Single-bead capacitive detector for microbiological applications.

Abstract

The present invention provides an improved capacitive bead sensor for detection and / or quantification of target analytes in a sample, with a detection limit down to a single bead, reusable for multiple bead tests, or continuous flow of beads, and easily manufacturable and automatable, which allows sensitivity down to single molecule detection without the need for enzymatic amplification such as PCR, through the use of various structural advantages and electronic signal amplification techniques that further enable multiplexed target detection across a range of nucleic acid targets as well as across entire target classes, for example allowing simultaneous detection of viral nucleic acid and host antibodies against that virus.

Description

[Technical field]

[0001] Related Applications This application claims the benefit of and priority to U.S. Provisional Patent Application No. 63 / 304,312, filed January 28, 2022, the entire contents of which are incorporated herein by reference.

[0002] FIELD OF THEINVENTION The present invention relates generally to nucleic acid and antibody / antigen molecule capture by magnetic beads and methods for detection / quantification of these molecules, even at the single bead level and thus single molecule detection. [Background technology]

[0003] background The capture and detection of nucleic acids (NA) from biological samples has become an important aspect of early disease detection and monitoring. In traditional clinical laboratory evaluation, both molecular and serological (antibody / antigen) tests are used together in providing a definitive diagnosis. Ideally, such tests are performed together from one sample. As an example, PCR or LAMP tests can provide qualitative or perhaps semi-quantitative information on the presence of viral RNA, while separate antibody detection can establish a person's past infection status. Quantification of captured NA is important, for example, in determining the viral load of an infectious pathogen or for gene expression analysis of multiple captured NA.

[0004] Multiplex testing, variant identification, and genotyping have also become very important, as seen in the recent COVID-19 pandemic, where "variants of concern" (e.g., delta, omicron) have caused major health problems and many deaths. Direct single molecule detection with ultra-high sensitivity may enable the analysis of rare mutations that can cause cancer, detection of trace contamination of biopharmaceuticals with pathogenic DNA or RNA, very early or late detection of viral infection in patients with extremely low viral loads (see "single-molecule-PCR" U.S. Pat. No. 5,811,235, incorporated herein by reference), and template generation in cloning and DNA and RNA sequencing. Drawbacks of ultra-sensitive methods may be cross-contamination with amplicons, accidental contact of suspects at the crime scene, and continued testing positive for the presence of pathogenic RNA or DNA for long periods of time after the patient has recovered from the infection and is no longer infectious. These are specific characteristics of enzymatic amplification methods.

[0005] These highly important health tests are currently available only in highly specialized laboratories and require highly trained staff as well as significant equipment and infrastructure, all of which contribute to significant costs.

[0006] The use of magnetic beads for nucleic acid capture is known. The optimal magnetic bead diameter for this ranges from 400 nm to 2800 nm. Non-magnetic particles (e.g., glass, silica, polystyrene, titania, silver, or gold) with even smaller diameters, typically down to 40 nm, are also used for nucleic acid capture. A notable use of nanoparticles in diagnostics is lateral flow tests, where they are conjugated with antibodies for colorimetric detection of antigens, or vice versa, with antigens for antibody testing using the same technology.

[0007] The captured DNA is eluted and, if the target DNA is present and hybridizes with the matching oligomer probe and primers, it is typically amplified exponentially in a polymerase chain reaction (PCR) or loop-mediated isothermal amplification (LAMP) assay for optical detection. If the captured target is RNA, a reverse transcription step (RNA to cDNA) is first used.

[0008] Conventional enzymatic amplification is used to detect target NAs, if present. However, these PCR and LAMP methods have difficulty in accurate quantification of target NAs in samples due to the stochastic nature of exponential amplification, and it remains difficult to measure and quantify samples with low concentrations. Single molecule detection using PCR and LAMP is also challenging due to enzyme inhibition by certain samples, and the length of the primer probes, which are typically >18 (generally 20+) and >50 nucleotides, respectively. This prevents binding to highly significant mutations, such as those in the receptor binding domain (RBD) of the spike gene of the SARS-CoV-2 virus, where many of the single nucleotide mutations (e.g., in delta or omicron variants) can occur within 15-20 nucleotides of each other. Droplet / digital PCR and sequencing can address single molecule detection and identification, but those methods typically require large and extremely expensive laboratory equipment and trained expert operators.

[0009] Detection of magnetic beads carrying target molecules on complementary metal oxide semiconductor (CMOS) chips has been proposed, e.g., Hall sensors as described in Florescu et al. (2010) "On-chip magnetic separation of superparamagnetic beads for integrated molecular analysis" J Appl. Phys. 107(5):054702, and GMR-SV sensors as described in Murmann et al. (2013) "A 256 pixel magnetoresistive biosensor microarray in 0.18 μm CMOS", IEEE J Solid-State Circuits 48(5):1290-1301, the entire contents of each of which are incorporated herein by reference.

[0010] However, each of these CMOS chips requires extra magnetic and metal special layers which are expensive and generally not available in standard high volume CMOS foundry processes.

[0011] Chang and Lu (2013), “CMOS capacitive biosensors for highly sensitive biosensing applications”, Annual Int Conf IEEE Eng Med Biol Soc. 2013:4102-5, the entire contents of which are incorporated herein by reference, propose capacitive detection of magnetic beads on a CMOS chip.

[0012] Detection of non-magnetic particles such as glass, silica, titania, silver, and gold nanoparticles typically requires expensive photodetectors, lasers, X-ray or SEM, or mass spectrometry equipment. Some lateral flow tests allow visual detection by the human eye, where a large number of typically 40 nm nanoparticles are captured at a detection (or control line) to allow visual detection. However, the sensitivity of these methods is limited to a few thousand particles.

[0013] U.S. Patent No. 10,746,683 (Cummins et al.), the entire contents of which are incorporated herein by reference, describes an interdigitated electrode (IDE) capacitance sensor for detecting both the magnetic and non-magnetic particles listed above. This is a disposable sensor with a lower detection limit of 200 beads, in which the water or buffer carrier evaporates and the beads remain permanently attached to the sensor surface. U.S. Patent No. 10,160,966 (O'Farrell et al.), the entire contents of which are incorporated herein by reference, discloses beads that can be provided to a sample and peptide nucleic acid (PNA) probes attached to these beads that hybridize to and capture target nucleic acid molecules, in a disposable configuration. U.S. Patent No. 11,459,601 (O'Farrell et al.), the entire contents of which are incorporated herein by reference, discloses various assay steps for nucleic acid detection. Paramagnetic transport (T) beads with attached first PNA probes and captured nucleic acids are magnetically removed from the sample and moved through various washing and tethering steps. Reporter (R) beads with attached second PNA probes are then tethered to the captured target nucleic acid, if present, to create a target-specific sandwich assay. This is then moved to a washing chamber and onto a CMOS sensor chip, where the R beads are eluted and detected. This is also a disposable configuration. [Prior art documents] [Patent documents]

[0014] [Patent Document 1] U.S. Patent No. 5,811,235 [Patent Document 2] U.S. Patent No. 10,746,683 [Patent Document 3] U.S. Patent No. 10,160,966 [Patent Document 4] U.S. Patent No. 11,459,601 [Non-patent literature]

[0015] [Non-Patent Document 1] Florescu et al. (2010) “On-chip magnetic separation of superparamagnetic beads for integrated molecular analysis” J Appl.Phys.107(5):054702 [Non-Patent Document 2] Murmann et al. (2013) “A 256 pixel magnetoresistive biosensor microarray in 0.18μm CMOS”, IEEE J Solid-State Circuits 48(5):1290-1301 [Non-Patent Document 3] Chang and Lu (2013), "CMOS capacitive biosensors for highly sensitive biosensing applications", Annual Int Conf IEEE Eng Med Biol Soc.2013:4102-5 Summary of the Invention

[0016] The systems and methods of the present disclosure generally relate to the detection and / or quantification of target analytes in a sample.Various structures and technologies discussed herein, alone or in combination, can provide sensitivity down to single molecule detection without the need for enzymatic amplification such as PCR.The systems and methods of the present invention can use various structural advantages and electronic signal amplification technologies that further enable multiplex target detection across a variety of nucleic acid targets as well as across all target classes, for example, allowing simultaneous detection of viral nucleic acid and host antibodies against that virus.

[0017] Aspects of the invention may include nucleic acid detection methods that use capacitive sensors to detect target nucleic acids present in a sample at 100 copies / mL or less without enzymatic amplification. In various embodiments, the techniques discussed throughout may be used alone or in combination to achieve detection of 150 copies / mL or less, 250 copies / mL or less, 500 copies / mL or less, 1000 copies / mL or less, 5000 copies / mL or less, 10,000 copies / mL or less, 50,000 copies / mL or less, 100,000 copies / mL or less, 1,000,000 copies / mL or less, etc., depending on the desired sensitivity, precision, and complexity of the assay. In certain embodiments, the method achieves detection of at least 10 copies / mL of target nucleic acid. 9 In various embodiments, the method can be operable to achieve a signal amplification of at least 10 8 , at least 10 7 , at least 10 6 , at least 10 5 , at least 10 4 , at least 10 3 , or at least 10 2 Similarly, the method may be operable to achieve signal amplification upon detection of the target nucleic acid equivalent to at least 30 PCR cycles, at least 25 PCR cycles, at least 20 PCR cycles, at least 15 PCR cycles, at least 10 PCR cycles, or at least 5 PCR cycles.

[0018] The nucleic acid detection by the capacitance sensor may include a full-scale range of at least about 8 pF, at least about 7 pF, at least about 6 pF, at least about 5 pF, at least about 4 pF, at least about 3 pF, at least about 2 pF, or at least about 1 pF. The method may include converting the signal from the capacitance sensor using a capacitance-to-digital converter. The capacitance-to-digital converter may be a sigma-delta 24-bit capacitance-to-digital converter. The nucleic acid detection by the capacitance sensor may include a resolution of at least about 0.5 aF and an accuracy of at least about 4 aF. In various embodiments, the nucleic acid detection by the capacitance sensor may include a full-scale range of at least about 15 pF by including an internal offset capacitor in the reference input.

[0019] The method of the present invention may further include binding a target nucleic acid to a reporter (R) bead, passing the R bead through a sensor region including two capacitive electrodes on a substrate in communication with a signal processing circuit, and using the signal processing circuit to detect the R bead as it passes through the gap. The two electrodes may be spaced apart to form a gap such that only a single bead moves across the gap at a time. The two electrodes may form a trench through which the R bead passes. The method may further include passing the R bead through a plurality of sensor regions, and detecting the R bead as it passes through the plurality of sensor regions by applying one or more of a maximum likelihood estimation (MLE) machine learning algorithm, a partial response maximum likelihood (PRML) algorithm, or a Viterbi algorithm to signals received by the signal processing circuit from the plurality of sensor regions.

[0020] In some embodiments, detecting the target nucleic acid may include binding the target nucleic acid to a reporter (R) bead and detecting the R bead using a capacitance sensor. Detecting the R bead using a capacitance sensor may include detecting the R bead flowing past the sensor in a fluid. The fluid may be an oil. The oil may be a silicone oil. The fluid may include a fluorinated carbon. The fluid may be selected from dodecafluoro-2-methylpentan-3-one and methoxy-nonafluorobutane. The method may include fluorinating the R bead. The method may include forming a water layer around the R bead flowing in the fluid to amplify the detection signal with the capacitance sensor.

[0021] In certain embodiments, the method may include binding the target nucleic acid to the R-bead using a bead-bound peptide nucleic acid (PNA). The bead-bound PNA may include a ligand or linker PNA that does not interact with the target RNA. The method may include detecting a plurality of different targets present in the sample using one or more capacitance sensors, the plurality of different targets including at least the target nucleic acid. The plurality of different targets may include a plurality of different target nucleic acids. The plurality of different target nucleic acids may be from different pathogens or different variants of a pathogen. The pathogen may be SARS-CoV-2. The plurality of different targets may include a protein. The protein may be an antibody. The protein may be an antigen. In some embodiments, the target nucleic acid may be from a dengue virus and the antibody may be a dengue-specific antibody. In certain embodiments, the target may be a drug, including but not limited to Onpattro, Patisiran, givosiran, lumasiran, and inclisiran (siRNA), or an antibody cocktail such as REGEN-COV (casilivimab and imdevimab).

[0022] The target may be an mRNA vaccine, such as Tojinameran. Co-monitoring of therapeutic drug and viral load is advantageous, and the multiplexed nature of the method and system may enable such co-monitoring. The stability of some RNA-derived drugs is a concern both before and after use, and monitoring techniques for these concentrations or levels both from blood or saliva samples before injection and after use are beneficial. The same technology can be used for QC of these technologies in pharmaceutical companies. The target may be synthetic and may include target nucleic acid derivatives that may include non-standard bases such as Nl-methylpseudouridine. The target may also be a phospholipid or glycolipid. The target may be on the surface of a cell, a viral particle, or a lipid nanoparticle, where the cell, viral particle, or lipid nanoparticle allows for tethering.

[0023] Detecting the multiple different targets may include binding each of the multiple targets to a different bead and detecting each of the different beads using a capacitance sensor. The method may further include distinguishing the different beads based on different capacitance detection signals from the capacitance sensor. In some embodiments, the method may include selectively releasing the R bead from the substrate-bound target based on the different sensor-bound targets to which the R bead is bound, and detecting the effect on capacitance that releases the R bead.

[0024] Detecting the multiple different targets may include binding each of the multiple different targets to a different probe on the substrate. The substrate may be operably associated with one or more capacitance sensors. The method may further include binding one or more reporter (R) beads to each of the different substrate-bound targets. In some embodiments, the method may include selectively releasing R beads from the substrate-bound targets based on the different substrate-bound targets to which the R beads are bound, flowing the released R beads in a fluid past a capacitance sensor, and detecting the released R beads flowing past the sensor. Each of the different beads may have a different size, and the method may include microfluidically directing each of the different beads to a different capacitance sensor based on its size. The directing step may include one or more of inertial, dielectrophoretic, or magnetophoretic methods.

[0025] Detecting the target nucleic acid may further include binding the target nucleic acid to a probe on the substrate, and binding a plurality of reporter (R) beads to the substrate-bound target nucleic acid. The plurality of R beads may bind to different sequences in the substrate-bound target nucleic acid. The substrate may be operably associated with a capacitance sensor. The method may further include releasing a plurality of R beads from the substrate-bound target nucleic acid, flowing the released R beads in a fluid past the capacitance sensor, and detecting the released R beads flowing past the sensor.

[0026] In certain embodiments, detecting the target nucleic acid may include binding transport (T) beads to the target nucleic acid in the sample to form a T-bead complex, binding the T-bead complex to a reporter (R) bead to form an R-bead complex, eluting the R and T beads from the R-bead complex, detecting the R beads using a capacitance sensor, and returning the T beads to the sample to bind to another target nucleic acid.

[0027] Aspects of the invention may also include systems and architectures as described herein operable to perform any of the above methods.

[0028] An embodiment of the invention may include a system for nucleic acid detection, the system including a capacitance sensor further including a signal processing circuit and a sensor region including two capacitance electrodes on a substrate in communication with the signal processing circuit, the two electrodes spaced apart to form a gap, only a single bead moves across the gap at any one time, and the electrodes detect the bead and / or any molecules bound to the bead. The system may be reusable after being used to detect the bead. The capacitance sensor may be approximately 2 mm x 2 mm. The beads or particles may be contained in water, buffer, or oil or fluorocarbon, or organic solvent. The beads or particles to be detected may be a surrogate for a microbiological analyte from an upstream assay. The beads or particles may be magnetic or non-magnetic.

[0029] The gap may be twice the bead diameter. In some embodiments, the substrate may be 3D printed. The electrodes may be formed by sputtering or ink deposition and patterning. The spacing between the electrodes may be about 0.8 μm to 20 μm. The substrate may be a CMOS semiconductor chip. The electrodes may be etched in a metal layer and the spacing between the electrodes may be about 40 nm to about 5 pm. In certain embodiments, the systems and methods of the present invention may further include one or more of ultrasonic shaking, dielectrophoretic bead steering, oil syringing and / or adjustment of bead zeta potential to prevent bead aggregation and / or guide movement of particles and beads through the system.

[0030] The objective of the present invention is to make available a vital health test outside the laboratory, in a simple portable format, for use in clinical settings and at home, requiring little or no training for operation. However, the same technology can reduce the size, scale, power consumption and logistics costs of the central laboratory. A further objective of the present invention is to allow simultaneous and rapid detection of nucleic acids and antibodies from patient samples. This simultaneous detection can be advantageous in cases such as dengue outbreaks, where it is essential not only to detect and differentiate dengue virus (DENV) RNA variants 1, 2, 3, 4 (indicative of current active infection), but also to detect and identify DENV antibodies. In triage, it can help to establish whether the patient had a history of past dengue infection before making any clinical decision regarding the application of a dengue vaccine.

[0031] The systems and methods described herein recognize and address certain shortcomings in the techniques discussed above. For example, Florescu's beads are extremely large, typically 2.8 μm or larger, requiring precise manipulation of the beads onto a 4.5 μm diameter circular Hall sensor. This precise manipulation of a single particle is not practical for automated commercial NA assays. Murmann's magnetic particles are smaller (50 nm), but have a lower limit of detection of 2000 particles, making them impractical for most nucleic acid testing and gene expression analysis. The beads described in Chang and Lu are 10 μm in diameter, too large for most molecular analysis assays, requiring electromagnetic manipulation of the beads into the precise center of a coil. This is also not practical for commercial or automated assays. Additionally, the sensing mechanism described in Chang and Lu is not capacitive, as the capacitance of Chang and Lu actually decreases in the presence of beads. Thus, there is an unmet need for a reusable sensor that can detect a continuous flow of beads and detect a single bead. Additionally, there is an unmet need for a continuous flow of beads to enable single molecule analysis. There is also an unmet need for continuous flow of T-beads and R-beads to enable single molecule analysis, genotyping, and variant identification, as well as an unmet need for simple, portable means and methods for performing nucleic acid detection, variant identification, viral load quantification, and single molecule detection analysis in non-laboratory settings, such as point-of-care and community clinics, to enable rapid clinical intervention, treatment, and real-time monitoring in pandemic disease outbreak situations. [Brief description of the drawings]

[0032] [Figure 1] A beaded analyte droplet on a sputtered ITO electrode on a PET film is shown.

[0033] [Diagram 2] The capacitance-to-digital converter and the connection socket to the PET film are shown.

[0034] [Diagram 3] 2 illustrates a 24-bit sigma-delta capacitance-to-digital converter architecture and circuitry.

[0035] [Figure 4] FIG. 1 shows an ink-printed laser-ablated electrode capacitor on a 3D printed substrate.

[0036] [Diagram 5] FIG. 5 is a side view of the ink-printed laser ablation electrode of FIG. 4.

[0037] [Figure 6] FIG. 2 shows an etched aluminum electrode IDE capacitor on a CMOS silicon chip.

[0038] [Figure 7] 1 shows the formation of aluminum and copper electrode CMOS semiconductors.

[0039] [Figure 8] 1 shows a prior art analyte sensor with tethered beads and its equivalent circuit.

[0040] [Figure 9] 7 is a photograph of beads specifically captured on the circular CMOS sensor of FIG. 6.

[0041] [Figure 10] 1 shows end-to-end assay performance in a graph of capacitance versus RNA.

[0042] [Figure 11] 1 shows signal amplification by specific tethering of multiple beads to one sensor.

[0043] [Figure 12] A PNA probe targeting a conserved region of the SARS-CoV-2 virus is shown.

[0044] [Figure 13] 1 shows further PNA probes targeting regions of the SARS-CoV-2 virus.

[0045] [Figure 14] 1 shows further PNA probes targeting regions of the SARS-CoV-2 virus.

[0046] [Figure 15] Antigenic peptide sequences of the dengue envelope and NS1 proteins are shown.

[0047] [Figure 16] PNA probes targeting dengue virus variants 1, 2, 3, and 4 are shown.

[0048] [Figure 17] PNA probes targeting dengue virus variants 1, 2, 3, and 4 are shown.

[0049] [Figure 18] Shown are 30 CMOS sensors with PNA probes spotted on some of the sensors.

[0050] [Figure 19] 4 shows the spotting patterns of sensors for dengue RNA and antibody detection.

[0051] [Figure 20] 1 shows dengue RNA and antibody assays arrayed on a CMOS sensor chip.

[0052] [Figure 21] Some of the IDE capacitors in FIG. 6 are shown in rows on a CMOS silicon chip.

[0053] [Figure 22]22 is a photograph of the chip of FIG. 21 with clear silicone oil on some of the sensors.

[0054] [Diagram 23] FIG. 22 is a schematic diagram of a cylinder of oil above one of the circular sensors of FIG. 21.

[0055] [Figure 24] 24 shows the equivalent circuit of FIG. 23 with C silicone oil and C nitride.

[0056] [Diagram 25] 22 is a photograph of the bead-containing PD5 oil film on the circular sensor of FIG. 21.

[0057] [Figure 26] FIG. 26 is a schematic diagram of a cylinder of oil with beads of FIG. 25 on a circular sensor.

[0058] [Figure 27] The equivalent circuit of Figure 26 with C silicone oil, C nitride, and C beads is shown.

[0059] [Figure 28] Provides sensor measurements of PD5 oil capacitance with and without beads.

[0060] [Figure 29] Shown is a 0.4 μm bead in a 0.8 μm trench between metal electrodes.

[0061] [Diagram 30] A simulation of the electric field of this bead in oil passing between electrodes.

[0062] [Diagram 31] 1 shows the calculated capacitance of a bead as it passes between the electrodes.

[0063] [Diagram 32]FIG. 32 shows the boundary element model used in the simulations of FIGS. 30 and 31.

[0064] [Diagram 33] 29 shows the ground plane beneath the electrode of FIG.

[0065] [Diagram 34] The bead capacitance calculations in Figure 33 are shown with the ground plane down.

[0066] [Diagram 35] 1 shows the real-time capacitance of a series of beads flowing between electrodes.

[0067] [Diagram 36] Beads are shown flowing between multiple electrodes and then down holes in the substrate.

[0068] [Figure 37] Shows the ink electrodes and laser-formed through-holes in the 3D printed substrate.

[0069] [Figure 38] 1 shows ink-printed electrodes with narrow gaps formed by laser ablation.

[0070] [Figure 39] 1 shows an assay according to a particular embodiment involving recycling of T-beads.

[0071] [Diagram 40] For example, an embodiment is presented that allows single molecule analysis of SARS-CoV-2.

[0072] [Diagram 41] 1 illustrates an embodiment of a home respiratory saliva self-test for COVID / Influenza / Hay Fever.

[0073] [Diagram 42] Fluorinated reporter beads are shown flowing through a fluid phase transition.

[0074] [Diagram 43] 1 shows a microfluidic embodiment for flowing beads across a phase transition.

[0075] [Diagram 44] Shown is an R bead tethered to a D bead that provides additional amplification.

[0076] [Diagram 45] Shown is the R bead-D bead tethering complex flowing between the sensors.

[0077] [Figure 46-1] (a) CMOS sensor, (b) RD bead complex tethered to a bead substrate. [Figure 46-2] Same as above.

[0078] [Figure 47] Shown are reduced size R nanoparticles (1-100 nm) tethering D beads to a surface.

[0079] [Figure 48] 1 shows an R nanoparticle having a linker-type molecular shape.

[0080] [Figure 49] A linker that is a ligand that does not interact with the target RNA is shown.

[0081] [Figure 50] Shown is R nanoparticles tethering D beads to the surface of magnetic particles. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0082] Detailed Description The present invention provides an improved capacitive bead sensor that can be easily manufactured and automated, and can be reused for multiple bead tests or continuous flow of beads. The system and method described herein allows detection down to a single bead, and therefore a single target molecule. The electronic and other signal amplification techniques discussed below allow the sensitivity to be achieved without enzymatic amplification, thereby avoiding some of the disadvantages of enzymatic methods as described above. Furthermore, the system and method described herein can achieve the desired sensitivity while maintaining a wide dynamic range, which is particularly useful in multiplex analysis of target molecules at various concentrations in a sample.

[0083] As discussed in more detail below, the systems and methods described herein may enable, among other things, one or more of: a) simultaneous capture and detection of nucleic acid molecules, antibody molecules, and antigen molecules from a biological sample; b) quantification of these molecules down to single digit copy levels (e.g., for clinical viral load determination); c) gene expression analysis by simultaneous quantification of multiple PNA captured genes on multiple capacitive bead sensors; d) single molecule detection analysis with multiple PNA probes on multiple beads recycled through the sample and assay steps; and e) discrimination between current and past infections by simultaneous detection of both RNA and antibodies from the same sample. Bead detection and signal amplification

[0084] Figure 1 shows six sputtered indium-tin oxide (ITO) capacitance electrode sensors on a 150 μm thick polyethylene terephthalate (PET) film. Each capacitor is approximately 2 mm × 2 mm. A connector socket (on the right) connects the capacitors to a signal processor, which may include a capacitance-to-digital converter, analog and digital filters, non-volatile memory, signal processor, and wireless communications (Bluetooth, WiFi, GSM, GPRS, etc.). The capacitances shown (4.45 pF, 4.587 pF, 4.59 pF) are baseline values ​​without liquid or beads. The slight value increases (S1-S3-S5) reflect longer wires to reach each capacitor. These baseline values ​​can be stored as offsets in the non-volatile memory of the signal processing circuit to allow nulling of all sensors to a zero baseline by offset subtraction from each of the associated sensor readings (as shown in Figure 28). The sensors are arranged in a straight line to facilitate integration with the droplet dispensing chip for automation of the assay in some applications.

[0085] FIG. 2 shows capacitance-to-digital converters on a printed circuit board (PCB) and the small sockets that connect them to the six capacitors of FIG.

[0086] FIG. 3 shows the circuit architecture of a sigma-delta capacitance-to-digital converter as known in the art, e.g., "Oversampling Delta-Sigma Data Converters" (Candy et al., Wiley-IEEE Press, 1992). Using standard switched-capacitor techniques operating at modulator clock frequencies in the KHz to MHz range, センサ C 基準 The difference is amplified and integrated and then applied to a comparator. The comparator output is V 基準The comparator bitstream output is applied to a decimation filter to produce a digital result with up to 24-bit resolution. Typical conversion times are 1mS to 100mS. The 1 million-fold electronic amplification provided by a multi-stage integrating amplifier can contribute to the overall assay amplification.

[0087] Figure 4 shows a 3D printed polylactic acid (PLA) substrate with interdigitated electrodes formed using ink printing and laser ablation patterning. The measured baseline capacitance (including finger and inter-wire parasitics) is in the region of 4.5 pF to 6.1 pF. The fringe electric field sensing or transduction portion of the capacitance may be in the region of 0.4 pF to 1.6 pF depending on the electrode spacing, which may be typically 0.8 μm to 20 μm depending on the spot size of the ablation laser.

[0088] Figure 5 is a side-view photograph of the ink electrodes in S4 of Figure 4. Laser ablation of the ink to create the electrode structures results in trenches between the electrodes, with ink thickness heights programmable from a few microns to hundreds of microns during 3D printing. Electrohydrodynamic inkjet printing is another method to create fine-line (e.g. 1 pm) electrode features without the need for laser ablation.

[0089] Figure 6 shows an interdigitated electrode (IDE) capacitor formed by subtractive etching of the metal top layer of a complementary metal oxide semiconductor (CMOS) chip. In this embodiment, a sigma-delta capacitance-to-digital converter may be formed in the same CMOS chip, below or adjacent to the sensor. This significantly reduces parasitic capacitance and noise (thermal and electromagnetic), thereby providing a wider dynamic range of measurements.

[0090] A cross-sectional view of the metal layers forming a CMOS sensor is shown in FIG. 7 (an illustrative example from Burghartz et al., IEEE 2004, DOI 10.1109 / TED.2004.823325). This reference also describes typical CMOS process formation steps and is incorporated herein by reference in its entirety. Electrode spacing as low as 20 nm can be achieved with modern dry plasma etching and highly directional anisotropic deep reactive ion etching (DRIE) methods of semiconductor manufacturing. The dashed line (a) in FIG. 7 shows the narrow space in the top metal that, in certain embodiments, can be used as a trench for beads to flow between without passivation. This can be advantageous to increase assay amplification while eliminating a fabrication step (passivation) to achieve tighter spacing. If the beads are covered with silicone oil that can flow, there is no oxidation or corrosion of the metal. The dashed line (b) in FIG. 7 is a variation using copper plating for much thicker metal structures, similar to the 3D printed wrench in FIG. 5. This copper plating example from Burghartz is for RF inductor coil formation. In certain embodiments, the system of the present invention can use this standard bulk RF-CMOS thick metal technology as electrode trench structures for measuring beads in trenches and for measuring bead flow in oil, thus avoiding the requirement for special CMOS layers or non-standard semiconductor processing.

[0091] 8 shows a capacitive bead detection method (see U.S. Pat. No. 10,746,683, which is incorporated herein by reference in its entirety). The bead can simply arrive at the sensor surface (8b) or be specifically tethered to the sensor surface (8c) by, for example, probe-0, which in turn is tethered by a further probe-1 attached to the bead. The capacitive fringe field (electric field 8b dotted line) is a function of C due to the dielectric constant of the bead material. ビーズ The equivalent circuit is shown in Figure 8(a). Since the field lines pass through the insulating nitride passivation layer, C 窒化物 C ビーズ and appear in succession.

[0092] Figure 9 is a photographic representation of Figure 8(c) tethering, in this case showing a reporter bead (R bead) specifically captured at the center of the circular sensor of Figure 6 by a PNA probe designed to target the HIV gag gene. These are proxy beads for HIV-RNA in an upstream assay from a patient sample, so the sensor capacitance, which is a quantification of the number of beads, is also a direct digital readout of the HIV viral load in the patient sample. This is shown in the end-to-end assay performance graph of "Capacitance vs. RNA" in Figure 10. The top photo shows beads specifically captured on the sensor surface. The bottom photo shows no beads being captured when using an off-target PNA probe.

[0093] Figure 11 shows a target RNA strand tethered to the sensor surface (by probe-0), with several beads tethered to the RNA by probes 1, 2, 3, 4, etc. These probes can be designed to target different conserved regions of the target DNA or RNA, thus giving very high specificity of target capture. It also results in significant signal amplification due to the increased number of capture beads tethered on the sensor. This can mimic dendritic amplification and aid in overall assay amplification to quantify DNA or RNA even at very low copy number levels without enzymatic amplification. Nucleic Acid and Antibody Assay Peptide and Probe Design

[0094] The sequence of the HIV Gag gene PNA probe in Figure 9 has been published by Zhao et al. in Nature Communications (nComms 5079), the contents of which are incorporated herein in their entirety. Fully synthetic PNA probes can be chemically synthesized by FMOC chemistry using solid phase peptide synthesis (SPPS).

[0095] Figure 12 shows a 14 base PNA probe targeting a unique region of the SARS-CoV-2 viral genome conserved across the Wuhan reference sequence and the alpha, beta, gamma, delta and omicron BA.1 and BA.2 variants. This probe may be tethered to a sensor surface, for example as probe-0 in Figure 11.

[0096] Figures 13 and 14 show additional PNA probes (probe 6, probe 7) designed to target regions of the SARS-CoV-2 receptor binding domain (RBD) to distinguish the Omicron BA.1 and BA.2 variants from the alpha, beta, gamma, and delta variants. Probe 8 further qualifies and identifies the BA.2 variant by targeting its S gene dropout deletion. As known in the art, additional probes can be similarly applied to identify other variants, where variant identification and virus genotyping can be achieved with a small number of carefully designed probe sequences (e.g., by applying a logic table such as that shown in Figure 41). For example, in 2004 "A universal microarray for detection of SARS coronavirus" (doi:10.1016 / j.jviromet.2004.06.016), Long et al. published 20 probe sequences that allowed "not only the detection of SARS-CoV but also the identification of the genotypes of six mutated bases associated with different phases of the SARS epidemic" (2003 / 2004).

[0097] Figure 15 shows synthetic peptide sequences of the envelope (E) and nonstructural (NS-1) proteins of the Dengue Virus (DENV) (Nagar et al., DOI10.1155 / 2020 / 1820325). These can also be chemically synthesized by FMOC SPPS. They are based on linear epitopes of DENV envelope and nonstructural proteins that are used to test and diagnose dengue-specific IgM and IgG antibodies in patient antisera. The nine sequences shown are from among the "most antigenic / reactive" to provide highly specific antibody detection (by Nagar et al.).

[0098] Figures 16 and 17 show an example of achieving 100% detection of the four dengue reference genomes NC_001477, NC_001474, NC_001475, NC_002640 (DENV1, 2, 3, 4) using just two PNA probes (14 bp and 15 bp) targeting the conserved regions shown. This is not possible with longer PCR probes and primers, which are typically >25 nt and would therefore span some of the mutations shown, resulting in less than 100% specificity, and lower capture efficiency. The high capture efficiency of PNAs may also contribute to the overall assay signal amplification and therefore sensitivity.

[0099] Multiplex Testing Figure 18 is a photograph showing 30 circular CMOS sensors, as shown individually in Figure 6, grouped together with PNA tethering probes spotted on various sensors. These probes can be seen in the bead circle in Figure 9 and can specifically capture reporter R beads, as shown in Figure 11. Thirty sensors allow for a much higher degree of multiplexing and genotyping than is possible with LAMP or PCR assays. This is shown in the exemplary embodiment of Figure 19, where 12 sensors are spotted with duplicates of six R bead DENV PNA probes (PNA1-6). These include two PNA probes of Figure 16 / 17 for 100% dengue detection and one PNA probe each for DENV1 / 2 / 3 / 4 mutation identification. For highly specific DENV antibody detection, nine sensors are spotted with the E and NS1 peptide sequences of Figure 15: EP1, EP2, EP4, EP7, EP10, EP12, NS1-1, NS1-3, NS1-4. The bead quantification capabilities of each sensor can be used to establish E and NS1 quantities. Combined with the machine learning capabilities of the CMOS chip electronics and processor, this can aid in improving assay sensitivity and selectivity, as well as predicting the potential onset of "severe dengue" and the severe and potentially fatal dengue hemorrhagic fever (DHF) and dengue shock syndrome (DSS). Two PNA positive control sensors (both spotted with 18s rRNA complementary probes for blood samples, or RNAse P complementary probes for saliva samples) are included. Two "PNA-neg" negative control sensors are included, e.g., for single base mismatch confirmation test controls. Sensors spotted for IgM and IgG positive controls, as well as aG negative antigen controls are also included. Two blank sensors for temperature and CMOS process variation correction and offset removal are used.

[0100] This highly multiplexed and nucleic acid / antibody-specific simultaneous detection capability of this assay is implemented by 30 co-located sensors, as shown in Figure 20, and parallel fully synthetic peptide sequences and peptide-nucleic acid (PNA) probes, each synthesized by the same FMOC chemistry.

[0101] In certain embodiments, the multiplexed detection may be performed using flow-through detection. For example, each of the above spots may not be a sensor. Instead, the tethering of the R-beads to the substrate may be mediated by a carefully designed PNA linker or ligand sequence that does not interact with the RNA in the sample. These are conjugated to a PNA probe or an antigenic peptide sequence that targets the RNA. A careful design of the linker PNA allows for continuous and selective elution of the R-beads from the substrate as the system is heated and the melting temperature of the linker PNA is reached. In this way, the R-beads attached to the spots mediated by the RNA target can be selectively and continuously released and subsequently measured by a downstream sensor. Furthermore, the R-beads attached to the spots mediated by the antibody can be selectively released and subsequently measured by a downstream sensor when the Tm of the associated linker PNA is reached. The sensor may be an in-line sensor that detects the R-beads passing by them. In certain embodiments, the sensor may be a PNA functional sensor that binds to a specific R-bead. Alternatively, the substrate may be the surface of a magnetic bead. Detection of beads flowing in silicone oil:

[0102] FIG. 21 shows another embodiment of the sensing capacitor of FIG. 6 on a CMOS silicon chip. In this embodiment, a sigma-delta capacitance-to-digital converter is formed in the CMOS chip to the left of the sensing capacitor. Multiple diode temperature sensors integrated across the chip ensure that differences in sensor readings due to temperature variations across the chip can be measured and compensated for. The sensors are in an offset row and column formation to facilitate integration with droplet dispensers for some assay automation applications, or multi-channel assay methods where, for example, the upstream nucleic acid and antibody assays may be in different channels.

[0103] Figure 22 is a photograph of the chip from Figure 21 with a film of transparent PD5 silicone oil on several sensors of various diameters on the right half of the chip. The film of oil is located in the fringe field transduction portion of the IDE electrode capacitive sensor and therefore increases the sensor capacitance due to the oil dielectric constant K of about 2.5-3, which is higher than air (K=1).

[0104] FIG 23 is a schematic diagram of a cylindrical portion of an oil film being sensed on one of the circular sensors. FIG 24 shows the capacitance of the insulating nitride passivation layer, C 窒化物 and the oil capacitance C シリコーンオイル FIG. 23 shows an equivalent circuit of the circuit shown in FIG.

[0105] Figure 25 is a photograph of a PD5 oil film containing titanium dioxide (TiO2) beads on one circular sensor. These titanium dioxide analyte beads in the oil film add additional capacitance because their dielectric constant is about 30 to about 100, which is much higher than the oil dielectric constant (2.5 to 3). In fact, the K of TiO2 can be 10 times higher (up to K = 1000) at lower frequencies (100 Hz, according to Wypych et al., DOI 10.1155 / 2014 / 124814). Thus, an additional 10-fold assay amplification is possible by slowing down the sigma-delta converter modulator frequency at the expense of a slower conversion time of up to 1 second.

[0106] FIG. 26 is a schematic diagram of FIG. 25 showing a cylindrical portion of beaded oil on a circular sensor.

[0107] FIG. 27 is an equivalent circuit for the analyte sensor of FIGS. 25 and 26, where C 窒化物 Paired with related C シリコーンオイル and C ビーズ Indicates capacitance.

[0108] Figure 28 shows sensor measurements of oil capacitance with and without beads. At time zero, all sensors are baselined to zero with air dielectric. After 2 minutes, a PD5 oil film (without beads) is applied to sensors S0 and S30 (as a negative control), and after 1 minute, PD5 oil containing 1 μm TiO2 beads is applied to S6 and S25. As can be seen at the top of the graph, the beads cause a 7.6 fF increase in capacitance. Visual counting of the beads under a microscope (as in the photograph in Figure 25) gives an estimate of about 6000 beads, i.e., about 1.26 aF per bead. This corresponds to about 12 electrons for a switched-capacitor voltage of 1.6 V applied between the electrodes by the modulator of the sigma-delta converter.

[0109] Referring to the S0 and S30 negative control sensors (0.96 mm diameter), the addition of silicone oil at 2 minutes increases the capacitance by 132 fF. For sensors S5 and S7 with diameters of 0.54 mm and 0.3 mm, the capacitance increases by 45 fF and 12.5 fF, respectively. These measurements are summarized in Table 1 below and show that regardless of sensor diameter, the silicone-oil film increases by approximately 183 fF / mm2 compared to the air dielectric. Table 1 [Table 1]

[0110] These measurements of beads in oil illustrate various embodiments of the systems and methods of the present invention. They can be used to eliminate the need for liquid evaporation, which is advantageous in allowing the reuse of the sensor for multiple measurements. It can allow for liquid flow control and movement of beads by syringe, which is useful for detecting non-magnetic beads and particles, or for magnetic particles. The graph shown in FIG. 28 also shows the change in capacitance as the oil film thickness changes and as beads flow out of the sensor, demonstrating dynamic real-time bead flow detection capabilities. Electrical characterization of single beads and bead flows in oil:

[0111] FIG. 29 shows a 0.4 μm bead in a 0.8 μm trench between electrodes.

[0112] Figure 30 shows a COMSOL electrostatic simulation of these electrodes with 1 volt applied and this bead (K=100) in oil (K=2.5) passing between the electrodes. The electric field gradient lines are shown with 0.1 V steps.

[0113] FIG. 31 shows the COMSOL capacitance calculations of a bead as it passes between the electrodes for bead dielectric constants of 2.5, 10, 30, 100, and 1000. The capacitance peaks at 4 attofarads (aF) for a 0.4 μm bead (K=100) in oil as it passes through a 0.8 μm electrode gap. This is because the capacitance is higher for the 0.4 μm bead than for the 1.2 aF for the larger 1 μm TiO2 bead (from the graph in FIG. 28) due to the smaller electrode spacing and stronger electric field across the bead. This may further contribute to the overall signal amplification of this nonenzymatic detection / quantification assay. The increased electric field across the bead increases the dipole polarization of the molecules within the bead, thereby increasing its capacitance. The baseline capacitance of 399 aF is quite low in this simulation, due in part to the short electrodes in the model with no other wiring or ground planes nearby, as shown in the boundary element model (BEM) in FIG. 32 used in this simulation.

[0114] In practice, the electrodes may have traces several millimeters long connecting them to a capacitance-to-digital converter and / or may be formed on a physical substrate, e.g., a 3D printed base, or a CMOS silicon substrate. Electrically these substrates may appear as a large ground plane. This is shown in FIG. 33, where a 0.4 μm bead passes through a 0.8 μm electrode gap, with a 2 μm ground plane underneath, as may be typical for a CMOS silicon substrate. For K=100 beads in oil (K=2.5), the COMSOL bead capacitance is again 4 aF, as shown in FIG. 34. However, the baseline capacitance in FIG. 34 is 4.274 femtofarads (fF) (i.e., more than 10 times higher than FIG. 31) due to increased parasitic capacitance to nearby ground planes. Longer traces (e.g., 10 mm to 15 mm) can increase this baseline even further, e.g., 4.5 pF to 6.1 pF in the embodiment shown in FIGS. 1 and 4. It is noteworthy that a single bead capacitance of 4 aF is detectable even with these huge baseline variations from attoFarads (aF) to picofarads (pF). This is due to the wide dynamic range of the Sigma-Delta 24-bit capacitance-to-digital converter. This embodiment has a full-scale range of 8 pF, a resolution of 0.5 aF, and an accuracy of 4 aF. An additional internal offset capacitor is added to the C 基準 (Figure 3), the baseline range can be extended, for example, up to 17pF or 20pF. This makes it possible to manufacture capacitive sensors in many different ways, even with large parasitics, on silicon substrates or CMOS chips, with or without ground planes, by 3D printing.

[0115] FIG. 35 shows a time domain graph of the converter output for very low attofarad level measurements. A large amount of heat and 1 / F noise is evident. Beads passing between the electrodes can appear as short "blips" (tips circled in the waveform) roughly corresponding to a capacitance signal of 3aF to 6aF. Identifying these blips is not straightforward. Simple peak detection can miss ambiguous blips (exemplified by the peaks marked with "?" in FIG. 35). In certain embodiments, moving average filtering can be used to aid in identification, but can miss short blips. Thus, in some embodiments, multiple electrodes can be used on the bead channel to help read the same bead multiple times as shown in FIG. 36. A maximum likelihood estimation (MLE) machine learning algorithm can then be applied to the reads to improve bead counting accuracy. Partial response maximum likelihood (PRML) and Viterbi algorithms further improve bead identification and reduce error rates. These digital filtering improvements in signal to noise ratio, which can themselves be embedded within the CMOS sensor chip, can further contribute to overall assay amplification without enzymatic amplification.

[0116] Many amplification methods and steps are described in this disclosure: PNA high capture efficiency, bead mass, CMOS 10 6 Electronic amplification, multi-bead tethering signal amplification, increased bead electric field in the trench, MLE and PRML / Viterbi digital signal processing. The combined effect of these was 10 9 ~10 12The goal is to reach an amplification of 100 copies / mL, which is roughly equivalent to 30-40 PCR cycles, respectively. Thus, using the fully synthetic assay described herein, analytical sensitivity levels of <100 copies / mL can be achieved, but with a wide dynamic range of several orders of magnitude, ideal for clinical RNA viral load monitoring. The assay of the present invention can further eliminate the requirement for enzymatic amplification, and therefore also eliminates many of the inhibition issues and complexities of PCR and LAMP assays. Being fully synthetic, the assay also eliminates the requirement for dry ice transportation and storage in a refrigerator or freezer. The long shelf life, potentially up to several years, as in electronic devices, is another possible advantage compared to enzymatic assays, where reagents may degrade more quickly and without special handling. Through-hole bead flow assay method

[0117] Figure 36 shows beads flowing along channels or trenches on a substrate between multiple electrode pairs and then down holes in the substrate and out the back, in some embodiments this can be DRIE etched silicon through holes as described in WO 2004 / 109770.

[0118] Figure 37 shows ink electrodes and exemplary laser-formed through-holes in a 3D printed PLA substrate, which can be advantageous for in-line bead measurements where beads can flow within a tube or channel.

[0119] FIG. 38 shows a narrow 0.8 μm electrode spacing between A and B ink printed electrodes, formed by laser ablation with a 0.8 μm laser spot size. The beads then flow in the channel and "drop" through holes between the electrodes (created by laser ablation of the 3D printed base). Smooth bead flow can be enhanced by syringe handling push / pull control of the oil or other liquid flow carrying the beads, and / or by dielectrophoretic steering, where charged beads (e.g., zeta potential of -30 mV) respond to positive or negative voltages applied externally or along the flow path, e.g., by the + / - tip shown in FIG. 38. Bead aggregation can be minimized by adjusting the zeta potential of the beads, e.g., -30 mV, by a carboxylate coating that can repel the beads from each other, and / or by ultrasonic shaking of the device to aid monodisperse flow through the sensor.

[0120] To provide multiplexing capabilities to the flow-through detector system, multiple flow-through sensors can be placed in the same device. Several separate groups of differently sized R beads can be used, corresponding to each variant to be tested. These R beads may be spatially separated to allow each size group to be delivered to a different through-hole sensor. The method of separation can be by inertial separation, or dielectrophoretic or magnetophoretic separation. The beads can thereby be differentiated, allowing differential detection of each target on each bead in a multiplex assay. Recycling beads

[0121] In certain embodiments, it may be advantageous to recycle beads through the system depending on the workflow. An exemplary bead recycling scheme is shown in Figure 39. The system 3901 has four main areas: sample capture 3903, assay assembly 3905, elution 3907, and sensor detection 3909. Depending on the sensing method selected, recycling of T beads may be advantageous as T beads are continuously supplied until the targeted RNA is depleted (Figure 39). In 3903, T beads capture the target RNA via the T probe. R probes and R beads are added to the next stage of assay assembly at 3905, at which point the complete assay is magnetically transported to an elution point 3907 on the sensor. The eluted R beads are then attached to the sensor 3909 via specific PNA-PNA binding. Over time, the number of R beads accumulates on a particular sensor until the source target RNA is no longer available. The final (total) signal is then captured. As shown in Figure 39, after elution step 3907, the T-beads can be recycled 3911 and returned to the initial capture step 3903. Before returning the T-beads to the initial capture step 3903, they can be washed to ensure that the beads are free of false probes, R-beads or RNA.

[0122] If the CMOS detector is a single bead continuous flow sensor as described above, without sensor surface binding, assay assembly and elution are performed similarly with T beads recycled after elution. In such an embodiment, the R beads and T beads are bypassed after elution, with the R beads passing through the single bead CMOS detector and the T beads returning to the sample in the initial capture step until the target RNA is depleted. The signal is then read at each event to effectively count the number of beads that have passed through the CMOS detector. Single molecule analysis:

[0123] As discussed throughout, various signal amplification techniques can be used alone or in combination to increase sensitivity, for example, for single molecule detection. Single molecule detection has many applications, including early detection of infection and accurate quantification of viral load, for example, in SARS-CoV-2 outbreaks. Figure 40 shows an embodiment for single molecule analysis, for example, of SARS-CoV-2.

[0124] In early 2022, Lai et al. (DOI: 10.1101 / 2022.01.08.22268865) published a list of 48 markers or nucleotide mutations in the SARS-CoV-2 genome that correspond to the variants known at that time: alpha, beta, gamma, delta, lambda, mu, epsilon, iota, eta, kappa, and early omicron variants. Assuming a "brute-force" set of 48 probes targeting each of these mutated regions in the genome, probe-1 is a capture probe designed to uniquely target all SARS-COV-2 variants, for example as shown in Figure 12. Probe-1 can be bound to paramagnetic transport (T) beads in a continuous flow assay to capture and extract SARS-CoV-2 RNA from the sample. R beads can also be flowed sequentially into the tethering chamber, first with probe 2, then probe 3, then probe 4, etc., in groups with different probes attached, up to probe 48. Variants in patient samples can then be identified by search and refinement analysis, as shown in part in Figures 13 and 14. In some embodiments, the number of probes required can be reduced by applying a logical reduction analysis of "truth tables" to identify variants, as shown in Figure 41 with the mapping of L452R, T478K, N501Y, and H655Y mutations to identify the Omicron variant.

[0125] FIG. 41 shows an assay embodiment for home respiratory testing from a self-testing saliva sample. In addition to the COVID-19 detection and SARS-CoV-2 variant identification described above, separate probes are added to identify and identify influenza A, influenza B, RSV and other viral pathogens, as well as peptide sequence probes of epitope antigens for detecting immunoglobulin E (IgE) antibodies to identify allergic rhinitis ("hay fever") that may cause sneezing and stuffy nose attacks. The assay sensitivity down to single molecules of the embodiments discussed above, the lack of enzyme amplification (which can complicate home use), and multiplexing across classes of targets can all be combined to enable such testing.

[0126] Detection of beads flowing in other low dielectric constant liquids: Other examples of low dielectric constant liquids are dodecafluoro-2-methylpentan-3-one (Novec 649), and fluorinated carbons such as methoxy-nonafluorobutane (Novec 7100). These are clear, colorless, low odor liquids used as advanced heat transfer fluids with desirable environmental and electrical properties. Novec 649 is commonly used for cooling electronics. It has a very low dielectric constant (@1kHz) of 1.8, making it an advantageous carrier medium for capacitive sensing of beads. Additionally, it is denser and less soluble in water, and as a fluorinated material, it will fractionate from both aqueous and oil-based liquids. In certain embodiments, beads released into the flow past the sensor can be monodispersed within the liquid to aid in assay sensitivity and single molecule detection. In some embodiments, phase transitions can be used since the sample type is typically aqueous.

[0127] FIG. 42 is a schematic diagram for making fluorinated R beads at the interface between aqueous and fluorinated solvents according to certain embodiments. Beads can be modified with ligands that allow for migration across the interface. The interface shown in FIG. 42 is followed by beads modified to fluorinated R beads, for example with heptafluorobutylamine (demonstrated with Dynabeads), and beads modified with fluoro-PNA (e.g., Fluor-PNA2 forms micelles and / or aligns at the aqueous / fluorinated fluid interface). This scheme shows how biphasic soluble PNAs (PNA and fluorinated ends) can be used to tag aqueous-dissolved R beads, allowing for the subsequent migration of the R beads to fluorinated solvents. Similar chemistry is available to manufacture lipophilic beads for migration to oil. In some embodiments, R beads can have amphiphilic properties and can be monodispersed in both aqueous and organic or fluorinated solvents. Additionally, the use of amphiphilic surfactants dissolved in one or both solvents can facilitate migration of beads between solvent phases. This approach has been demonstrated by Wei et al. (DOI: 10.1021 / ja039874m, incorporated herein by reference in its entirety) for the transfer of gold nanoparticles from an aqueous phase into an immiscible ionic liquid.

[0128] In certain embodiments, for example, the transfer of beads to the final fluorinated solvent may be by transfer between multiple solvents (e.g., from aqueous to aprotic / organic, fluoroether / ester / ketone, fluorinated or low dielectric constant solvents). Such transfer can be performed using a microfluidic approach as shown in Figure 43. Fluidics can be used to introduce different solvents that overlap for mixing or micelle formation as the beads move through the solvent phases. This approach may also facilitate the use of phase transfer molecules or amphiphilic surfactants in each of the separate solvents. This approach is guided by the Hansen solubility parameters of the beads and the solvent to obtain sufficient solubility of each to facilitate the bead phase transfer. Similar methods can be used in microfluidic bead fabrication to transfer beads from an organic solution to an aqueous solution.

[0129] Another way to move the R beads from one solvent to another could be to use their high dielectric constant or high zeta potential to deflect the R beads from the "sample" or aqueous flow into a parallel flow of Novec 649 or other fluid, where the beads move laterally across the flow line until they enter Novec 649 and are deflected to an outlet with a sensor.

[0130] In some embodiments, magnetically sensitive R beads can be deployed in the device.Magnetic phoretic separation can be used to move beads in fluid, as shown in the literature (Pamme&Manz, 2004, DOI 10.1021 / ac049183o, the entirety of which is incorporated herein by reference).Magnetic phoretic separation can be used to deflect R beads from aqueous phase to Novec 649 or oily or other phase.

[0131] The transition of the beads from the aqueous phase to the oil phase allows for the formation of a thin shell layer or droplet of aqueous solution around the beads while they are in the oil phase. In some embodiments, flowing this aqueous coated bead through a capacitance sensor can provide an amplified signal over that of standard beads due to the high dielectric constant of the aqueous shell layer on the beads, thereby further increasing the sensitivity of the system. Combining tethering with bead flow assays:

[0132] In the embodiment of FIG. 44, the R beads eluted from the upstream assay may then bind to further beads to amplify the signal ("D beads", again mimicking dendritic amplification). This can be done using PNA-PNA bonds between the beads. This allows the R beads to be made smaller, allowing more of them to be attached to single molecules (e.g. RNA). FIG. 45 shows an R bead-D bead complex through a flow path, where two CMOS sensors are placed on top of each other to partially or fully form a flow path, and the capacitance signal is measured by both chips independently as the bead passes each sensor. The measurement by one sensor is digitally compared to the measurement from the other sensor. The same arrangement can be used to measure single beads.

[0133] R beads can be used to mediate the capture of R bead and D bead complexes on a substrate. The complexes can have properties that mediate their distinction from uncomplexed beads. These can include capture of the complexes on a substrate (e.g., mediated by R bead PNA) that can complete an assay where the substrate is the surface of the sensor. The properties of the complexes can include PNA-PNA binding, dielectrophoresis, or electrophoresis. This property can attract the complexes to the sensor. This property can cause the complexes to deviate in a flow path where the complexes pass through the sensor and the uncomplexed particles flow through an outlet. This property can temporarily hold the complex or elements of the complex on the substrate while the uncomplexed particles are washed and / or removed through an outlet. For example, if PNA-PNA that binds to the substrate is used, after the uncomplexed particles are removed through an outlet, the temperature of the system can be increased to elute one or other members of the complex from the substrate, which then pass through the sensor and can be measured as described elsewhere herein. The capture and subsequent release of multiple beads can achieve the aforementioned tethering signal amplification in flow-through embodiments.

[0134] In certain embodiments, the R beads can be reduced to a single PNA, and binding of this PNA to another substrate can be used to mediate capture of the D beads. The substrate can be a sensor chip (i.e., functionalized with PNA for PNA-PNA capture) as shown in Figure 46a, or the substrate can be additional magnetic beads as shown in Figure 46b. This can provide for magnetic bead tethering and manipulation, taking advantage of the very strong PNA-PNA binding within each member of the tethered complex. Thus, when using a flow-through sensor, the advantage of signal amplification by tethering can still be obtained.

[0135] Figure 47 is an example of much smaller R beads (1-100 nm) tethering D beads to a surface.

[0136] Figure 48 shows that the R nanoparticles can have other shapes, or can be linker-type molecules that allow attachment of one or more probes, or can be long PNA probes with two regions. Probe 2 can be a probe or ligand that is also functionalized on the R nanoparticle but does not interact with the native RNA target. This can be achieved by careful probe design to avoid sequences found in the native RNA target and avoid interactions with the PNA probe. The ligand can be such that it does not interact with normal base pairing on the target RNA, including but not limited to biotin-streptavidin, as shown in Figure 49.

[0137] Figure 50 shows small R nanoparticles (1-100 nm) tethering D beads to the surface of magnetic particles for further manipulation before sensing. Incorporation by Reference

[0138] References and citations to other documents, such as patents, patent applications, patent publications, journals, books, articles, web content, etc. are made throughout this disclosure. All such documents are incorporated herein by reference in their entirety for all purposes.

[0139] Equivalent Various modifications of the invention and many further embodiments thereof, in addition to those shown and described herein, will become apparent to those skilled in the art from the entire contents of this document, including references to the scientific and patent literature cited herein. The subject matter herein contains important information, exemplification, and guidance that can be adapted to the practice of this invention in its various embodiments and equivalents.

Claims

1. 1. A method for nucleic acid detection, the method comprising detecting a target nucleic acid present in a sample at or below 100 copies / mL without enzymatic amplification using a capacitance sensor.

2. In detecting the target nucleic acid, at least ９ 2. The method of claim 1, operable to achieve a signal amplification of

3. 10. The method of claim 1, operable to achieve signal amplification in the detection of the target nucleic acid equivalent to at least 30 PCR cycles.

4. 10. The method of claim 1, wherein the nucleic acid detection by a capacitive sensor comprises a full scale range of at least about 8 pF.

5. The method of claim 4 , further comprising converting a signal from the capacitive sensor using a capacitance-to-digital converter.

6. The method of claim 5, wherein the capacitance-to-digital converter is a sigma-delta 24-bit capacitance-to-digital converter.

7. 5. The method of claim 4, wherein nucleic acid detection by a capacitive sensor comprises a resolution of at least about 0.5 aF and an accuracy of at least about 4 aF.

8. 5. The method of claim 4, wherein the nucleic acid detection by a capacitive sensor includes a full-scale range of at least about 15 pF by including an internal offset capacitor at the reference input.

9. Binding the target nucleic acid to a reporter (R) bead; passing the R bead through a sensor area including two capacitive electrodes on a substrate in communication with signal processing circuitry; The two electrodes are spaced apart to form a gap; passing the R bead through the gap, where only a single bead moves between the gaps at a time; detecting the R bead as it passes through the gap using the signal processing circuitry; The method of claim 1 further comprising:

10. passing the R beads through a plurality of sensor regions; Detecting the R beads as they pass through the plurality of sensor areas by applying one or more of a maximum likelihood estimation (MLE) machine learning algorithm, a partial response maximum likelihood (PRML) algorithm, or a Viterbi algorithm to signals received by the signal processing circuitry from the plurality of sensor areas; The method of claim 9 further comprising:

11. Detecting the target nucleic acid Binding the target nucleic acid to a reporter (R) bead; Detecting the R beads using a capacitance sensor; The method of claim 1 , comprising:

12. The method of claim 11 , wherein detecting the R beads using the capacitance sensor comprises detecting the R beads flowing past the sensor in a fluid.

13. The method of claim 12 , wherein the fluid comprises oil.

14. The method of claim 13, wherein the oil is a silicone oil.

15. The method of claim 12 , wherein the fluid comprises a fluorinated carbon.

16. The method of claim 15, wherein the fluid is selected from dodecafluoro-2-methylpentan-3-one and methoxy-nonafluorobutane.

17. The method of claim 14 further comprising fluorinating the R beads.

18. The method of claim 13 , further comprising forming a water layer around the R beads flowing in the fluid to amplify a detection signal in the capacitive sensor.

19. 13. The method of claim 12, further comprising binding the target nucleic acid to the R bead using bead-bound peptide nucleic acid (PNA).

20. 20. The method of claim 19, wherein the bead-bound PNA is a ligand linker PNA.

21. 10. The method of claim 1, further comprising detecting a plurality of different targets comprising at least the target nucleic acid present in a sample using one or more capacitive sensors.

22. 22. The method of claim 21 , wherein the plurality of different targets comprises a plurality of different target nucleic acids.

23. 23. The method of claim 22, wherein the plurality of different target nucleic acids are from different pathogens or different variants of a pathogen.

24. 24. The method of claim 23, wherein the pathogen is SARS-CoV-2.

25. 22. The method of claim 21 , wherein the plurality of different targets comprises proteins.

26. 26. The method of claim 25, wherein the protein is an antibody.

27. 26. The method of claim 25, wherein the protein is an antigen.

28. 27. The method of claim 26, wherein the target nucleic acid is derived from a dengue virus and the antibody is a dengue-specific antibody.

29. detecting said plurality of distinct targets, binding each of said plurality of targets to a different bead; detecting each of the distinct beads using the capacitive sensor; 22. The method of claim 21 , comprising:

30. 30. The method of claim 29, further comprising distinguishing the different beads based on different capacitive detection signals from the capacitive sensors.

31. 22. The method of claim 21, wherein detecting the plurality of different targets comprises binding each of the plurality of different targets to a different probe on a substrate.

32. The method of claim 31 , wherein the substrate is operably associated with one or more capacitive sensors.

33. 32. The method of claim 31, further comprising binding one or more reporter (R) beads to each of the different substrate-bound targets.

34. selectively releasing the R beads from the substrate-bound targets based on the different substrate-bound targets to which the R beads are bound; flowing the released R beads in a fluid past the capacitive sensor; detecting the released R beads flowing past the sensor; 34. The method of claim 33, further comprising:

35. 30. The method of claim 29, wherein each of the different beads has a different size.

36. 36. The method of claim 35, further comprising microfluidically directing each of the different beads to a different capacitive sensor based on its size.

37. 37. The method of claim 36, wherein the guiding step includes one or more of an inertial method, a dielectrophoretic method, or a magnetophoretic method.

38. Detecting the target nucleic acid Binding the target nucleic acid to a probe on a substrate; binding a plurality of reporter (R) beads to the substrate-bound target nucleic acid; The method of claim 11 further comprising:

39. 40. The method of claim 38, wherein the plurality of R beads binds to different sequences in the substrate-bound target nucleic acid.

40. The method of claim 38 , wherein the substrate is operably associated with the capacitive sensor.

41. Releasing the plurality of R beads from the substrate-bound target nucleic acid; flowing the released R beads in a fluid past the capacitive sensor; detecting the released R beads flowing past the sensor; 40. The method of claim 39, further comprising:

42. The method of claim 9 , wherein the two electrodes form a trench through which the R bead passes.

43. Detecting the target nucleic acid binding transport (T) beads to the target nucleic acid in the sample to form a T-bead complex; binding the T-bead complex to a reporter (R) bead to form an R-bead complex; eluting the R beads and the T beads from the R bead complex; Detecting the R beads using a capacitance sensor; returning the T-beads to the sample to bind additional target nucleic acids; The method of claim 1 , comprising:

44. 1. A system for nucleic acid detection, comprising: A signal processing circuit; a sensor region including two capacitive electrodes on a substrate in communication with the signal processing circuitry; the two electrodes being spaced apart to form a gap; A system in which only a single bead moves across the gap at any one time, and the electrodes detect the bead and / or any molecules bound to the bead.

45. 45. The system of claim 44, wherein the system is reusable after being used to detect beads.

46. 45. The system of claim 44, wherein the capacitance sensor is approximately 2 mm by 2 mm.

47. 45. The system of claim 44, wherein the beads or particles are contained in water, a buffer, or an oil or a fluorocarbon, or an organic solvent and flow through the sensor region.

48. 48. The system of claim 47, wherein the beads or particles being detected are surrogates for microbiological analytes from an upstream assay.

49. 48. The system of claim 47, wherein the beads or particles can be magnetic or non-magnetic.

50. 45. The system of claim 44, wherein the gap is twice the bead diameter.

51. 45. The system of claim 44, wherein the substrate is 3D printed.

52. 52. The system of claim 51, wherein the electrodes are formed by sputtering or ink deposition and patterning.

53. 53. The system of claim 52, wherein the spacing between the electrodes is between about 0.8 μm and 20 μm.

54. 45. The system of claim 44, wherein the substrate is a CMOS semiconductor chip.

55. 55. The system of claim 54, wherein the electrodes are etched in a metal layer, and the spacing between the electrodes is between 40 nm and 5 μm.

56. 45. The system of claim 44, further comprising one or more of ultrasonic agitation, dielectrophoretic bead steering, or adjustment of bead zeta potential to prevent bead aggregation and / or oil syringing.

Images