3- / 4- Expected reversible peaks for the redox pair, with a reduction peak current of 42 μA, an oxidation peak current of 44 μA, and a peak potential separation of 150 mV. After adding the peptide layer, electron transfer from the redox probe to gold was inhibited, confirming the formation of a continuous peptide layer on the gate electrode. The permeability of ions through the peptide layer is sharply reduced, and the corresponding current response is correspondingly reduced. When the nanobody is immobilized via spyTag-spyCatcher coupling chemistry, the penetration of the redox probe is further reduced, and the current is reduced to almost zero. In addition, EIS is used to characterize these subsequent modifications by drawing representative Bode plots and Nyquist plots of the Au gate electrode after each step ( Figure 3B and 3C ). Using the Randles equivalent circuit ( Figure 3C The equivalent circuit includes the electrolyte resistance (R s ), the double layer capacitance formed at the electrode / electrolyte interface (C dl ), the charge transfer resistance of the electrode (R ct ) and the Weibull impedance generated by the diffusion of ions from the electrolyte to the electrode surface. After anchoring the spyTag peptide to the electrode, the diameter of the semicircle in the high-frequency region of the Nyquist trace increased significantly, indicating increased impedance. Modeling of the Nyquist plot showed that R ct The increase from 0.5 to 7.5 kΩ confirmed the charge blocking behavior of the peptide layer ( Figure 3D After further immobilization of spyCatcher / nanobody (with BSA blocking agent), the impedance increased further and the Nyquist plot became a large semicircle (R ct ≈14.4 kΩ). These measurements show that the biofunctional layer on gold increases the charge transfer resistance and decreases the capacitance, demonstrating successful biofunctionalization at each step. The successful immobilization of the nanobody on the gate surface was also confirmed by XPS ( Figures 3E to 3G ). As subsequent layers of peptide and nanobody are added to the gate surface, the intensity of both the N1s XPS spectrum and the C1s XPS spectrum increases, while the intensity of the Au 4f spectrum decreases due to coverage of the gold surface.

[0269] Detection of SARS-CoV-2 using SpyDirect Nanobody Gate

[0270] To evaluate the performance of the SpyDirect sensor, a well-characterized GFP-recognizing nanobody was used. First, a GFP nanobody-functionalized gate was used to record a baseline after incubation with PB (40 mM, pH 7.4). The same gate was then incubated with increasing concentrations of GFP protein for 10 min and washed thoroughly. The corresponding transfer curves were recorded (data not shown). The current of the transfer curves with increasing GFP concentrations decreased and exhibited a threshold voltage (V th ) shifts significantly towards more negative values. D The trend can be attributed to the specific interaction between GFP nanobody and GFP rather than OECT instability, as the OECT device was stable over a long period of operation (data not shown). In contrast, no significant changes in the transfer characteristics were observed when the GFP nanobody functionalized gate electrode was incubated with a solution containing 20nM to 325nM lysozyme (data not shown). Lysozyme was chosen as a negative target because it is abundant in saliva (salivary lysozyme concentration is 154nM), which is intended to be used as a medium for SARS-CoV-2 applications. In V D =-0.1V, V G = -0.05V to obtain the g of the blank sensor m The maximum value was chosen to determine the sensor calibration curve and maximize sensor performance. The normalized response (NR) was calculated to measure sensor performance. The sensor responded significantly to GFP concentrations as low as 1.2 aM (NR = 20%) and showed a dynamic range spanning 10 orders of magnitude ( Figure 3C ). For these devices, the lowest limit of detection (LOD) was calculated to be 5.5×10 -19 M, corresponds to 1 to 2 GFP molecules in 5 μL sample. The sensor response to non-target protein (lysozyme at 325 nM concentration) was negligible (NR<10%). The specificity of the sensor was independent of the gate voltage (data not shown).

[0271] After validating the sensitivity and specificity of the GFP OECT sensor (in the case of target protein dissolved in buffer), the performance of the sensor was tested in untreated human saliva with the addition of GFP. The NR increased with increasing concentrations of GFP in saliva in a manner similar to the measurements performed with GPF in buffer, especially in the low target concentration range (data not shown). The sensitivity to GFP in saliva (LOD = 7.4 × 10 -19 M) and in buffer (LOD = 5.5 × 10 -19 M), indicating that the complexity of saliva has a negligible effect on the sensitivity and selectivity of the nanobody sensor.

[0272] After demonstrating that the SpyDirect OECT sensor can selectively detect GFP in buffer and saliva with ultrahigh sensitivity (data not shown), the OECT sensor was modified for COVID-19 diagnosis by replacing the GFP nanobody with a nanobody (VHH72) that recognizes the receptor binding domain (RBD) of the spike protein from both SARS-CoV-1 and SARS-CoV-2 with similar affinity. First, a gate functionalized with the SARS-CoV-1 nanobody was used to record the baseline after incubation with phosphate buffer (PB, 40mM, pH7.4). The same gate was then incubated with increasing concentrations of the SARS-CoV-2 spike protein for 10min and washed thoroughly. The corresponding transfer curves were recorded ( Figure 4A The transfer curves for increasing SARS-CoV-2 spike concentrations showed a decrease in current and a threshold voltage (V th ) shifts significantly towards more negative values. D The trend can be attributed to the specific interaction between the SARS-CoV-1 nanobody and the SARS-CoV-2 spike protein, rather than OECT instability, because the OECT device was stable over a long period of operation (data not shown). In contrast, no significant changes in the transfer characteristics were observed when the SARS-CoV-1 nanobody functionalized gate electrode was incubated with a solution containing 20nM to 650nM lysozyme (data not shown). Lysozyme was used as a negative target because it is abundant in saliva (salivary lysozyme concentration is 154nM), which is intended to be used as a medium for SARS-CoV-2 applications. In V D =-0.1V, V G = 0.05V to obtain the blank sensor's g m The maximum value is chosen to determine the sensor calibration curve and maximize sensor performance ( Figure 4B ). m Decreases as more proteins bind, e.g. Figure 4B As shown. The normalized response (NR) was calculated to evaluate the sensor performance. The sensor responded significantly to SARS-CoV-2 spike concentrations as low as 1.2aM (NR = 20%) ( Figure 4C ). This concentration is equivalent to 3 protein particles in 5 μL sample solution, indicating that the OECT sensor can achieve single-molecule sensitivity. In addition to the high sensitivity (LOD = 6 × 10 -22 In addition to SARS-CoV-2, the OECT sensor also achieved a dynamic range of 10 orders of magnitude in this complex biological medium. In contrast, the OECT sensor showed only negligible response (NR < 15%) to non-targets in saliva (20 to 650 nM lysozyme) ( Figure 4C).

[0273] In OECT, considering the gate-electrolyte-channel structure ( Figure 4D ) in the OECT: the gate / electrolyte interface and the channel / electrolyte interface. According to the Bernard model, the OECT operation involves a purely capacitive process of the ionic circuit. Due to the insulating nature of these protein target molecules, the capacitance of the gate electrode decreases during the binding event (data not shown). Depending on the magnitude of the capacitance at each surface, the applied gate voltage is distributed at the two interfaces. The voltage drop at the gate electrode and the channel is described by the following equation:

[0274]

[0275] Among them, V CH is the voltage drop at the channel surface, j is a complex number, ω is the angular frequency, C CH is the channel capacitance, C G is the gate capacitance, and R E is the resistance of the electrolyte. Under DC conditions, the above equation can be simplified to:

[0276]

[0277] Therefore, the corresponding voltage drop at the gate electrode V G It can be described as:

[0278] V G =1-V CH (3)

[0279] When the target protein molecule binds to the gate electrode, C G Reduce to C' G,eff , which results in the applied V G A larger drop in the gate / electrolyte interface ( Figure 4E ). Therefore, the Nanobody-protein binding at the gate electrode leads to a less pronounced doping capability and therefore to a less efficient transistor operation, and a reduced leakage current is observed.

[0280] The high sensitivity and selectivity were further demonstrated by randomly detecting SARS-CoV-2 spikes in saliva at concentrations of 1.2 fM and 1.2 pM, respectively. Figure 5A). SpyDirect gates functionalized with either SARS-CoV-1 nanobody or GFP nanobody were incubated with 5 μL of SARS-CoV-2 spikes in saliva, and the response of the SARS-CoV-1 nanobody gate to SARS-CoV-2 spikes had a larger NR than the response of the GFP nanobody gate. The SpyDirect nanobody gate was further challenged to detect SARS-CoV-2 spikes from untreated wastewater. Random concentrations of 1.2 pM and 1.2 nM of SARS-CoV-2 spikes in wastewater were distinguished by the SpyDirect SARS-CoV-1 nanobody gate using the GFP nanobody gate as a control ( Figure 5B ). SpyDirect SARS-CoV-1 Nanobody Gates Show Improved Long-Term Stability ( Figure 5C and 5D ). Compared to freshly prepared SARS-CoV-1 nanobody gates, the response of the HDT SAM-based SARS-CoV-1 nanobody gates to 1.2pM SARS-CoV-2 spike protein in saliva dropped to 92% response after 3 days of storage in PBS buffer; after 7 days of storage, these gates could no longer distinguish SARS-CoV-2 spike from lysozyme due to similar NRs. In contrast, the SpyDirect gates maintained a 67% response after 7 days of storage and were still able to discriminate spike protein at a concentration of 1.2pM (NR=20%) from lysozyme (NR<12%). In addition, the simplified biofunctionalization of the SpyDirect OECT sensor also improves the ease of use and storage of the sensor, because the peptide-coated gate electrode can be stored under ambient conditions for a specific period of time without losing its functionality before nanobody immobilization. The SpyDirect OECT SARS-CoV-1 nanobody sensor was fully functionalized and effectively detected the SARS-CoV-2 spike protein target after seven days of storage after SpyDirect peptide immobilization ( Figure 5E ). Compared with the freshly prepared sensor, similar NR was obtained.

[0281] The HDT SAM-based OECT sensor in the previous report achieved an LOD of 1.3 × 10 -21 In contrast, after exposing the SARS-CoV-1 nanobody gate to raw saliva, universal transport medium (UTM), and untreated wastewater, the SpyDirect OECT sensor showed lower background noise levels ( Fig. 6AThe lower background noise level in raw saliva makes the SpyDirectOECT sensor more sensitive than its HDT SAM sensor in saliva (LOD of 6 × 10 -22 With 1.3×10 -21 M).

[0282] Finally, as a proxy for viral load in patient samples, the SpyDirect sensor was tested against SARS-CoV-2 recombinant lentivirus pseudotyped with the spike protein in human saliva, UTM, and untreated wastewater. The OECT sensor showed a specific response to the SARS-CoV-2 pseudotyped lentivirus but not to the negative control in saliva ( Figure 6B ). The NR increased with the copy number of SARS-CoV-2 pseudotyped lentivirus, but showed negligible changes when the same gate electrode was exposed to an equivalent copy number of virus particles that did not display the SARS-CoV-2 spike protein. The sensor can detect 25 copies of virus particles in a 5μL sample volume (average NR change of 28±6%), which is greater than the background noise from the negative control (16%). The LOD of the sensor for SARS-CoV-2 pseudotyped lentivirus in saliva was 200 copies / mL. In addition, the SpyDirect-SARS-CoV-1 Nanobody Sensor showed excellent detection capabilities in response to SARS-CoV-2 pseudotyped lentivirus in both UTM and wastewater. Using the GFP nanobody gate as a control, the SpyDirect SARS-CoV-1 Nanobody Sensor can reliably detect as low as 10 copies ( Figure 6C right Fig.6D , HDT SAM sensor 1000 copies) and as low as 500 copies in wastewater ( Fig. 6E right Fig. 6F , HDT SAM sensor 50,000 copies) of SARS-CoV-2 pseudotyped lentivirus.

[0283] The excellent sensor performance of the SpyDirect gate can be attributed to its excellent electrochemical and chemical stability. Compared with the fresh gate (237k ohm), the average Rct of the SpyDirect nanobody gate showed negligible change after 3 days of storage in PBS (242k ohm), and a 12.5% ​​decrease after 7 days of storage ( Fig. 7A ). In contrast, the HDT SAM-based nanobody gates showed a dramatic increase in mean Rct values ​​after storage (to 26-fold after 3 days of storage and to 64-fold after 7 days of storage, respectively) ( Figure 7BXPS further confirmed the negligible changes in the chemical structures of C1s, O1s, and N1s of the SpyDirect-based nanobody gate after long-term storage ( Figures 7C to 7E However, for the HDT SAM-based nanobody surface, a shift in the XPS spectra of C1s, O1s, and N1s was observed after 1 week of storage in PBS ( Figures 7F to 7H ).

[0284] The poor chemical stability of HDT SAM may be mainly due to the instability of the nanobody surface based on HDT SAM. As demonstrated in previous studies, the Bio-SAM layer (peptide-spyTag / SpyCatcher-nanobody) has a high flexibility ranging from 12 to 20nm in a wet state using the autocatalytic attachment of spyCatcher nanobody fusion protein. Once HDT SAM is attacked, the Bio-SAM layer will collapse and randomly cover the gold electrode to form an additional insulating layer. Due to the loss of flexibility, this further hinders the diffusion of redox probes to the electrode surface, and a sharp increase in Rct is observed. Due to the passage of time, new surfaces are exposed, and the transformation of XPS energy spectrum occurs.

[0285] discuss

[0286] Conventional biofunctionalization methods use thiol SAMs as linkers to build immunosensors. However, this strategy results in low sensor stability when used under ambient or biological conditions. SpyDirect, described in this article, is a simplified biofunctionalization approach for OECT immunosensors. SpyDirect uses cysteine-terminated spyTag peptides to directly attach nanobodies to the gold surface of the gate electrode. The self-catalytic spyCatcher-nanobody module is designed to improve nanobody orientation and stacking. This strategy produces sensors that display flexible ultra-high-density arrays of highly efficient and directional nanobody bioreceptors. Due to these features, the SpyDirect OECT sensor can detect less than 200 copies of SARS-CoV-2 pseudotyped lentivirus from raw saliva, UTM, and untreated wastewater in less than 15 min. Thus, the sensor achieves 6×10 -22 The results show that SpyDirect biofunctionalization can provide a low LOD of 10 M and cover a dynamic range of more than 10 orders of magnitude. Compared with conventional biofunctionalization methods using thiol SAMs, SpyDirect biofunctionalization exhibits lower background noise, improved long-term stability, and easier manufacture, storage, and use. Therefore, SpyDirect biofunctionalization represents an improved approach for the next generation of sensor devices (including but not limited to immunosensors).

[0287] Example 2. Ultrafast detection driven by convection

[0288] Materials and methods

[0289] Material

[0290] Sodium chloride, Tween-20, glycerol, HEPES, bovine serum albumin (BSA), (3-glycidyloxypropyl)trimethoxysilane (GOPS), ethylene glycol (EG), 1,6-hexanedithiol (HDT), 3-mercaptopropionic acid (MPA), and PBS (pH 7.4) were purchased from Sigma Aldrich and used without treatment. Solutions were prepared with ultrapure water (Millipore Milli-Q). p(C6NDI-T) was synthesized according to a previously reported procedure. The synthetic route of p(g3C2T2-T) is provided below. The materials used for protein purification are as follows: agar, LB broth, 2xYT broth, kanamycin, glucose, isopropyl β-d-1-thiogalactopyranoside (IPTG), BugBuster (Novagen), cOmplete protease inhibitor cocktail (Sigma), Benzonase (Novagen), hen egg lysozyme (Fluka), tris(2-carboxyethyl)phosphine (TCEP), tris(hydroxymethyl)aminomethane hydrochloride (Tris-HCl), imidazole, glycerol, dithiothreitol (DTT), ethylenediaminetetraacetic acid (EDTA), d-dethiobiotin, 10K Amicon ultra spin concentrator (Milipore). Viral target proteins were purchased from Sino Biiological: SARS-CoV-1RBD (40150-V08B2), SARS-CoV-2RBD (40592-V08H), and SARS-CoV-2S1 (40591-V08B1).

[0291] Synthesis and characterization of p(g3C2T2-T)

[0292] Scheme 1. Synthetic route for producing p(g3C2T2-T)

[0293]

[0294] ((((([2,2'-bithiophene]-3,3'-diylbis(oxy))bis(ethane-2,1-diyl))bis(oxy))bis(ethane-2,1-diyl))bis(oxy))bis(ethane-2,1-diyl)bis(4-methylbenzenesulfonate)(1)

[0295] 2,2'-((((([2,2'-bithiophene]-3,3'-diylbis(oxy))bis(ethane-2,1-diyl))bis(oxy))bis(ethane-2,1-diyl))bis(oxy))bis(ethane-1-ol) 1 (11.63 g, 25.1 mmol, 1.00 eq.) was dissolved in 30 mL of pyridine and cooled to 0°C. 4-Toluenesulfonyl chloride (11.51 g, 5.25 mmol, 2.4 eq.) was added in portions. The solution was stirred for 30 minutes and then placed in a refrigerator overnight. Upon completion, 200 mL of water was added, followed by 100 mL of 2M HCl, and the aqueous phase was extracted three times with dichloromethane. The combined organic layers were then washed with water and then dried over sodium sulfate. Excess solvent was removed under reduced pressure to give the crude product (17.81 g, 1.30 mmol, 92% yield) as a colorless oil, which was used without further purification. 1H NMR (400 MHz, CDCl3) δ 7.82-7.72 (m, 4H), 7.37-7.27 (m, 4H), 7.06 (d, J = 5.6 Hz, 2H), 6.83 (d, J = 5.5 Hz, 2H), 4.27-4.15 (m, 4H), 4.17-4.06 (m, 8H), 3.77-3.55 (m, 12H), 2.40 (s, 6H).

[0296] 3,3'-Bis(2-(2-(2-ethoxyethoxy)ethoxy)ethoxy)-2,2'-bithiophene(2)

[0297] Sodium hydride (3.36g, 84mmol, 10.50eq.) is dissolved in 200mL anhydrous tetrahydrofuran and cooled to 0°C. Then anhydrous ethanol (4.67mL, 79.9mmol, 10.00eq.) is added dropwise. After that, the reaction mixture is stirred at 65°C for 60 minutes. Then the reaction mixture is cooled to 0°C, followed by dropwise addition of a solution of compound 1 (6.17g, 8.00mmol, 160mmol, 1.00eq.) in 100mL tetrahydrofuran. The mixture is stirred overnight at 55°C. Water is added and the aqueous phase is extracted three times with dichloromethane. The combined organic layer is washed with water and brine, and then dried over sodium sulfate. Excess solvent is removed under reduced pressure. The final product (1.70g, 3.28mmol, 41% yield) is obtained as an off-white solid by silica column chromatography eluting with ethyl acetate. 1H NMR (400MHz, CDCl3) δ7.06 (dd, J=5.6, 0.9Hz, 2H), 6.84 (dd, J=5.6, 0.9Hz, 2H), 4.27-4.20 (m, 4H), 3.89 (d d,J=5.6,4.5Hz,4H),3.78-3.70(m,4H),3.69-3.52(m,12H),3.59-3.43(m,4H),1.19(t,J=6.9Hz,6H).13C NMR (101MHz, CDCl3) δ151.85,122.01,116.69,114.90,71.50,71.05,70.83,70.81,70.79,70.72,70.13,69.94,66.73,15.27.

[0298] 5,5'-Dibromo-3,3'-bis(2-(2-(2-ethoxyethoxy)ethoxy)ethoxy)-2,2'-bithiophene(3)

[0299] Compound 2 (0.86 g, 1.66 mmol, 1.00 eq.) was dissolved in 100 mL of anhydrous tetrahydrofuran. The resulting solution was degassed for 15 minutes and cooled to -20 ° C. Then N-bromosuccinimide (0.62 g, 3.48 mmol, 2.10 eq.) was added in portions to avoid light. The progress of the reaction was monitored by thin layer chromatography using ethyl acetate as eluent. Once the reaction was complete, after about 60 minutes, the reaction mixture was poured into a saturated sodium bicarbonate aqueous solution. The aqueous phase was extracted three times with ethyl acetate, and the combined organic layer was washed with water and brine, and then dried over sodium sulfate. Excess solvent was removed under reduced pressure. The final product (0.92 g, 1.36 mmol, 82% yield) was obtained as a white solid after column chromatography on silica using ethyl acetate as eluent. 1H NMR (400MHz, CDCl3) δ6.86 (s, 2H), 3.86 (dd, J = 5.7, 4.0Hz, 4H), 3.77-3.62 (m, 16 H),3.59(dd,J=6.1,3.6Hz,4H),3.52(q,J=7.0Hz,4H),1.20(t,J=7.1Hz,6H).13C NMR (101MHz, CDCl3) δ151.87,122.03,116.71,114.91,72.09,71.52,71.06,70.85,70.76,70.71,70.16,59.16.

[0300] Poly-3,3'-bis(2-ethoxyethoxy)-5,5"-dimethyl-2,2':5',2"-terthiophene (p(g3C2T2-T))

[0301] Compound 3 (180 mg, 0.266 mmol, 1.00 eq.), 2,5-bis(trimethylstannyl)thiophene (109 mg, 0.267 mmol, 1.00 eq.) and tetrakis(triphenylphosphine)palladium(0) (6.2 mg, 0.053 mmol, 0.020 eq.) were added to a 10 mL microwave vial, followed by 1.8 mL of anhydrous dimethylformamide and 1.8 mL of anhydrous chlorobenzene. The vial was sealed, and the solution was degassed for 10 minutes, then heated to 111 ° C overnight. After cooling to room temperature, the reaction mixture was precipitated into methanol. The crude product was purified by sequential Soxhlet extraction in hexane, acetone, ethyl acetate, and methanol and finally dissolved in chloroform. The solvent was removed from the chloroform fraction under reduced pressure, and the product was reprecipitated into methanol. The final product was collected by suction filtration and recovered as a dark blue solid (119 mg, yield 71%). GPC (DMF, 40°C) Mn = 174 kDa, 1H NMR (500MHz, CDCl3) δ7.08(s,2H),6.97(s,2H),4.37-4.31(m,4H),4.00-3.94(m,4H),3.81(dd,J=5.8,3.7Hz,4H),3.73(dd , J=6.0, 3.6Hz, 4H), 3.66 (dd, J=5.9, 3.8Hz, 4H), 3.57 (dd, J=5.9, 3.8Hz, 4H), 3.50 (q, J=7.0Hz, 4H), 1.19 (t, J=7.0Hz, 6H).

[0302] Fabrication of OECTs and gate electrodes

[0303] OECTs were fabricated on glass substrates using standard photolithography protocols. First, a layer of photoresist (LOR 5B and S1813) was spin-coated on the substrate, which was then exposed to UV light to form a pattern in the shape of the transistor components. After treating the substrate with a developer (MF-319), 10 nm of chromium (Cr) and 100 nm of gold (Au) were sputtered, and a standard stripping was performed in N-methyl-2-pyrrolidone (NMP) at 80°C. Two layers of parylene separated by an adhesive were then deposited, and a second photolithography step was performed using AZ 10XT as a photoresist to insulate the interconnects and the source and drain contacts. After reactive ion etching, the film was spin-cast. A second sacrificial layer of parylene C was stripped off to form a pattern of the OECT channel with a length of 10 μm and a width of 100 μm. Concentric gate electrodes were also fabricated on glass substrates using standard photolithography. A first layer of photoresist (AZ2020) was spin-coated and exposed to UV light using a contact aligner. The photoresist pattern was formed by AZ726 developer and the residual photoresist was removed by oxygen plasma. 10nm Cr and 100nm Au were deposited by sputtering and then stripped in hot NMP. The area of ​​the concentric gate electrode (25mm) was determined by considering the capacitive coupling requirement. 2 ).

[0304] Biofunctionalization of gate electrodes

[0305] Prior to biochemical functionalization, the gold grid electrodes were cleaned in an acidic solution by using cyclic voltammetry. The electrodes were connected in a three-electrode assembly and immersed in a 10 × 10 -3 m H2SO4 solution, with a charge of 100 mV s -1 A scan rate of 25 cycles from -0.2 to 1.5 V vs. Ag / AgCl was applied. -3 A solution of HDT-SAM was prepared in 100% ethanol of mHDT. The gold electrode was immersed in this solution for one hour, after which it was rinsed with ethanol and dried under nitrogen. The electrode was then incubated in PBS (0.1 mg mL-1 ) was incubated with a synthetic maleimide-modified SpyTag peptide solution for one hour. After rinsing the electrodes with PBS, the electrodes were exposed to a concentration of 20×10-3 m HEPES, 150×10-3 m NaCl, 0.05% v / v Tween-20, 0.02% w / v NaN3, 0.1% w / v BSA in binding buffer (pH 7.4). - 6 m of SpyCatcher / Ty1-Nanobody fusion protein for 1 h, and then the electrode was rinsed again with PBS. The functionalized gate electrodes were used on the day of their biofunctionalization.

[0306] MST

[0307] Recombinant nanobody Ty1 and GFP were labeled using the second generation NHS red labeling kit in BSA-free buffer 2.1. Excess dye was removed by size exclusion chromatography on an agarose 6Increase column. Fractions were pooled, concentrated, and stored in BSA-free buffer 2.1. Commercial SARS-Cov1 RBD, SARS-Cov2 RBD, and laboratory-prepared GFP fluorescent protein were unlabeled. -9 A 2× serial dilution of unlabeled protein was analyzed in 16 capillaries at a starting and ending concentration of 2.5×10 -9 Labeled Nanobodies were used at a final concentration of 100 m. The binding reactions were incubated at room temperature for 30 min before analysis. MST analysis was performed on a Monolith NT.115pico instrument (Nano Temper Technologies, Germany) at 20% LED power and HIGH MST power. Three replicates were analyzed for each binding pair.

[0308] KPFM

[0309] KPFM measurements were performed with a Dimension Icon SPM (Bruker) using a SCM-PIT-V2 tip (Bruker). The surface potential reflected by the contact potential difference (VCPD) represents the difference in the work function of the sample and the KPFM tip, as shown in the equation below, and can provide insight into the band structure of the surface:

[0310] V CPD =(Φ 针尖 -Φ 样品 ) / (-e)

[0311] where Φtip and Φsample are the work functions of the tip and sample, and e is the electron charge.

[0312] XPS

[0313] XPS spectra were obtained using a Kratos AXIS Supra instrument equipped with a monochromatic Al Kα X-ray source (1468.6 eV). The source was operated at 75 W under ultrahigh vacuum conditions (about 10 to 9 mbar). The spectra were recorded in hybrid mode using electrostatic and magnetic lenses and 300×700 μm aperture slits. High-resolution spectra were obtained at a fixed analyzer pass energy of 20 eV. The sample was mounted in floating mode to avoid differential charging. The spectrum was calibrated to the C1s reference at 284.8 eV. Background subtraction was performed using the Tougaard method, and deconvolution was performed using the Gaussian and Lorentz methods via XPSPEAK41 software.

[0314] SIMS

[0315] To characterize the elemental composition of the films, secondary ion mass spectrometry (SIMS) studies were performed. P-type and n-type polymers were spin-coated onto gold electrodes. The films were immersed in a 100×10 -3 m NaCl solution to achieve passive swelling. A second set of films of each polymer type was immersed in the electrolyte and biased with a doping voltage (|0.5| V vs. Ag / AgCl) for 5 min. The films were quickly removed and inserted into the SIMS assembly for analysis. Depth profiling experiments were performed on a dynamic SIMS instrument from Hiden Analytical (Warrington, UK), which operates under ultra-high vacuum conditions (typically 10 to 9 Torr). A continuous Ar+ beam with an energy of 4 keV was used to sputter the surface, while selected ions were sequentially collected using a MAXIM spectrometer equipped with a quadrupole analyzer. The raster over the sputtered area was estimated to be 750×750 μm 2 In order to avoid edge effects in depth profiling experiments, it is necessary to acquire data from a small area located in the middle of the eroded area. Using appropriate electronic gating, the acquisition area from which depth profiling data are obtained is approximately 75 × 75 μm 2 .

[0316] QCM-D

[0317] QCM-D measurements were performed using a Q-Sense analyzer (QE401, Biolin Scientific). The piezoelectric active gold sensor (0.7854 cm2) was pre-coated with a monolayer of 1,6-hexanedithiol (HDT) and then mounted in the QCM-D assembly. After the frequency (f) and dissipation (D) changes of the crystal in PBS stabilized, a SpyTag peptide solution (0.1 mg mL in PBS) was added. -1 ) at 100 μL min -1The flow rate was pumped into the fluid chamber at 1.5 ℃. After the solution completely covered the sensor surface, the pump was stopped and the sensor was covered with the peptide solution. Then, the sensor surface was rinsed with PBS flushed with fluid for 15 min to remove excess peptide from the surface. The same procedure was performed to expose the surface of quartz to SpyCatcher / Ty1-nanobody fusion protein (20×10 - 6 m, in binding buffer). The seventh overtone was selected to present the QCM-D data. The mass change (Δm) experienced by the sensor was calculated using the Sauerbrey equation:

[0318] Δm=[(-17.7) / n]Δf n

[0319] Where n is the number of overtones, and Δf n is the change in QCM-D frequency at that chosen overtone.

[0320] EIS

[0321] EIS was used to examine the formation of the biochemical layer on the gold electrode. The three-electrode assembly was connected to a potentiostat (AutolabPGstat128N with Nova software, MetroOhm). A platinum wire and an Ag / AgCl electrode were used as the counter electrode and reference electrode, respectively. The gold grid electrode (working electrode) was immersed in a 10×10 -3 m [Fe(CN)6]3- / 4- in 5 mL PBS (pH 7.4). Impedance spectra were measured at zero DC offset and open circuit potential, and 10 mV AC modulation was applied in the frequency range of 0.1 to 100,000 Hz. The energy spectrum was analyzed using Nova software and appropriate equivalent circuit modeling.

[0322] Proteins and peptides

[0323] Nanobody-SpyCatcher fusion proteins were designed based on the following: available structures (Ty1 nanobody: 6ZXN; SpyCatcher: PDB 4MLI), while placing the nanobody at the N-terminus of the fusion protein to orient the common Ty1 target binding interface away from the sensing surface toward the bulk solution. The protein sequence was reverse transferred and codon optimized with an internal Python script based on DNAChise for expression in E. coli. The plasmids used for protein expression were gene synthesized by TwistBioscience (USA) in a custom expression vector pJE411c with kanamycin resistance and modified with an RBS insulator (BCD2) box to achieve improved translation initiation. The plasmids were transformed into E. coli BL21 (DE3) and the starter culture was inoculated overnight from a single colony. 1 L production culture was grown in a 50 mg L-1 The cells were inoculated 1:100 in 2xYT medium containing kanamycin and 1% glucose and grown at 37°C and 250 rpm to an OD600 of 0.8. -3 m IPTG and incubated at 25°C with shaking for 18 h. Cells were harvested by centrifugation at 6000 g for 10 min at 4°C, washed once with cold PBS, and resuspended in lysis buffer [25×10-3 m Tris-HCl (pH 7.4), 500×10 -3 m NaCl, 10×10 -3 m imidazole, 10% glycerol, cOmplete protease inhibitor (Roche), 25U mL -1 Benzonase HC(Milipore), 2×10 - 3 m DTT] and homogenized using a cell disruptor (Constant Systems, UK). Early purification of GFP nanobodies and msfGFP used ultrasonic lysis. Lysates were clarified by centrifugation at 87,000 g for 45 min, supernatants were filtered through Miracloth filters (Milipore), and purified using StrepTrap HP or HisTrap HP columns (GE Healthcare). Affinity chromatography was performed on FPLC (GE Healthcare), depending on the purification tag. Strep tag binding buffer was 100 × 10 -3 m Tris-HCl (pH 8), 150 × 10 -3 m NaCl, 1×10 -3 m EDTA, 5% glycerol, 0.5×10 -3 m TCEP and 2.5 × 10 -3 Perform elution with 25 × 10 mM desthiobiotin. His-tag binding buffer is pH 7.4 - 3 m Tris-HCl, 500×10 -3 m NaCl, 10×10 -3 m imidazole, 10% glycerol, 2×10 -3 m DTT, and elution was performed with a four-step imidazole gradient up to 0.5 m. Fractions were pooled and concentrated using a 10K Amicon ultra (Milipore) before gel filtration on a Superdex75 16 / 600 column (GE Healthcare) to 20 × 10 -3 m HEPES, 300×10 -3m NaCl, 10% glycerol, 50×10 -6 m EDTA. After rotary concentration, aliquots were snap frozen in liquid nitrogen and stored at -80 ° C. Protein purity, quality and accurate molar mass were monitored by SDS-PAGE and SEC-MALS (size exclusion chromatography multi-angle light scattering) on ​​Dawn HeleosII and OptiLab T-rEx (Wyatt, USA). Protein concentration was determined by absorbance at 280 nm on a Nanodrop spectrophotometer using sequence-specific extinction coefficients.

[0324] Fluorescence microscopy

[0325] Imaging was performed on a DMI8 inverted fluorescence microscope (Leica Microsystems) combined with a pE-4000 fluorescence illumination system (CoolLED), and images were processed using ImageJ software. After ACET-promoted incubation, the presence of proteins was investigated by monitoring the binding of GFP and mCherry to the GFP nanobody gate electrode solution. Green and red emission light from the electrode surface was monitored.

[0326] Protein diluent

[0327] All proteins were thawed on ice and centrifuged at 15,000 rpm for 30 to 45 min at 4°C to remove any potential aggregates (although no aggregation was observed). -3 Protein dilutions were prepared in 1% 10-3 m HEPES, 150×10-3 m NaCl, 0.05% v / v Tween-20, 0.02% w / v NaN3, 0.1% w / v BSA. Saliva incorporation was measured using a modified lysis buffer (20×10-3 m HEPES, pH 7.4). -3 m HEPES, 500×10 -3 m NaCl, 1% v / v Triton-X, 0.02% w / v NaN3, 0.1% w / v BSA) supplemented with cOmplete protease inhibitor cocktail with EDTA (Sigma) at four times the manufacturer's recommended concentration. Dilutions in Sino Biologicals storage buffer were verified to not produce any sensor response on their own. Weak background sensor signals were recorded from dilutions in DTT. Therefore, in-house proteins were stored or exchanged into DTT-free buffer prior to use. Higher concentration proteins from in-house production were first diluted to an intermediate concentration that could still be verified and calibrated by spectrophotometry. -9m, a four-fold dilution series was prepared in 96-well microplates over 23 steps. For measurements in saliva, the diluent was mixed with saliva at 3:1 (3 volumes of protein diluent, 1 volume of saliva) before measurement. As part of the registered program approved by the Institutional Biosafety and Bioethics Committee (IBEC) of King Abdullah University of Science and Technology (KAUST) (project numbers 18IBEC11 and 20IBEC25), saliva samples were collected by healthy volunteers in the morning before eating or brushing their teeth. All volunteers provided signed consent to participate in the study.

[0328] OECT characterization and sensor operation

[0329] The steady-state characteristics of the transistor were recorded using a Keithley 2602A source meter unit operated by custom LabVIEW software. When the source electrode was common ground, a drain voltage (VD) and a gate voltage (VG) were applied. A PDMS hole (1 cm in diameter, 2 mm thick) was placed on the top of the OECT and filled with 100 μL of PBS (pH 7.4,, ionic strength 0.162 m). The p(g3C2T2-T)OECT was operated by: VG was varied between 0.2 and -0.6 V, while VD was scanned from 0 to -0.6 V. For p(C6NDI-T)OECT, VD was scanned from 0 to 0.6 V while VG was varied from 0 to 0.9 V. ID and gate current (IG) were monitored simultaneously. After the functionalization step, the reference (blank) response of the sensor was obtained by immersing the gate electrode in PBS. The same electrode was then incubated with sample (10 μL) under ACET flow for 2 min, rinsed in buffer 2.1 and twice in PBS, and mounted on top of the channel in a parallel manner to complete the OECT biosensor. To quantify the sensor response and minimize device-to-device variation, the normalized response (NR) of the OECT was calculated by normalizing the protein-induced ID change at a single VD and VG to its value measured after exposure to a blank solution (I0):

[0330] NR=|ID-I0| / I0

[0331] where ID is the current response of the sensor to the analyte solution to which the gate is exposed. After each sensing measurement, the response of the channel gated with Ag / AgCl was recorded.

[0332] Results and discussion

[0333] Design of Nanoantibody-OECT Protein Sensor

[0334] OECTs have two main components: an organic semiconductor thin film in the channel, i.e., p-type p(g3C2T2-T) or n-type p(C6NDI-T) (100 μm in width, 10 μm in length, and ≈85 nm in thickness), and a nanobody-functionalized gold gate electrode (25 nm in diameter). 2 )( Fig. 8A ). The structure of p-type p(g3C2T2-T) or n-type p(C6NDI-T) is shown below.

[0335]

[0336] Both polymers are mixed ionic and electronic charge conductors. After electrochemical doping, their films contain trace amounts of their respective dopant ions throughout their bulk (data not shown), and the capacitance scales with film thickness (data not shown), as revealed using ex situ secondary ion mass spectroscopy (SIMS). For the p(g3C2T2-T) channel, a negative gate voltage (VG) pushes anions into the film, which compensates for holes injected from the metal contact, turning on the device with a threshold voltage (Vth) of -0.13 V (data not shown). The gate potential is modulated by The maximum transconductance, quantifying the maximum change in drain current (ID), is -60 mS at VG = -0.35 V in saturation (VD = -0.35 V). For p(C6NDI-T) OECTs, positive VG drives the cation electron coupling in the channel, switching the current above Vth of 0.27 V with gm of ≈15 μS (VG = 0.5 V, VD = 0.5 V) (data not shown). Under these operating conditions, both OECTs have an on / off ratio of ≈515.

[0337] The operational stability of these channels was evaluated by monitoring the ID for 1 h during application of a square wave shared voltage pulse of magnitude 0.5 V and a duty cycle of 10 s at the gate electrode. These devices were the most stable among identical devices made from a range of other organic semiconductors commonly used in OECTs, with p(g3C2T2-T) retaining 99% of its original channel current after 1 h, while p(C6NDI-T) retained 95% (data not shown).

[0338] Figure 1BThe bioconstruct on the gate electrode is schematically shown. The biorecognition unit of the sensor is the recognition domain of a single-chain antibody (called a nanobody). The nanobody is recombinantly produced as a fusion protein connected to the SpyCatcher domain via a flexible linker. The SpyCatcher domain autocatalytically covalently binds to a commercially synthesized SpyTag peptide, which is chemically immobilized and assembled on a 1,6-hexanedithiol (HDT)-based monolayer on top of a gold electrode. This indirect functionalization strategy enables the immobilization of nanobodies on the surface of electronic devices via HDT. Alpaca-derived nanobody Ty1 was used, which specifically recognizes the receptor binding domain (RBD) of the SARS-CoV-2 spike glycoprotein. The affinity and specificity of this nanobody were evaluated using microscale thermophoresis (MST) analysis, where it showed a strong affinity for SARS-CoV-2 RBD (KD = 24.9 ± 10.3 × 10 -9 m), and minimal interaction with SARS-CoV-1 RBD or green fluorescent protein (GFP) (data not shown).

[0339] To characterize the gold electrode surface and monitor its functionalization with HDT, SpyTag peptide, and nanobody-spyCatcher fusion, a combination of techniques, namely electrochemical impedance spectroscopy (EIS), quartz crystal microbalance and dissipation monitoring (QCMD-D), Kelvin probe atomic force microscopy (KPFM), and X-ray photoelectron spectroscopy (XPS) ( Figures 9A to 9F ). Figures 9A to 9C The representative Nyquist diagram of the gold grid electrode after each functionalization step is presented. As each (bio) chemical layer is formed on the surface, the semicircle of the Nyquist trace becomes larger. Using equivalent circuit fitting, when the nano antibody layer covers the gold electrode surface, these changes are shown as an increase in the charge transfer resistance of the system (from 0.14kΩ to 4.83kΩ) and a decrease in its double-layer capacitance (from 3.26μF to 1.12μF). The topographic images obtained via KPFM measurements show that the chemical layer and the bioreceptor make the gold electrode surface slightly rougher than its original form (for Au and Au / HDT / nano antibody constructs, root mean square roughness (RMS) = 1.57nm and 1.72nm). The electrode surface functionalized with nano antibodies seems to be relatively homogeneous, without any large aggregates, indicating that there are uniformly immobilized species (data not shown). Using KPFM, the contact potential difference of the gold electrode was measured and found to increase with the attachment of each functional layer (VAu=21.57mV, VHDT=60.61mV, Vnanobody=194.41mV). These results indicate that the functionality of the gold electrode is reduced due to the insulating chemical SAM and protein layers reducing the electron injection efficiency.

[0340] QCM-D enables real-time monitoring of the formation of these layers. Fig.9D The frequency changes of the crystal oscillations are shown after injection of SpyTag peptide and Ty1-SpyCatcher onto the HDT-coated gold surface (top). These changes correspond to 68 ng cm after the SpyTag peptide binds to the HDT layer. -2 After the introduction of Ty1-nanobody-SpyCatcher protein, 264 ng cm -2 The mass density increases to 5.7×10 per square centimeter of gold. 12 XPS confirmed the presence of HDT layer and biolayer on the gold surface ( Fig.9E ). The high-resolution C1s spectrum of HDT-functionalized gold exhibits two deconvoluted peaks. The main one is attributed to the six-carbon chain of the HDT molecule (CC, at 284.7 eV). A smaller peak appears at 286.2 eV, corresponding to the CO bond. This peak is attributed to the residual ethanol used during the functionalization and rinsing steps. After immobilization of the SpyTag peptide, new peaks at 285.2 eV and 288.5 eV appear in the C1s spectrum, corresponding to the CN bonds and C-OOR bonds derived from the amino acids of the peptide, respectively. In contrast, the high-resolution N1s spectrum of the HDT-functionalized gold surface has no features ( Fig.9F ). For the peptide-functionalized electrode, a nitrogen peak was observed at about 400 eV, and its intensity increased after nanobody functionalization. The presence of nanobody-SpyCatcher on the surface was confirmed by the increased signal of CO bonds in the entire C1s spectrum. Overall, these findings suggest that the functional layer is built on top of the gold electrode, such as Figure 8B shown.

[0341] ACET integrated sample incubation

[0342] Concentric grid electrode design ( Fig. 8A) allows the application of an AC potential through the solution to accumulate target molecules (such as spike proteins) on immobilized nanobodies. Due to the Joule heating effect, the inhomogeneous AC electric field will increase the local temperature. The local temperature increase will produce gradients in the density, dielectric constant and conductivity of the solution. Due to the combined effect of these gradients and the applied electric field, the electrohydrodynamic force is applied to the fluid, which stirs the solute molecules by the induced flow. However, the self-heating induced by ACET must be evaluated, especially in high ionic strength solutions (such as PBS (pH 7.4, ionic strength 0.162m)), to ensure that ACET does not generate too much heat and is safe to use in biological media. In order to estimate the amplitude of the temperature increase caused by Joule heating, numerical simulations are performed using a theoretical model containing fully coupled electrostatics, energy and Stokes systems. The accuracy of this numerical model has previously been verified by the thermal reflection method of a similar system. According to these simulations, the optimal AC application conditions are 6Vpp (pp: peak to peak) and 100kHz. Under these conditions, ACET at the gate electrode is predicted to result in an average temperature increase in the buffer of only a few degrees Celsius, while the average velocity of water can reach 15 μm s -1 (Data not shown). This speed can be sufficient to quickly transport the target protein to the sensor surface. In order to experimentally verify the simulation, the magnitude of the voltage applied to the gate electrode is changed, and the electrochemical impedance spectrum of the electrode is recorded immediately after stimulation. Increasing the amplitude from 2VPP to 6VPP has a relatively small effect on the spectrum (i.e., the diameter of the semicircle decreases with the increase of solution temperature) (data not shown). However, at a critical voltage of 8Vpp and above, the impedance decreases sharply, indicating that the biochemical layer is removed from the gold electrode surface. In order to determine how long the stimulation should be applied, different incubation times (i.e., 30, 60, and 120s) are used to perform sensing measurements (data not shown). The results show that if the gate electrode is incubated for 2min with the target sample under ACET, the sensor response is maximized (data not shown).

[0343] ACET applied using these conditions was evaluated to determine if it allowed specific binding events to occur on the gate electrode. A GFP nanobody construct specific for GFP was immobilized on the gate electrode to detect GFP, as this binding event could also be monitored using fluorescence microscopy (data not shown). -9m) After 2min of ACET-assisted incubation, a fluorescent signal was detected on the electrode. In contrast, under the same conditions, the GFP-nanoantibody electrode did not show a fluorescent response to the mCherry protein, which is a red fluorescent protein with low sequence homology to GFP but very high structural similarity. Similar measurements were performed using diffusion-assisted incubation, i.e., the protein solution was left on the electrode surface for 10min by some intermittent manual mixing. In contrast to the ACET-assisted experiment, some nonspecific binding of mCherry on the electrode surface was observed. It is speculated that ACET can help wash off nonspecific or low-affinity antigens that are only partially released in traditional washing steps by generating a continuous flow parallel to the electrode surface. Therefore, compared with sensors with diffusion-based incubation, sensors with ACET-assisted incubation are expected to have lower background signals. After determining the ACET application parameters and verifying their effects on reducing the binding of nonspecific analytes, the Ty1-nanobody-functionalized gate electrode and the GFP-nanobody-functionalized gate electrode were included in 10 μL of buffer solution containing increasing concentrations of target analyte (SARS-CoV-2 spike protein, S1) for 2 min while applying AC simulation ( Figure 8C ). The GFP-Nanobody electrode was used as a negative control because the GFP Nanobody should not bind to the Spike protein, but some nonspecific binding was observed in subsequent QCM-D experiments (data not shown).

[0344] Comparison of the performance of p-type and n-type OECTs in protein detection

[0345] After an ACET-assisted sample incubation step, gate electrodes were integrated with both types of OECT channels and device current-voltage characteristics were recorded. Figures 10A to 10D In the Figure 2, transfer curves (ID vs. VG) were plotted, which were recorded each time the gate was exposed to a new solution. Binding of SARS-CoV-2S1 to the gate electrode modulated the current flowing in both channels. For p(g3C2T2-T)OECTs, protein binding caused ID( Fig. 10A ) increases, accompanied by a shift in the transfer curve towards more positive VG values ​​(data not shown). For the p(C6NDI-T) device, target binding induces a similar shift, however, the ID undergoes a significant decrease in the low VG regime due to the bell-shaped profile of the transfer curve ( Fig. 10C). At high VG, the ID values ​​increased with the concentration of SARS-CoV-2S1. The shift in the transfer curve with analyte binding is attributed to the positively charged nature of the protein at pH 7.4. The bound protein induces an electric dipole at the gate / electrolyte interface and pushes the threshold voltage to a more positive value. In contrast, when the GFP-nanobody gate electrode was incubated with the same protein solution, both sensors produced only a weak response ( Fig. 10B and 10D ).

[0346] To compare the sensing performance of these devices, the normalized change in leakage current (NR = |ID-I0| / I0) under selected bias conditions was plotted as a function of protein concentration. Before exposing the Nanobody-OECT sensors to solutions containing the analyte, their reference (blank) responses (I0) were recorded. The VD value in the linear regime and the VG value that maximized the binding-induced change in current were selected. The current change was maximized near Vth for p(g3C2T2-T) and at two voltages for p(C6NDI-T); at VG = 0.5V, close to Vth, and at VG = 0.9V. Figures 10E to 10G These calibration curves (normalized response (NR) versus protein content) are shown. These calibration curves also include the channel current measured with an Ag / AgCl gate electrode. After each protein sensing measurement, an Ag / AgCl gate was used to account for any contribution to the sensor response associated with channel material instability. These control measurements are necessary given the S-shaped ID versus VD curves at high gate voltages (data not shown). As an additional investigation of stability, EIS, electrochemical QCM-D, and in situ Fourier transform infrared spectroscopy measurements were performed on the n-type films (data not shown). The results indicate that the high doping voltage does not lead to material degradation (data not shown). Under both gate bias conditions, the 100×10 -18 The p(C6NDI-T) current changes in response to the spike protein of m are sufficiently distinguishable from the response of the GFP gate electrode ( Fig.10F and 10G ). Only at or above 100×10 -15 m, the response of the p(g3C2T2-T) sensor to the target is significantly different from that of the GFP electrode gate device ( Fig.10E). Comparing these figures, it is concluded that the NR values ​​of n-type OECTs are higher than those of p-type devices, especially when reporting at VG = 0.9V. Although the p-type OECTs outperform the n-type in terms of transistor gain (gm 60mS vs. gm 15μS), the n-type OECTs reduce the detection limit and show a higher current response to changes in protein concentration when operated using the same biofunctionalized gate. The high performance of the p(C6NDI-T)OECTs is complemented by a low power requirement (100nW) compared to the p-type (100μW), which makes the p(C6NDI-T)OECTs compatible with integrated circuit designs and handheld battery-driven readers.

[0347] The binding between the nanobody and SARS-CoV-2S1 leads to an increase in impedance at the gate electrode. The capacitance of the protein-bound gate electrode decreases (C g,eff <C g ), resulting in a voltage drop at the gate electrode / electrolyte interface (V G 压降 ) and the change in the capacitive coupling between the gate and the channel ( Fig.10J The difference in sensor behavior between the two OECTs can be understood when comparing the gate-to-channel capacitance ratio (Figure 3K). In order to make an OECT with high gain, the goal is to make Cg as large as possible compared to the channel capacitance (e.g., C g =100C ch , Fig.10L However, for devices with a large gate / channel capacitance ratio, C g Small changes in the voltage drop across the channel (V CH 压降 ) has an impact. This is the case of the p(g3C2T2-T) channel, which has a very small capacitance (33±5nF) compared to the capacitance of the gate electrode (1.23μF). Due to its higher channel capacitance (146±11nF) and the resulting smaller C g / C ch , the p(C6NDI-T)-based OECT will generate V CH 压降 and ID, even for some binding events at the gate. Devices with smaller gate-to-channel capacitance differences enable more sensitive sensors. However, if the channel has a higher capacitance than the gate (e.g., Cg=0.1Cch, device 1 in FIG31 ), the OECT is not an effective amplifying transducer. Fig.10LWe show that by varying the channel material and / or the channel / gate geometry, it is possible to design immunosensors with specific sensitivity to low or high concentrations of analytes and adjust their dynamic range. It is important to note that capacitance reduction is not the only parameter governing the sensor's operating mechanism, as bound proteins at the gate electrode also change the electrochemical potential of this terminal.

[0348] Effect of ACET on the performance of n-type OECT sensors

[0349] The ACET-facilitated sample incubation method (2 min) was compared with the previously described conventional method (manual pipetting up and down for 30 s every 3 min during a 10 min incubation) ( Fig.10F For 10H, and Figure 10G ACET-driven SARS-CoV-2S1 detection performance is similar to that of conventional diffusion-based platforms, achieving a detection limit as low as 1×10-15m at VG = 0.5V ( Fig.10F and 10H As discussed above, the dense nanobody layer on the electrode surface produces a very high local receptor concentration (>100×10 -6 m), so that binding is no longer determined by the dissociation constant of the receptor-target interaction (25 × 10 -9 m). Instead, binding becomes diffusion-limited. Improved mixing can directly affect the diffusion-limited transport of the target to the gate surface. Compared with conventional incubation methods, ACET not only has the advantage of greatly shortened incubation time, but also gives higher NR values, especially for low protein concentrations. In addition to the diffusion-limited transport of the target to the gate surface, manual incubation may also lead to the loss of the target from the surface due to adsorption to the pipette tip during repeated mixing. In addition, the ACET flow parallel to the electrode surface can improve the wash-off of nonspecific or low-affinity antigens, resulting in lower background signals compared to diffusion-based sensing. Overall, ACET accelerates the transport of target molecules to the electrode, shortens the incubation time from ten minutes to 2min, and minimizes manual intervention during the sample incubation step.

[0350] ACET-assisted protein sensing in complex biological media

[0351] Recent studies compared saliva samples and nasopharyngeal samples from the same patients and concluded that saliva samples reduced variability and improved titer detection and consistency with RT-PCR. Self-collection of saliva is simple, non-invasive, and minimizes interaction between patients and healthcare providers, making saliva a preferred medium for diagnostic applications. Therefore, the ACET-integrated nanobody-OECT sensor faces the challenge of using human saliva as a sample medium. Three Ty1-functionalized gate electrodes were sequentially exposed to increasing concentrations of SARS-CoV-2S1 protein spiked into a 1:3 (v / v) saliva:lysis buffer mixture. As before, NR increased with increasing target concentration ( Fig.11A and 11B The obtained calibration curve ( Fig.11A ) shows that the spike protein concentration detected by Ty-1-nanobody-OECT is as low as 10×10 -15 M. As negative controls, three GFP-nanobody functionalized gate electrodes were exposed to the same spike protein / saliva samples. These control gate electrodes showed (albeit lower) responses to increasing spike protein concentrations.

[0352] Repeated exposure to the complex saliva:buffer:mixture may eventually lead to the accumulation of nonspecific background binding to any component of the biolayer or HDT surface. In contrast, in clinical applications, each gate would only encounter a single patient sample. This clinical scenario was outlined by exposing individual gate electrodes to randomly selected saliva samples with different SARS-CoV-2S1 concentrations ( Fig. 11C and 11D In addition, the same number of GFP-nanobody functionalized gate electrodes were used as controls to verify the specificity ( Fig. 11C and 11D small circle in the middle). Fig. 11C and 11DThese individual gate electrode measurements in the Figures show that Ty-1-nanobody OECTs always give higher NR values ​​than GFP-nanobody gated channels (compare small circles with large circles in each figure), but the NR values ​​vary greatly and do not seem to correlate with target concentration. Protein sensing in saliva (i.e., untreated samples (no heating, filtration, or centrifugation steps), whose complex content varies from patient to patient), has therefore proven challenging. It is speculated that the absolute NR values ​​may also be affected by variations in the Au electrode surface or biofunctionalization. Therefore, EIS and XPS measurements were performed on different gates prepared in parallel under the same conditions before and after target binding, and indeed considerable variability in capacitance was observed (data not shown). Further modifications to the manufacturing and biofunctionalization methods may help to address this gate-to-gate variation as well as background binding. The larger size of these gate electrodes may be a factor that increases the standard deviation of the sensor output. Larger geometries reduce surface homogeneity and make it more difficult to manufacture gate electrodes with immobilized chemical and biological layers of the same quality. However, comparison of several target and negative control measurements for multiple gates per sample can effectively block these two confounding factors. The difference between the means of the target and negative control measurements does correlate with protein concentration over a wide dynamic range ( Fig.11E ). Thus, these results demonstrate the sensitivity and selectivity of ACET-assisted nanobody OECTs in complex biological media and at physiologically and clinically relevant protein concentrations.

[0353] Finally, the nanobody / OECT sensor performance was compared to other similar thin-film transistor- and electrode-based devices developed for immunosensing applications and the performance of this technology was benchmarked in terms of power consumption, LOD, incubation time, and dynamic range ( Fig.11F and 11G 3). Although several biosensors have shown advantages, including single-molecule detection, nW power requirements, and 1 min incubation time, the device described here can detect concentrations as low as 30 × 10 in complex media using only 100 nW of operating power and 2 min incubation with 10 μL of sample. -18 M, up to 300×10 -9 Target proteins of M.

[0354] Table 3. Performance parameters of selected electrochemical immunosensors for various protein antigens

[0355]

[0356]

[0357]

[0358]

[0359]

[0360]

[0361]

[0362]

[0363]

[0364] *Reference List:

[0365] in conclusion

[0366] Described in this paper is a label-free electrochemical immunosensing technique that has ultrafast detection capabilities, 100 × 10 -18 m detection limit and large dynamic range (from ×10 -18 m to ×10 -9 m). New solution-processable n-type organic semiconductors enable transistors to operate in enhancement mode with high sensitivity and selectivity under very low bias conditions. ACET-induced mixing confinement reduces the time required for immune complex formation at the nanobody-functionalized gate electrode because it affects the diffusion-limited transport of proteins to the gate electrode surface. ACET also enhances detection sensitivity with lower standard deviation of sensor response compared to operation with diffusion-controlled incubation. ACET-enhanced nanobody-OECT biosensor can detect specific protein molecules from untreated saliva samples after only 2 minutes of incubation under ambient conditions. Measurements in complex samples use averaged measurements and off-target controls. This platform can detect any other target by changing the biorecognition unit, and its reusability, simple operation and detection speed can be used for routine biomarker screening.

[0367] Example 3. OECT functionalization with CA9 / sgRNA

[0368] Materials and methods

[0369] Construct design

[0370] Based on the available structures in the Protein Data Bank (PDB), the Cas proteins were fused to SpyCatcher, 3C protease cleavage site and 8x histidine purification tag (dSauCas9: 5AXW, dLwaCas13a: not available, dLbuCas13a: 5XWP, and dEsCas13d: 6E9F) at the N-terminus (dSauCas9) or C-terminus (all Cas13 proteins). The protein sequences were codon optimized for expression in E. coli and commercially gene synthesized (Twist) in the high-copy cloning vector pJEx411c (a modified version of the original DNA 2.0 / ATUM) containing kanamycin resistance and RBS insulator (BCD2) cassettes to improve translation initiation 68. The protein sequences are shown in Table 4.

[0371]

[0372]

[0373]

[0374]

[0375]

[0376] SpyTag peptides were amino-terminally maleimide-labeled and commercially synthesized (GenScript Biotech, Singapore). Lyophilized as received, dissolved in PBS, and stored at -20°C

[0377] Guide RNA and target RNA sequences were commercially synthesized as gBlocks containing a T7 promoter sequence followed by a 23 nt long spacer sequence and a canonical scaffold sequence of crRNA or direct repeat (DR) sequence, and flanked by a T7 terminator sequence at 3'. The gRNA sequences are shown in Table 5.

[0378]

[0379]

[0380]

[0381]

[0382]

[0383]

[0384]

[0385]

[0386]

[0387]

[0388] Bioinformatic evaluation of spacer sequences

[0389] Bioinformatics analysis was performed on a set of 36 probes and primers from standard RT-qPCR protocols and literature reviews. First, the location and frequency of single nucleotide polymorphisms (SNPs) in the SARS-Cov2 genome determined primers that were more likely to bind to this virus. Then, primer cross-references were evaluated by comparing the human genome and common human coronaviruses: MERS (MK129253.1), SARS-Cov1 (NC_004718.3), HCoV-229E (NC) 002645.1), Hepatitis B (NC_003977.2), HIV1 (NC_001802.1), NL63 (NC_005831.2), and OC43 (AY585228.1).

[0390] Protein expression and purification

[0391] The bacterial expression plasmid was transformed into Rosetta 2 (DE3) competent cells (Millipore). One colony was inoculated into 2xYT medium enriched with 1% glucose, kanamycin (50 μg / mL) and chloramphenicol (34 μg / mL) for initial culture, and then inoculated into 1L 2xYT production medium enriched with the same. Protein expression was induced at 18°C ​​for 18h with 400 μM IPTG. Cell lysis was performed at 20 kPsi in a cell disruptor. The lysate was clarified at 45,000 rpm for 30 min. The following was performed at The supernatant was processed on FPLC (GE Healthcare): affinity purification was first performed using a HisTrap column (GE Healthcare) and the target protein from the column was eluted with 80 to 240 mM imidazole. The fractions were then pooled and further purified by cation exchange using a HiTrap heparin column (GE Healthcare). After elution with nM NaCl, the fractions were pooled and concentrated using a 50K Amicon ultra centrifugal filter unit. Further purification was performed by gel filtration on a high load Superdex 200 into 20 mM HEPES-NaOH (pH 7.5), 500 mM NaCl, 1 mM TCEP, 5% glycerol. After spin concentration, aliquots were snap frozen in liquid nitrogen and stored at -80°C.

[0392] RNA production

[0393] Guide RNA and target RNA were transcribed using NEB HiScribeT7 High Yield RNA Synthesis Kit with 2.5% Cy5-UTP for Cy5 random labeled nucleic acids according to the manufacturer's protocol. Transcripts were purified according to the instructions of the RNAeasy kit, resuspended in double distilled water and stored at -20°C. RNA quality was visualized by denaturing gel electrophoresis (100 V for 1 h). Nucleic acid concentration was determined by absorbance (A260 nm) using NanoDrop one.

[0394] SEC-MALS

[0395] The purified protein was injected at 3 mg / ml by Agilent HPLC 1100 and separated by size in a Superdex200 10 / 300 column equilibrated in SEC-MALS buffer (20 mM HEPES-NaOH (pH 7.5), 500 mM NaCl, 1 mM TCEP) in advance, and then entered the sample compartment of a DAWN-EOS multi-angle laser light scattering detector and an Optilab-DSP relative refraction interferometer (Wyatt). Peak alignment and band broadening correction between UV, MALS and RWe detectors were performed using Astra software.

[0396] Binary complex assembly and precipitation assay

[0397] Immediately after thawing, warm the guide RNA to 70°C for 5 min. Perform complex formation at room temperature in complexation buffer at a 1:1 ratio of protein to guide RNA for 1 h to a final concentration of 5 μM.

[0398] EMSA

[0399] The binding reaction was performed in an RNase-free tube in which the binding buffer was dispensed, and the Cas: guide RNA complex was serially diluted (2xDF) to a final volume of 30 μL. The Cy5-labeled target or non-target, which was pre-warmed to 75°C for 5 min, was incorporated at a constant concentration. The binding reaction was incubated at room temperature for 1 h. The sample was diluted with non-denaturing loading buffer, loaded into a 3% (1xTAE or TBE) agarose gel, and run at 100V at 4°C for 2 h. Gel imaging was performed using an iBright FL1500 imaging system (ThermoFisher) by a fluorescence imaging protocol (Exc. 610 to 660nm, Em. 710 to 730nm).

[0400] MST

[0401] Short RNA targets were commercially synthesized by Integrated DNA Technologies (IDT, Inc.) and designed to contain a single nucleotide labeled with Cy5 at the 5' end. In vitro transcribed targets were randomly labeled and Cy5-UTP incorporated during transcription by including a mixture of UTP Cy5-labeled / unlabeled nucleotides (2.5:10) in the IVT reaction. The pre-assembled complex Cas:guide RNA was serially diluted in binding buffer, and then a constant concentration of target was pipetted in. The reactions were incubated at room temperature in the dark for 1 h. Measurements were performed using the PicoRed laser of the NanoTemper MST instrument at 40% LED power, medium MST power for short targets and positive IVT targets, and 5% LED power and medium MST power for IVT negative controls to achieve 20,000 counts at 250 pM in all cases. In total, 16 concentrations were analyzed per binding test, with three replicates for each concentration.

[0402] OECT manufacturing

[0403] OECTs were fabricated using photolithography and parylene C stripping techniques. Briefly, a layer of photoresist (AZ5214) was spin-coated on a glass substrate and treated with UV light using a mask aligner. After sputter coating the substrate with 10 nm of Cr and 100 nm of Au, a standard stripping process was performed in hot DMSO to remove the UV exposed areas. Then, a second layer of photoresist (AZ9260) was spin-coated on the same substrate to insulate the gold connection pads, followed by parylene C stripping. The OECT channel (10 μm in length and 100 μm in width) was formed by reactive ion etching techniques. The p-type accumulation mode organic film (p(g0T2-g6T2)) was dissolved in a chloroform solution (5 g / L) and then spin-coated (800 rpm, 45 s) on the substrate to produce a film thickness of approximately 70 nm in the channel.

[0404] The gate electrode was fabricated on a 175 μm thick Kapton (polyimide) substrate. The Kapton substrate was sputter coated with 10 nm of Cr and 100 nm of Au and patterned using a process cutter to produce the final form. The square geometry at the tip of the electrode defines the sensor active area (0.8 x 0.8 mm). The electrode was electrochemically cleaned in 10 mM sulfuric acid (H2SO4) using cyclic voltammetry (CV). The potential was scanned twenty times from -0.2 V to 1.2 V at a scan rate of 100 mV s-1.

[0405] Quartz Crystal Microbalance with Dissipation (QCM-D) Monitoring

[0406] Performed by collaborators from Prof.Inal's laboratory. QCM-D measurements were performed using a Q-sense analyzer (QE401, BiolinScientific). The piezoelectric active gold sensor (0.7854cm2) was pre-coated with a monolayer (SAM) of 1,6-hexanedithiol (HDT) self-assembly, and then placed in the QCM-D assembly. After the frequency (Δf) and dissipation (ΔD) in PBS were stabilized, the peptide solution (0.1mg / mL SpyTag peptide in PBS) was pumped into the fluid chamber at a flow rate of 100μL / min. After the solution completely covered the sensor surface, the pump was stopped to modify the surface of the sensor in static mode. After the sensor was incubated in the peptide solution for one hour, the sensor surface was rinsed with PBS for 15min to remove excess peptides from the surface. The same procedure was performed to expose the surface to 5μM of the complex protein (Cas: guide RNA) in the binding buffer. The 7th frequency was selected to present QCM-D data. The changes in mass and thickness during functionalization were calculated according to the Sauerbrey equation (1):

[0407] Δm=(-17.7) / n Δf_n (1)

[0408] Where n is the overtone number (-17.7), which is calculated considering the resonant frequency, effective area, density, and shear modulus of the quartz crystal sensor. The mass and thickness of each biochemical layer are extracted by using their molecular weight according to the following equation.

[0409] ds = Δm / Ar100 (2)

[0410] Where Δm is the change in mass (ng), A is the area (cm2), and r is the estimated density of the layer (g cm-3, assumed to be the same as water, taking 1 g cm-3).

[0411] Electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV)

[0412] Electrochemical impedance spectroscopy (EIS) and cyclic voltammetry (CV) techniques are used to check the formation of biochemical layer on the gold electrode. The three-electrode assembly is connected to a potentiostat (Autolab PGstat128N, with Nova software, MetroOhm) to carry out these two measurements. Platinum wire and Ag / AgCl electrode are used as counter electrode and reference electrode, respectively. The gold grid electrode (working electrode) is immersed in 5mL 10mM PBS (pH 7.4) containing 10mM [Fe (CN) 6] 3- / 4- to measure. For CV measurements, the potential window is scanned from -0.2V to 0.5V at a scan rate of 100mVs-1. Impedance spectrum is measured in a frequency range of 0.1 to 100000Hz at zero DC offset and open circuit potential and 10mV alternating current (AC) modulation. Use Nova software to analyze data using appropriate equivalent circuit modeling.

[0413] Gate biofunctionalization

[0414] A HDT-SAM solution was prepared in 100% ethanol containing 1 mM HDT. The gold electrode was immersed in this solution for one hour, then rinsed with ethanol and placed under nitrogen. 128 The electrodes were then dried under 4% CO 2 0 ...

[0415] Sensor Operation

[0416] The steady-state characteristics (IV) of transistors were recorded using a Keithley 2602A source meter unit operated by custom LabVIEW software. When the source electrode served as a common ground in two circuits, a drain voltage (VD) and a gate voltage (VG) were applied. Steady-state measurements of OECTs based on p(g0T2-g6T2) were performed by changing VG (0.2 to 0.6V, step length 0.05V) and VD (0 to -0.6V, step length 0.05V), and leakage current (ID) was obtained simultaneously.

[0417] A PDMS hole (1 cm in diameter, 2 mm thick) was placed on top of the OECT and filled with 100 μL of PBS to perform IV measurements. After the functionalization step, a reference (blank) response of the sensor was obtained by immersing the gate electrode in PBS. The same electrode was then incubated for 10 min with a 5 μL droplet of a binding buffer containing the protein, rinsed in the buffer and rinsed twice in PBS (phosphate buffered saline, pH 7.4, ionic strength 0.162 M), and then mounted in parallel on top of the trench to complete the OECT biosensor.

[0418] result

[0419] Sensor design: mechanism of action and characteristics

[0420] The study aims to integrate the high specificity and programmable features of RNA-targeted Cas systems with the high sensitivity of organic electrochemical transistors (OECTs) to rapidly (<15 min), sensitively, and selectively detect RNA in raw samples. The sensor includes a biorecognition unit consisting of three building blocks: a maleimide-modified SpyTag peptide, a SpyCatcher-fused Cas protein, and a guide RNA (gRNA). In addition, various versions of the sequence (long and short) were generated and used as targets or negative controls [data not shown].

[0421] The SpyTag / SpyCatcher system derived from the fibronectin binding protein FbaB of Streptococcus pyogenes spontaneously and irreversibly forms isopeptide bonds under different conditions (i.e., pH, buffer, and temperature). Its robustness and speed of bond formation (minutes) facilitate the biofunctionalization of the disclosed receptor modules on gate electrodes.

[0422] A flexible linker composed of glycine-serine amino acids was chosen to dynamically separate the Cas module from the anchoring SpyCatcher. With this flexibility, the aim is to increase the packing density on the gate surface while reducing steric hindrance between compact receptor units.

[0423] The ribonucleoprotein complex Cas:gRNA was preassembled in solution. The receptor module was then biofunctionalized on top of a gold gate electrode in a three-stage step. First, a 1,6-hexanedithiol (HDT) self-assembled monolayer (SAM) was formed on top of the gold electrode. Second, the maleimide-modified SpyTag peptide was chemically immobilized to produce a chem-SAM layer. Finally, the Cas receptor module was coupled via the SpyTag / SpyCatcher isopeptide covalent binding system.

[0424] After biofunctionalization of the gate, the sensor worked in three steps: 1) sample incubation (10 min) on top of the gate electrode with gentle mixing by pipetting up and down, 2) to increase specificity, the gate electrode was washed with binding or wash buffer and with PBS to separate nonspecific binders, 3) the gate was placed on top of the OECT base to complete the transistor setup and the electrical response was measured in PBS.

[0425] Selecting the best recognition unit (Cas protein) and target sequence to detect SARS-Cov 2

[0426] Many Cas proteins have been reported in the literature. A 5-step pathway was set up to select the top four Cas effectors that best met the sensor requirements: 1) Only Cas effectors that could directly target RNA were retained. 2) Cas effectors that required a protospacer adjacent motif (PAM) or protospacer flanking sites (PFS) were filtered out. 3) Effectors with in vitro experimental characterization were prioritized, 4) Cas proteins whose guide RNA sequences and features were reported were selected, and 5) Cas effectors with available kinetic and mechanistic information were prioritized.

[0427] Four Cas effectors were selected by this strategy: a) Staphylococcus aureus (SauCas9), b) Lactobacillus waarsenii (LwaCas13a), c) Lactobacillus buccae (LbuCas13a) and d) Eubacterium inermis (EsCas13d). These four Cas effectors belong to Class 2, which means that they work independently of other proteins but perform different functions through their multiple domains. LbuCas13a, LwaCas13a and EsCas13d belong to the Cas13 family, which targets RNA, has paralogous RNA cleavage activity and HEPN-dimer nuclease cleavage motifs 135 to 137. In contrast, SauCas9 can bind to and cleave both DNA and RNA and has two different nuclease domains (RuvC and HNH). The nuclease LwaCas13a is well known for its use in the SHERLOCK system, and unlike other selected effectors, it does not have a disclosed structure. EsCas13d is the smallest of the selected effectors and is also the only one in the d subtype of the Cas13 family. In one report, the protein showed better activity in vitro than in mammalian cells. Fortunately, the modularity of the disclosed sensing platform allows the four receptor units to be easily interchanged and assessed under similar conditions.

[0428] The next step was to design appropriate guide RNA sequences for each Cas effector. The sensor was designed to detect the SARS-Cov 2 isolate Wuhan-Hu 1 [Accession number: NC_045512.2]. In order to select the region to be targeted on the viral genome, SARS-Cov-2 targeting primer and probe sequences reported in the literature and those used by the standardized RT-qPCR protocol were collected. Given the dependence of Cas proteins on binding to targets with lower secondary structure, each reported primer sequence was evaluated against its status in the published SARS-Cov2 RNA secondary structure predictors. To ensure selectivity, the spacer sequence should also have a low probability of binding to human genes or other human-related viruses. Bioinformatics analysis and filtering were performed on all possible spacer sequences, taking into account their conservation within the SARS-CoV-2 genome and their differences from the human genome or viral sequences.

[0429] The six best spacer sequences and the four most mentioned Cas proteins were used for further experiments. Since the goal is to detect the direct binding of RNP complexes to RNA targets, the catalytic death patterns of the selected Cas proteins are designed and produced for these experiments. The three guide RNAs studied (N1, N2 and N3) target different regions of the SARS-Cov2 N gene, two guide RNAs (E1 and E2) target two different regions of the SARS-Cov2 gene E, and a guide RNA (RPP30) (designed as a control) targets a region of human ribonuclease RPP30. All 24 guide RNAs are constructed according to the literature for the corresponding Cas reports of the spacer sequences adding 23-nt lengths, either in the same direction (DR) or CRISPR RNA (crRNA).

[0430] Production, characterization and stabilization of identification units

[0431] All four catalytically dead (dCas) effectors were expressed and purified in E. coli. Further characterization was performed by size exclusion chromatography combined with multi-angle light scattering (SEC-MALS) [data not shown]. This analysis returned single peaks for three of the four effectors, with molecular weights within the expected range (<|10%| error). The only exception was dLwaCas13a, which exhibited two peaks with a mass difference of about 20% between the two peaks and a mass difference of -14% with the nearest peak of the expected molecular weight (which only represents about 30% of the total mass). This mass value is consistent with impurities present at similar molecular weights as observed in SDS-PAGE analysis. SDS-PAGE also indicated significant impurities in dSauCas9, but this significant impurity was not apparent under native conditions in the SEC-MALS chromatogram.

[0432] The disclosed system requires two different types of RNA: guide RNA and target RNA. Guide RNAs are primarily in vitro transcribed (IVT) and size and quality confirmed by electrophoresis (data not shown). Two different types of targets are used: a) a short sequence of 60 nucleotides in length containing a 23nt spacer / target sequence right in the middle, and b) a longer version consisting of a partial segment of the target gene sequence with one or more 23-nt spacers located at different positions along the target. The short targets are commercially synthesized, while the long targets are IVT'd.

[0433] The study aimed to evaluate the stability of ribonucleoprotein complexes at high concentrations of the complexes themselves and against three typical components of irritating buffers (such as those required for the dissolution of viral capsids): a) pH, b) detergent, c) salt concentration. Four buffers were defined and the protein stability therein was evaluated by observing protein precipitation. Under all conditions, dEsCas13d was the most stable effector, with no signs of precipitation under any of the conditions tested. The other three dCas proteins were sensitive to low salt concentrations and moderate to high precipitation. Generally speaking, all Cas proteins are stable at high salt, pH 7.4 and 0.01% Igepal or 1% Triton. The only exception is apo-dLwaCas13a, which precipitates in a buffer containing Triton.

[0434] Electrochemical characterization of OECT biofunctionalization

[0435] Subsequent studies evaluated whether gate electrode biofunctionalization directly affects the sensitivity and general functionality of OECTs. Good coverage of receptor units on the gate electrode is necessary, but is a limiting step in cases where large amounts of sample or long incubation times are required. The coupling of receptor units on the gate electrode can be evaluated by electrochemical impedance spectroscopy (EIS). Different combinations of time and concentration of the complex were tested, and the impedance spectra of each of these combinations were plotted on the Nyquist plot [data now shown]. This figure shows the expected behavior of having higher coverage / impedance (semicircles with larger diameters) at higher concentrations or longer incubation times. 5 μM and 1 h (respectively) were then selected as the preferred receptor unit concentration and incubation time for biofunctionalization. This condition allows high gate coverage while requiring only moderate Cas-gRNA complex concentrations. The study also characterized the binding of the Cas complex on the surface by monitoring the mass increment using a dissipative quartz crystal microbalance (QCMD) [data not shown]. These experiments showed a steady mass increase after coupling dCas to the sensor surface for all four Cas effectors.

[0436] Initial OECT sensing experiments: yeast tRNA and Igepal in binding buffer to reduce noise and improve selectivity

[0437] Next, the normalized response (NR) at different target or negative control concentrations was analyzed to characterize target binding on the sensor by cyclic voltammetry (CV). First, 10 min was set as the target incubation time required for moderate signal. However, the data showed that both sensitivity and selectivity needed to be improved. A new binding buffer (WBB3) was formulated, which was enriched with yeast tRNA, higher salt, and a different detergent (Igepal CA-63 instead of Tween-20). The new formulation resulted in higher selectivity and sensitivity.

[0438] Binding is conserved at high detergent concentrations, whereas affinity depends on target length and the presence of yeast tRNA

[0439] In parallel with the OECT characterization work, follow-up studies performed biochemical characterization of the binding properties of candidate Cas effectors in solution. Electrophoretic mobility shift assays (EMSA) confirmed the selectivity of both dSauCas9 and dEsCas13d for target E2 [data not shown]. Interestingly, in both cases, heavier bands were observed at 0.5 μM dSauCas9:gRNA and 2 μM dEsCas13d:gRNA. While not being bound by theory, two possible explanations are: a) the complex multimerizes at high concentrations, or b) the ternary complex presents a heterogeneous conformational mode that migrates slower in native electrophoresis.

[0440] The equilibrium binding affinity of different Cas effectors was studied by microthermophoresis (MST). First, when a short target (from gene E2) and a short negative control (from gene N2) were combined in a binding buffer without yeast tRNA, the Cas effectors were combined. In this experiment, dSauCas9 and dEsCas13d showed the highest selectivity and a stronger affinity (about 10nM) for the target. Similarly, dLwaCas13a showed high selectivity, but the affinity was weak to nine times. Unlike the other three Cas effectors, dLbuCas13a binds to the target and the negative control with similar affinity, but there are differences in the direction of the binding curve (data not shown).

[0441] Based on their expression yield, SEC-MALS characterization, and target binding performance, dSauCas9 and dEsCas13d were selected for further analysis.

[0442] Since secondary structures are more likely to occur in longer RNA molecules, studies were performed to assess the affinity of dSauCas9 and dEsCas13d for longer targets, again in binding buffer, in the absence of yeast tRNA [data not shown]. Surprisingly, the performance of both effectors was reduced. In both cases, MST did not reach the response amplitude observed in the case of short targets (half that of dSauCas9). More importantly, dSauCas9 now exhibited nonspecific binding to the negative control while affinity was comparable to target binding.

[0443] dSauCas9 and dEsCas13d restored their specificity for long RNA targets and also increased the amplitude of the MST response in the presence of yeast tRNA [data not shown]. However, yeast tRNA blockade also reduced the binding affinity to the target to some extent. Therefore, yeast tRNA can block the interaction between non-target RNA and Cas: gRNA complex, but also compete with the interaction between target RNA and Cas complex to some extent.

[0444] Binding competition assays were performed to examine how yeast tRNA interferes with target binding to the complex. In this experiment, the concentration of yeast tRNA was variable, while the concentration of the ternary complex dSauCas9: guide RNA: target / negative control was kept constant [data not shown]. Yeast tRNA binds to both complexes with target or negative control RNA (and may outcompete both target binding and non-target binding). However, the observed half-maximal effective concentration (EC50) was different in the presence of target or non-target RNA: non-non-target RNA was more easily displaced, resulting in an EC50 value three times smaller. In general, tRNA can act as an effective blocking agent to prevent nonspecific binding to the Cas complex. 4 μM was used as the concentration of yeast tRNA in the new binding buffer. This concentration maximizes its blocking activity (due to the smaller EC50 (6.4 μM) required for the negative control) while still minimizing competition with target binding.

[0445] dSauCas9 and dEsCas13d target binding is conserved in lysis buffer

[0446] The aim was to perform the experiment in raw saliva samples. For this, a lysis buffer was required. Lysis buffers typically contain high concentrations of detergents that could damage or inhibit proteins. The compatibility of the dSauCas9 and dEsCas13d binding systems was tested in binding buffers containing high concentrations of detergent (1% Triton). The MST experiments were repeated with the simplest setup (for short targets, and in the absence of yeast tRNA but in the presence of high detergent concentrations) to directly explore the effects of detergents on protein performance. The MST experiments confirmed the stability and selectivity of the receptor units studied. Interestingly, in buffer with 1% Triton, the binding affinity of Cas effectors to target RNA was stronger (KD 6.3 nM (±1.3) for dSauCas9 and 0.8 nM (±0.5) for dEsCas13d) than previously assessed with only 0.5% Igepal (KD 9.5 nM (±1.3) for dSauCas9 and 10.7 nM (±7.7) for dEsCas13d) [data not shown].

[0447] dEsCas13d is the most sensitive and specific Cas effector in preliminary OECT experiments

[0448] Four Cas effector proteins were tested on the OECT sensor. A longer version of the SARS-Cov2 gene E2 was used as a target, and the human gene RPP30 was used as a negative control. Experiments were performed in lysis buffer (with 1% Triton) and in a mixture (4:1) of lysis buffer and raw saliva (spiked with target or negative control). Performance in pure buffer was generally less noisy than in buffer and saliva. dEsCas13d and dLwaCas13a stood out for their reproducibility, sensitivity, and selectivity.

[0449] discuss

[0450] Most Cas-based RNA detection methods use (indirect) collateral cleavage of a Cas protein-reported RNA molecule to report target detection. Some sensors are only able to detect DNA, cannot measure oligonucleotides in the original sample, require additional steps to achieve low noise, and operate at relatively high voltages because it uses OFETs instead of OECTs to amplify the target binding signal. The disclosed devices and methods combine programmable binding of CRISPR-Cas proteins to target RNA with large signal amplification of OECTs for rapid, reliable and sensitive detection of RNA in untreated samples.

[0451] Unlike other CRISPR-Cas based approaches that rely primarily on, typically, optimization for a single Cas effector, the disclosed modular design allows for the performance of four different RNA-targeting Cas effectors to be evaluated in parallel. As expected, the individual mechanisms and features of the proteins influence their stability, biochemical binding, and sensing properties. The data show that the dEsCas13d:guide RNA complex is a very stable complex and does not precipitate at high concentrations even in buffers containing harsh detergents.

[0452] Although preliminary, the OECT sensing experiments and MST results suggest a clear dependency benefit of yeast tRNA for improved selectivity of binding to longer targets.

[0453] This study biophysically characterized four Cas proteins, their stability as binary complexes, their binding activity in solution and when immobilized on the surface of the disclosed OECT sensor. In fact, one of the advantages of the disclosed multi-module sensor platform is the flexibility to test the performance of four different receptor units under similar conditions.

[0454] OECT results achieved detection limits in the high attomole to femtomolar range in about 15 min from sample to result in both lysis buffer and saliva. No sample pretreatment was required. Sample volumes were less than 10 μL. Testing of the dEsCas13d receptor unit indicated good guide RNA-dependent selectivity.

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Claims

1. A biosensor comprising an organic electrochemical transistor (OECT) and a biorecognition layer, wherein the OECT comprises a source electrode, a drain electrode, a channel and a gate electrode, wherein the source electrode and the drain electrode are electrically connected via the channel, and the gate electrode is removable and located at a position separated from the source electrode, the drain electrode and the channel.

2. A biosensor comprising an array of two or more OECTs and a biorecognition layer, wherein the array comprises two or more source electrodes, two or more drain electrodes, two or more corresponding channels and a common gate electrode, wherein each source electrode is paired with a drain electrode and is electrically connected to the drain electrode via a corresponding channel, and the common gate electrode is removable and located at a position separated from the source electrode, the drain electrode and the corresponding channel. 3 . The biosensor according to claim 1 , wherein the biorecognition layer is integrated on the gate electrode or the common gate electrode of the OECT.

4. The biosensor according to any one of claims 1 to 3, wherein the orientation of the biorecognition layer relative to the OECT surface can be represented by the following general formula: L1-AP1:AP2-L2-B Formula I wherein L1 is a first linker, AP1 is a first peptide binding partner; AP2 is a second peptide partner; AP1 and AP2 are binding partners, L2 is a second linker and B is a biological recognition element, wherein the binding of AP1 and AP2 produces a biological self-assembled monolayer (Bio-SAM), optionally wherein said L1 comprises an N- or C-terminal cysteine ​​residue.

5. The biosensor according to claim 3 or 4, wherein the biological recognition element is a two-domain or single-domain antibody fragment, such as a nanobody or a Cas protein.

6. The biosensor according to any one of claims 3 to 5, wherein the biological recognition element recognizes the SAR-CoV-2 receptor binding domain. 7 . The biosensor according to claim 3 , wherein the L1 is directly attached to a surface of the gate electrode or a common gate electrode.

8. The biosensor according to any one of claims 3 to 7, wherein the L1 comprises SEQ ID NO: 22 or SEQ ID NO:

23.

9. The biosensor according to any one of claims 3 to 8, wherein the AP1 to AP2 are selected from the group consisting of SpyTag / Spy Catcher peptide conjugates, snoopCatcher / snoopTag, MoonTag / MoonCatcher, SnoopTagJr / SnoopCatcher or DogTag and SdyTag / SdyCatcher peptide conjugates.

10. The biosensor of claim 9, wherein the AP1 comprises CGGSGSGSGAHIVMVDAYKPTK (SEQ ID NO: 24) or AHIVMVDAYKPTKGSGC (SEQ ID NO: 25).

11. A biosensor according to any one of claims 1 to 10, wherein the or each channel comprises a conducting polymer selected from the group consisting of: PEDOT:PSS, PEDOT-S, PEDOT:TOS, PEDOTOH:ClO 4 、PEDOT-co-PEDOTOH:ClO 4 , P3HT, PTHS, BBL, p(g2T-TT), PTHS-TMA+-co-P3HT, p(gNDI-g2T), p(g0T2-g6T2), P-90, p(g 3 C 2 T2-T) and p(C 6 NDI-T).

12. A method for integrating a biorecognition layer on an electrode, the method include: (i) incubating at least a portion of the surface of the electrode with a first incubation solution comprising a first conjugate to produce a first peptide-binding partner-modified surface via a first linker, wherein the first conjugate comprises the first linker and a first peptide-binding partner, and (ii) incubating the surface modified with the first peptide binding partner with a second incubation solution comprising a second conjugate and a blocking agent, wherein the second conjugate comprises a second binding partner and a biorecognition element, wherein the first peptide binding partner is conjugated to the second peptide to form a bond.

13. The method of claim 12, wherein the first peptide binding partner is a SpyTag peptide.

14. The method according to claim 12 or 13, wherein the second conjugate is a SpyCatcher-Nanobody conjugate.

15. The method of claim 14, wherein the Nanobody is a SAR-CoV-2 receptor binding domain binding Nanobody.

16. The method according to any one of claims 12 to 15, wherein the blocking agent is BSA.

17. A method for detecting the absence, presence or concentration of an analyte in a biological sample, the method comprising contacting the biological sample with the biosensor of claim 1 or 2, wherein the biosensor further comprises a reservoir, the method include: (v) incubating the gate electrode with the biological sample for a period of time sufficient to allow binding between the analyte and the biological recognition element; (vi) rinsing the gate electrode with a rinse buffer; as well as (vii) measurement signal ID, Wherein a difference between the signal ID and a background ID indicates the absence, the presence, or the concentration of the analyte in the biological sample, and wherein the biological sample is in liquid form.

18. The method according to claim 17, wherein the method further comprises the step of measuring the background ID by: (i) incubating the gate electrode with a blank solution; (ii) placing the gate electrode on top of the channel; (iii) applying VG and VD; (iv) measuring the background ID.

19. The method according to claims 17 and 18, wherein the method further comprises The following steps are involved: An electrolyte solution is added to the reservoir prior to any of steps (i) to (vii).

20. The method according to any one of claims 17 to 19, wherein steps (v) to (vii) are repeated one or more times.

21. The method of any one of claims 17 to 20, wherein the biological sample is (a) a body fluid selected from the group consisting of whole blood, plasma, serum, saliva, mucus, sputum, bronchoalveolar lavage (BAL), bronchial washings (BW), cerebrospinal fluid (CSF) and urine, or (b) a non-body fluid.

22. The method according to any one of claims 17 to 21, further comprising: The following steps are involved: The sample is processed into the biological sample prior to any one of steps (i) to (v), wherein the sample is selected from the group consisting of tissue, stool, rectal swab, nasopharyngeal swab and throat swab, optionally wherein the sample is mixed with a buffer prior to incubation.

23. The method according to any one of claims 17 to 22, further comprising: The following steps are involved: A protease inhibitor is added to the biological sample prior to any one of steps (i) to (v), wherein the sample is a saliva sample.

24. The method of any one of claims 17 to 23, wherein the volume of the biological sample is less than 20 μL, less than 10 μL, or less than 5 μL.

25. The method of any one of claims 17 to 24, wherein the gate electrode is incubated with the biological sample for a period of up to 60 minutes, up to 50 minutes, up to 40 minutes, up to 30 minutes, up to 20 minutes or up to 10 minutes.

26. An alternating current thermal flow (ACET) enhanced biosensor comprising an organic electrochemical transistor (OECT) based biosensor and a conductive layer for inducing ACET, wherein the OECT based biosensor comprises an OECT and a biorecognition layer, wherein the OECT comprises a source electrode, a drain electrode, a channel and a gate electrode, wherein the source electrode and the drain electrode are electrically connected via the channel, and the gate electrode is removable and located at a position separated from the source electrode, the drain electrode and the channel.

27. An ACET-enhanced biosensor according to claim 26, wherein the OECT-based biosensor comprises an array of two or more OECTs, wherein the array of OECTs comprises two or more source electrodes, two or more drain electrodes, two or more channels and a gate electrode, wherein each source electrode is paired with a drain electrode and is electrically connected to the drain electrode via a channel, and the gate electrode is removable and located at a position separated from the source electrode, the drain electrode and the channel.

28. The ACET enhanced biosensor according to claim 26 or 27, wherein the conductive layer is placed in close proximity to the gate electrode of the OECT.

29. The ACET enhanced biosensor according to any one of claims 26 to 28, wherein the conductive layer is disposed along an edge of the gate electrode and surrounds at least a portion of the gate electrode.

30. The ACET enhanced biosensor according to any one of claims 26 to 29, wherein the gate electrode is circular in shape, wherein the conductive layer has an arch shape surrounding at least a portion of the circular gate electrode.

31. The ACET enhanced biosensor of any one of claims 26 to 30, wherein the conductive layer comprises one or more extensions configured for contacting a power source for applying an AC potential.

32. The ACET enhanced biosensor according to any one of claims 26 to 31, wherein the biorecognition layer is integrated on the gate electrode of the OECT.

33. The ACET enhanced biosensor according to any one of claims 26 to 32, wherein the orientation of the biorecognition layer relative to the OECT surface can be represented by the following general formula: L1-AP1:AP2-L2-B Formula I Wherein L1 is the first linker, AP1 is the first peptide binding partner; AP2 is the second peptide partner; AP1 and AP2 are binding partners, L2 is the second linker and B is the biorecognition element, wherein the binding of AP1 and AP2 produces a bio-self-assembled monolayer (Bio-SAM).

34. The ACET-enhanced biosensor of claim 33, wherein the L1 comprises an N- or C-terminal cysteine ​​residue.

35. The ACET enhanced biosensor according to claim 33 or 34, wherein the L1 is directly attached to the surface of the gate electrode.

36. The ACET-enhanced biosensor according to any one of claims 26 to 32, wherein the orientation of the biorecognition layer relative to the OECT surface can be represented by the following general formula: N-L1-AP1:AP2-L2-B Formula I' wherein N is an organic molecule capable of self-assembly to form a first SAM, L1 is a first linker, AP1 is a first peptide binding partner; AP2 is a second peptide partner; AP1 and AP2 are binding partners, L2 is a second linker and B is a biorecognition element, wherein N and AP 1 Chemical conjugation produces a chemically self-assembled monolayer (Chem-SAM), and the combination of AP1 and AP2 produces a biologically self-assembled monolayer (Bio-SAM).

37. The ACET enhanced biosensor of claim 36, wherein N is an alkanethiol or a derivative thereof.

38. The ACET enhanced biosensor according to any one of claims 33 to 37, wherein the biological recognition element is a di-domain or single domain antibody fragment, such as a nanobody or a Cas protein.

39. The ACET-enhanced biosensor according to any one of claims 33 to 38, wherein the biological recognition element recognizes the SAR-CoV-2 receptor binding domain.

40. The ACET-enhanced biosensor of any one of claims 33 to 39, wherein the L1 comprises SEQ ID NO: 22 or SEQ ID NO:

23.

41. The ACET enhanced biosensor according to any one of claims 33 to 40, wherein the AP1 to AP2 are selected from the group consisting of SpyTag / Spy Catcher peptide conjugates, snoopCatcher / snoopTag, MoonTag / MoonCatcher. SnoopTagJr / SnoopCatcher or DogTag and SdyTag / SdyCatcher peptide conjugates.

42. The ACET enhanced biosensor of any one of claims 33 to 41, wherein the AP1 comprises CGGSGSGSGAHIVMVDAYKPTK (SEQ ID NO: 24) or AHIVMVDAYKPTKGSGC (SEQ ID NO: 25).

43. An ACET enhanced biosensor according to any one of claims 33 to 42, wherein the or each channel comprises a conducting polymer selected from the group consisting of: PEDOT:PSS, PEDOT-S, PEDOT:TOS, PEDOTOH:ClO 4 、PEDOT-co-PEDOTOH:ClO 4 , P3HT, PTHS, BBL, p(g2T-TT), PTHS-TMA+-co-P3HT, p(gNDI-g2T), p(g0T2-g6T2), P-90, p(g 3 C 2 T2-T) and p(C 6 NDI-T).

44. A method of detecting the absence, presence or concentration of an analyte in a biological sample, the method comprising contacting the biological sample with the ACET-enhanced biosensor of claim 26, wherein the ACET-enhanced biosensor further comprises a reservoir, the method include: (v) incubating the gate electrode with the biological sample for a period of time sufficient to allow binding between the analyte and the biological recognition element; (vi) rinsing the gate electrode with a rinse buffer; as well as (vii) measurement signal ID, Wherein a difference between the signal ID and a background ID indicates the absence, the presence, or the concentration of the analyte in the biological sample, and wherein the biological sample is in liquid form.

45. The method of claim 17, wherein before and / or during step (v), an AC potential is applied to the conductive layer of the biosensor, wherein the AC potential is sufficient to induce a micro-stirring effect.

46. ​​The method of claim 45, the AC potential being in the range of 1 Vpp to 8 Vpp or 2 Vpp to 6 Vpp, such as about 6 Vpp (pp means peak to peak).

47. The method according to any one of claims 44 to 46, wherein the method further comprises the step of measuring the background ID by: (i) incubating the gate electrode with a blank solution; (ii) placing the gate electrode on top of the channel; (iii) applying VG and VD; and (iv) measuring the background ID, Optionally, wherein an AC potential is applied to the conductive layer of the biosensor before and / or during step (i), and optionally, wherein the AC potential is the same as the AC potential applied to the conductive layer before and / or during step (v).

48. The method according to any one of claims 44 to 47, wherein the method further comprises The following steps are involved: An electrolyte solution is added to the reservoir prior to any of steps (i) to (vii).

49. The method of any one of claims 44 to 48, wherein steps (v) to (vii) are repeated one or more times.

50. The method of any one of claims 44 to 49, wherein the biological sample is (a) a body fluid selected from the group consisting of whole blood, plasma, serum, saliva, mucus, sputum, bronchoalveolar lavage (BAL), bronchial washings (BW), cerebrospinal fluid (CSF), and urine, or (b) a non-body fluid.

51. The method according to any one of claims 44 to 50, further comprising: The following steps are involved: The sample is processed into the biological sample prior to any one of steps (i) to (v), wherein the sample is selected from the group consisting of tissue, stool, rectal swab, nasopharyngeal swab and throat swab, optionally wherein the sample is mixed with a buffer prior to incubation.

52. The method according to any one of claims 44 to 51, further comprising: The following steps are involved: A protease inhibitor is added to the biological sample prior to any one of steps (i) to (v), wherein the sample is a saliva sample.

53. The method of any one of claims 44 to 52, wherein the volume of the biological sample is less than 20 μL, less than 10 μL, or less than 5 μL.

54. The method of any one of claims 44 to 53, wherein the gate electrode is incubated with the biological sample for a period of up to 5 minutes, up to 4 minutes, up to 3 minutes, up to 2 minutes, or up to 1 minute.

55. The method of any one of claims 44 to 54, wherein power consumption is ≤ 100 nW.

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