US20160003798A1

US20160003798A1 - Apparatuses and methods for determining analyte charge

Info

Publication number: US20160003798A1
Application number: US14/745,253
Authority: US
Inventors: Kee-Hyun Paik; Henrik H.J. Persson; Billy Tsz Cheong Lau; Hanlee P. Ji; Robert W. Dutton; Yang Liu; Ronald W. Davis
Original assignee: Leland Stanford Junior University
Current assignee: Leland Stanford Junior University
Priority date: 2014-06-19
Filing date: 2015-06-19
Publication date: 2016-01-07
Also published as: EP3158330A4; CA2951945A1; CN107003295B; CN107003295A; WO2015196148A1; EP3158330A1; US20210048426A1; US20170097332A1; EP3158330B1; CA2951945C

Abstract

The present disclosure provides a sensor including a pore and an applied electric field that is capable of detecting analytes such as nucleic acids. In accordance with various embodiments, the sensor comprises a fluidic chamber having electrically opposing portions with a membrane between, the membrane providing a pore suitable for the passage of an electrolyte between the electrically opposing portions of the fluidic chamber, and having at least one charged analyte tethered in proximity to the pore, a first circuit configured to apply an electric field capable of passing the electrolyte through the pore and pulling the at least one charged analyte into the pore, and a second circuit configured to measure a signal indicative of the charge of the at least one charged analyte. Also provided are methods for using the sensor, for example, to sequence a nucleic acid molecule.

Description

CROSS-REFERENCE

This application claims the benefit of U.S. Provisional Patent Application Ser. No. 62/014,595, filed on Jun. 19, 2014, with three Appendices; this provisional patent document and its appendixes are fully incorporated herein by reference.

FEDERALLY-SPONSORED RESEARCH AND DEVELOPMENT

This invention was made with Government support under contract HG000205 awarded by the National Institutes of Health. The U.S. Government has certain rights in the invention.

OVERVIEW

There is a great deal of interest in the field of biotechnology on nucleic acid sensors that can replace the currently popular optics-based biosensors. While there has been numerous theoretical advances in nucleic acids research, the cost of performing the methods developed (whether for diagnosis of a patient's ailment or investigation of a pathogenic trait) frequently hamper their adoption into clinical settings. For example, while the cost of human genome sequencing has been dramatically reduced from $3 billion to $20 thousand, it is still far too expensive to be used in a routine clinical environment. The optics-based sensing (that tends to be time consuming to operate, needs modified fluorescing reagents, requires bulky optical sources and needs costly imaging equipment) is seen as a major bottleneck in lowering the cost of genomics. While the integrated electrical sensors seem to provide many advantages over the optical ones, recent approaches have either severe limitation in manufacturability (requiring definition of exceedingly small features O(1 nm)) or have shown poor robustness.
The above discussion/summary is not intended to describe each embodiment or every implementation of the present disclosure. The figures and detailed description that follow also exemplify various embodiments.

SUMMARY

Aspects of the present disclosure are believed to be application to a variety of different devices, systems, arrangements involving sensor devices that detect analytes. In certain embodiments, the sensor devices can be used to detect analytes such as nucleic acids (e.g., that are more sensitive, more robust, and/or more easily manufactured). In various aspects, the present disclosure provides such sensors and methods for using the sensors, for example, to sequence a nucleic acid molecule.
Various aspects of the present disclosure are directed toward a sensor for detecting a charged analyte and methods of using the sensor. The sensor includes a fluidic chamber having electrically opposing portions (e.g., a top portion and a bottom portion) with a membrane between. The membrane provides a pore suitable for the passage of an electrolyte between the electrically opposing portions (e.g., from the top portion to the bottom portion) of the fluidic chamber and having at least one charged analyte tethered in proximity to the pore; a first circuit configured to apply an electric field capable of passing the electrolyte through the pore and pulling the at least one charged analyte into the pore. The sensor further includes a second circuit configured to measure a signal indicative of the charge of the at least one charged analyte upon the at least one charged analyte being pulled into the pore.
The second circuit optionally includes a sensing electrode for measuring the signal, wherein the sensing electrode is located at a distance away from the at least one charged analyte. The distance may be at least 2-times a Debye length associated with the at least one charged analyte. In various embodiments, respective sizes or diameters (and/or shapes) of the pore can vary as needed to pass certain types (having corresponding sizes) of charged analytes. As one example, certain charged analytes may be appropriate for pores sized between 25 nm and 2000 nm in diameter as used with a membrane having a thickness between 50 nm and 3 μm. For other analytes such as smaller-sized analytes, the pore has a diameter of at least about 10 nanometers (nm). The at least one charged analyte may have an electrical double layer (EDL) surrounding it and the electric field may be capable of de-screening the EDL. Further, the membrane and walls of the pore may have an EDL surrounding them and the electric field may be capable of de-screening the EDL. The electric field may be capable of generating a non-equilibrium transport condition. The membrane may be electrically insulating. Further, the membrane may be comprised of silicon dioxide (SiO2) or silicon nitride (Si3N4) (e.g., which are and/or can be electrically insulating material). The first circuit may include a first electrode in a first electrically opposing portion (e.g., the top portion) of the fluidic chamber and a second electrode in a second electrically opposing portions (e.g, the bottom portion) of the fluidic chamber. The second circuit may include an electrode embedded in the membrane in proximity to the pore. The second circuit may include an amplifier capable of amplifying the signal. The signal may be linearly proportional to the charge of the at least one charged analyte. The at least one charged analyte may be a nucleic acid molecule. Further, the at least one charged analyte may have a net charge of at least about 40 e−. However, the at least one charged analyte may have a net charge lower or higher than about 40 e−. The electrolyte may have an ionic strength of about 100 μM to about 1M. Further, the at least one charged analyte can be tethered in proximity to the pore by a molecular structure and/or by being immobilized. Further, the sensor may have a plurality of pores into which the plurality of charged analytes are pulled. Further, the plurality of charged analytes may be clonal. In another aspect, a device is disclosed in which the device has a plurality of the sensors detailed herein.
Other related embodiments are directed to a kit for detecting a charged analyte. The kit includes at least one charged analyte, and a sensor. The sensor includes a fluidic chamber having electrically opposing portions (e.g., a top portion and a bottom portion) with a membrane between, the membrane providing (or defining) a pore suitable for the passage of an electrolyte between the electrically opposing portions (e.g., from the top portion to the bottom portion) of the fluidic chamber and having the at least one charged analyte tethered in proximity to the pore (e.g., sufficiently the pore for the electrical field interactions described herein. The sensor also includes a first circuit configured to apply an electric field capable of passing the electrolyte through the pore and pulling the at least one charged analyte into the pore. The sensor also includes a second circuit configured to measure a signal indicative of the charge of the at least one charged analyte upon the at least one charged analyte being pulled into the pore. The first circuit may include a first electrode in a first electrically opposing portion (e.g., the top portion) of the fluidic chamber and a second electrode in the second electrically opposing portion (e.g., the bottom portion) of the fluidic chamber. The second circuit may include an electrode embedded in the membrane in proximity to the pore. The second circuit may include an amplifier capable of amplifying the signal. The signal may be linearly proportional to the charge of the at least one charged analyte. The at least one charged analyte may be a nucleic acid molecule. Further, the at least one charged analyte may have a net charge of at least about 40 e⁻. However, the at least one charged analyte may have a net charge lower or higher than about 40 e⁻. The electrolyte may have an ionic strength of about 100 μM to about 1M. Further, the sensor may have a plurality of pores into which the plurality of charged analytes are pulled. Further, the plurality of charged analytes may be clonal.
In accordance with other related embodiments, aspects of the present disclosure are directed to a method for detecting a charged analyte. The method includes providing a fluidic chamber having electrically opposing portions (e.g., a top portion and a bottom portion) between a membrane, one of the electrical opposing portions (e.g., the top portion) having an electrolyte, the membrane providing a pore suitable for the passage of an electrolyte between the electrically opposing portions (e.g., from the top portion to the bottom portion) of the fluidic chamber, and having at least one charged tethered in proximity to the pore. The method further includes applying an electric field to pass the electrolyte through the pore and pull the at least one charged analyte into the pore; and measuring a signal indicative of the charge of the at least one charged analyte upon the at least one charged analyte being pulled into the pore. The method optionally includes the at least one charged analyte being pulled to a position in proximity to a periphery of the pore. In this method, the first circuit may apply the electric field. The second circuit may measure the signal.
Additional aspects and advantages of the present disclosure will become readily apparent to those skilled in this art from the following detailed description, wherein only illustrative embodiments of the present disclosure are shown and described. As will be realized, the present disclosure is capable of other and different embodiments, and its several details are capable of modifications in various obvious respects, all without departing from the disclosure. Accordingly, the drawings and description are to be regarded as illustrative in nature, and not as restrictive.

BRIEF DESCRIPTION OF THE FIGURES

Various example embodiments may be more completely understood in consideration of the following detailed description in connection with the accompanying drawings (including those in the in the Appendices that form part of this patent document) in which:

FIG. 1 a shows an example schematic of an integrated biosensor and FIG. 1 b shows an example circuit diagram of the proposed integrated biosensor shown in FIG. 1 a, consistent with various aspects of the present disclosure;

FIGS. 2 a-h show an example sensor being used for sequencing by a synthesis scheme, consistent with various aspects of the present disclosure;

FIGS. 3 a-b show example contour plots of simulated electrostatic potential change due to the presence of the charged biomolecule with 0 V and 7 V external electrical biases applied, consistent with various aspects of the present disclosure;

FIG. 4 a shows an example schematic plot of a charge sensor device structure, consistent with various aspects of the present disclosure;

FIG. 4 b shows a fraction of charge in the sensing electrode shown in FIG. 4 a induced by the biomolecule when no oxide is in between the electrode and solution, consistent with various aspects of the present disclosure;

FIG. 4 c shows an example plot of the effect of dielectric formation on the sense electrode on sensing efficiency, consistent with various aspects of the present disclosure;

FIG. 5 shows an example pore array for proof of concept demonstration of charge sensing, consistent with various aspects of the present disclosure;

FIGS. 6 a-b show an example process of covalently attaching nucleic acids to the charge sensor surface, consistent with various aspects of the present disclosure;

FIGS. 7 a-c show an example optical microscopy study of solid-phase PCR amplified DNA on the sensor surface, consistent with various aspects of the present disclosure; and

FIGS. 8 a-b show an example demonstration of a charge sensor, consistent with various aspects of the present disclosure.

While various embodiments discussed herein are amenable to modifications and alternative forms, aspects thereof have been shown by way of example in the drawings and will be described in detail. It should be understood, however, that the intention is not to limit the invention to the particular embodiments described. On the contrary, the intention is to cover all modifications, equivalents, and alternatives falling within the scope of the disclosure including aspects defined in the claims.

DESCRIPTION

Various aspects of the present disclosure are directed toward integrated, highly-manufacturable, solid-state nucleic acid charge sensors for sequencing and DNA microarray applications. For instance, aspects of the present disclosure are directed toward apparatuses, methods and systems that include a fluidic chamber having a top portion and a bottom portion that hold charged analytes. Further, the apparatuses, methods and systems can include a membrane separating the top portion and the bottom portion of the fluidic chamber. The membrane includes an opening to provide a pathway between the top portion and the bottom portion of the fluidic chamber. Additionally, the apparatuses, methods and systems can include a first circuit that applies an electric field to tether a cluster of the biological molecules. Further, the apparatuses, methods and systems can include a sensor and an integrated circuit that determine a charge of the biological molecules while the cluster of the charged analytes are tethered.
In certain embodiments, the charged analytes are one or more of DNA molecules and RNA molecules. Additionally, in certain embodiments, the charged analytes are one or more of inorganic toxins (e.g., cadmium, fluorides, mercury, lead, arsenic, toxic element salts), drugs, proteins, other toxins, fungal spores, bacteria, viruses, heavy metals, and other similar charged analytes. Other embodiments of the present disclosure are further characterized as having an exterior portion of the membrane that includes a plurality of adapters which provide solid-phase amplification of the charge sensed by the sensor and the integrated circuit. Additionally, in certain embodiments, an exterior portion of the membrane includes a plurality of adapters that provide solid-phase amplification to create a clonal DNA cluster. Further, a polymerase chain reaction (PCR) primer can be attached to the tail end of the DNA. Further, the pore can be between 25 nm and 2000 nm in diameter, and the membrane can be between 50 nm and 3 μm thick.
Certain embodiments of the present disclosure include a membrane and walls of the opening that form an electric double layer (EDL). In such embodiments, the first circuit generates a non-equilibrium transport condition for descreening of the EDL. In other embodiments, the first circuit pivots the anchored DNA molecules into the pore in response to the electric field. Additionally, the sensor and an integrated circuit can determine the charge of the charged analytes to sense base incorporations of the charged analytes. Further, the first circuit can also include a cathode and an anode in the fluidic chamber to apply the electric field. One of the anode and the cathode is in the top portion of the fluidic chamber, and the other of the anode and the cathode is in the bottom portion of the fluidic chamber. In other embodiments, the first circuit immobilizes the cluster of the charged analytes as separated away from the walls of the pore, and the sensor and an integrated circuit are configured and arranged to determine the charge of the charged analytes. Additionally, certain embodiments can include an array of biological sensing devices. Aspects of the present disclosure can replace any biosensing that is currently done optically, chemically or radiologically. Additionally, applications include but are not limited to DNA, RNA or protein sequencing, DNA microarray and immunoassay.
Since nucleic acids have a 1e⁻ charge in its phosphate backbone, the net charge on a nucleic acid molecule is directly proportional to the number of bases in it. Thus, the ability to monitor the amount of charge on a nucleic acid's molecule can enable monitoring of the number of bases in a molecule; the knowledge about the number of bases in a DNA or RNA molecule, in turn, enables the detection of synthesis events for sequencing or hybridization events for microarrays.
Two major challenges to charge sensing via electronic charge sensors in an aqueous environment include excessive confinement requirement due to electric double layer's (EDL) shielding of analyte charge and difficulty of capturing the analytes for sensing. Surprisingly, aspects of the present disclosure are directed to a non-equilibrium transport phenomenon along with a strategic immobilization of analytes to circumvent the challenges for charge-based biosensors. Novel physics enable utilization of devices for various applications. Because charge is an inherent characteristic of nucleic acids, various aspects of the present disclosure enable fast, label-free detection of nucleic acids for cost-effective analysis. Aspects of the present disclosure directed toward optics-based methods of sensing can dramatically reduce the entry barrier to perform nucleic acids research over radioisotope labeled nucleic acids sensing. The regulatory simplification from not using radiation sources can provide a plethora of commercial analysis tools (the next generation sequencing, DNA microarray, real-time PCR, etc.).
Various aspects of the present disclosure are directed toward an integrated charge sensor chip that can include a source follower (SF) amplifier and a sense electrode in close proximity resting on a thin SiN_xmembrane. Such an integrated charge sensor can be a passive non-integrated sensor.
The embodiments and specific applications discussed herein may be implemented in connection with one or more of the above-described aspects, embodiments and implementations, as well as with those shown in the figures, including the description and figures shown in the Appendices as filed as part of the underlying provisional application, which form part of this patent document and are fully incorporated herein by reference.
Turning now to the figures, FIG. 1 a shows an example schematic of an integrated biosensor and FIG. 1 b shows an example circuit diagram of the proposed integrated biosensor shown in FIG. 1 a, consistent with various aspects of the present disclosure. The biosensor consists of an electrode embedded in a 100˜300 nm pore, that is easily definable by photolithography. The signal from the electrode is read through a thin-film unity gain c SF amplifier integrated in close proximity to the sensor. The clonal nucleic acid analytes are immobilized on one surface of the sensor in such a way that by applying the appropriate bias on the anode, the negatively charged DNA or RNA can be drawn into the pore for sensing. The membrane (200˜400 nm thick) provides the electrical confinement to generate the non-equilibrium transport condition necessary for descreening of EDL.
FIG. 2 shows an example sensor being used for sequencing by a synthesis scheme, consistent with various aspects of the present disclosure: a) the cutaway schematic of the charge sensor (the device consists of a pore in a thin membrane whose perimeter is coated with adaptors for solid-phase amplification. A sensor is embedded in the pore with a nearby integrated circuit); b) prepared library DNA is attached to an adaptor; c) Solid-phase amplification is done to create a clonal DNA colony; d) PCR primer is attached to the tail end of the DNA; and e) by applying an external electric field, the anchored DNA molecules pivot into the pore. During this time, a reference charge read is performed. The difference between the final charge level and the result of this reference read will determine whether or not nucleotides were incorporated or not. The high electric field that will help descreen the EDL is also used to pull the immobilized DNA strands into the sensor; f) One of the four nucleotides are introduced and allowed to be incorporated; g) The excess nucleotides are washed away or degraded away (e.g. with apyrase); and h) High electric field is applied to perform a charge read to sense base incorporations.
Solid-phase PCR amplification is used to generate a immobilized clonal DNA cluster (FIGS. 2 a-c), end primer will be introduced and excess un-hybridized primers will be washed away (FIG. 2 d). The charge level of the DNA molecules prior to a base incorporation will be measured and stored. One of the four bases will be introduced along with polymerase enzymes (FIG. 2 f). If the base is of the appropriate type, it will be synthesized into the DNA molecule (FIG. 2 g). The excess nucleotides will be washed and degraded away. The charge level of the DNA molecules post-base introduction will be read (FIG. 2 h). This level will be compared against the previous charge state stored to perform a base-calling. The cycle returns to step in FIG. 2 e) for other base types. A conceptually similar method can be employed for DNA microarray application where hybridization of target DNA or RNA would be done to synthesized probes. In certain embodiments, the pore can be large (>100 nm) with an embedded electrode. The DNA to be sensed is attached near the pore.
An example sensing scheme is based on delivery of charged molecules to the charge sensor via applied electric field which also suppresses the electrical charge-shielding in the confined geometry of a pore through a thin (ca. 100 nm) membrane. Because the electrostatic potential drop across the device is dominated by the pore, high electric fields (ca. 10⁶˜10⁷V/m) can be easily generated inside it. The resulting ionic current through the pore can disrupt the electrostatic screening of the molecules in the sensing region, making it possible to detect their charge hundreds of nanometers away. This is a surprising effect since under equilibrium conditions the Debye-Huckel screening model predicts that charge sensing is only possible within a distance of a few Debye lengths away from the target biomolecules (Debye length, λ_D, is ˜1 nm at physiological conditions).
In the presence of ionic current flow in nano-confined geometries, the effective ionic screening length can dramatically increase. By applying electrical biasing across aqueous pores, electro-diffusion current flow is present, particularly along the radial direction due to the presence of the charged biomolecules; this current significantly suppresses the charge-screening effect. This finding serves as the operation principle of our proposed devices, which aim at sensing the charge of biomolecule at distances 10-100 times the Debye length, XD.
FIG. 3 shows example contour plots of simulated electrostatic potential change due to the presence of the charged biomolecule with 0 V and 7 V external electrical biases applied, consistent with various aspects of the present disclosure: (a) for zero external electrical bias, the Debye-Huckel screening behavior is observed; and (b) for an electrical bias of 7 V applied across the pore, significant long-range electrostatic interaction is observed. The membrane is modeled as a solid dielectric layer of 500 nm thickness. The ionic solution is 1 mM KCl. The pore radius is set to 300 nm, corresponding to ˜30 Debye lengths at this molar concentration.
To demonstrate the operating principle, a cylindrically symmetric model system was simulated where a fragment of 60 bp double-stranded (DS) DNA is located at the center (the worst case scenario) of an aqueous nanopore, as schematically shown in FIG. 1. The Poisson-Nernst-Planck (PNP) equations along with the Stokes equations have been solved to model the ionic and fluidic transport across the pores using the general partial differential equation solver Prophet, to solve the nonlinear, coupled model equations. Example simulation results are presented in FIG. 3, in which the change of electrostatic potential due to the presence of the charged biomolecule is plotted. The validity of the simulator has been shown by its ability to accurately simulate DNA translocation behavior through gated pores of similar dimensions for actuation application.
Further simulations were performed to study more specifically a device structure, consistent with various asepects of the present disclosure. FIG. 4 a shows an example schematic plot of a charge sensor device structure, consistent with various aspects of the present disclosure. FIG. 4 b shows a fraction of charge in the sensing electrode shown in FIG. 4 a, induced by the biomolecule when no oxide is in between the electrode and solution, consistent with various aspects of the present disclosure. FIG. 4 c shows an example plot of the effect of dielectric formation on the sense electrode on sensing efficiency, consistent with various aspects of the present disclosure. The schematic plot shows the effect of dielectric formation on the sense electrode on sensing efficiency. This models the sensing metal electrode sandwiched between two insulating layers. Assuming the biomolecule charge is −Q, the induced charge in the metal electrode is Q′=βQ, which is essentially the amount of charge sensed by the amplifier circuitry (FIG. 2). The sensing efficiency or the fraction of induced charge, β, has been calculated for different pore radii and biasing conditions. The conclusion of this modeling was that, with modest DC electrical bias across the pore, the charge that could be sensed is around 20˜40% of the biomolecule charge for a 500 nm pore when the biomolecule is at the center of the pore. The charge can be delivered much closer to the sense electrode due to the way the DNA are immobilized and the external electric field profile affects them.
Various aspects of the present disclosure are directed toward a non-integrated passive charge sensor. The passive sensor can include a membrane with a 2D array of micropores with an embedded platinum sense electrode as shown in FIG. 5. An array of pores were used to enhance the net charge delivered to the sensor in absence of a dedicated integrated amplifier, such need for the array will disappear with circuit integration. FIG. 5 shows an example pore array for proof of concept demonstration of charge sensing, consistent with various aspects of the present disclosure: a) a schematic of the device (a 75 nm metal [70 nm Pt, 5 nm Ti]) electrode is embedded in a thin membrane comprising of an 80 nm thick Si₃N₄and a 70 nm thick SiO₂. The membrane has an array of pores defined by photolithography; b) top (SiO₂side) view scanning electron microscope image of the pore array (the Pt electrode visible is 400 μm×400 μm in size); c) bottom (Si₃N₄side) view scanning electron microscope image of the pore array (the Si substrate is visible around the border); d) top view scanning electron microscope image of an individual pore; and e) bottom view electron microscopy of an individual pore (the scale bars in d) and e) are 1 μm in size). Measured pore diameters are ca. 1.15 μm.
Densely populated immobilized DNA can be provided to a sensor and solid-phase PCR amplification or bridge amplification protocols have been developed. A gas phase silanization of the chip surface was performed with molecular vapor deposition of (3-aminopropyl)-trimethoxysilane (APTMS). Using the crosslinker N-(p-Maleimidophenyl)isocyanate (PMPI) a thiol-modified oligonucleotide that acts as the PCR primer is attached to the chip surface. There are a variety of other crosslinkers that can be used whose properties and effect on the immobilization should be further explored. The schematic of the attachment chemistry is shown in FIG. 6.
FIG. 6 shows an example process of covalently attaching nucleic acids to the charge sensor surface, consistent with various aspects of the present disclosure: a) the SiO₂surface is silanized in gas phase (the crosslinker PMPI is used to connect the amine group on the silane and the sulfhydryl group of the 5′ thiol-modified primer); and b) the product of the surface chemistry. Sensing chips were plasma cleaned, rehydrated and functionalized with Aminopropyltrimethoxysilane using a chemical vapor deposition system. The amino-functionalized surfaces were subsequently transformed into a thiol-reactive moiety by exposure to a 2.3 mM solution of N-(p-maleimidophenyl) isocyanate, PMPI in anhydrous toluene at 40° C. for 2 h under an argon atmosphere. The surfaces were subsequently washed with anhydrous toluene and dried in a stream of argon followed by DNA immobilization using thiolated oligonucleotides. Prior to immobilization the thiolated oligos were reduced using TCEP as a reducing agent and desalted using a spin column (MWCO=3000). Thiolated oligos can be spotted directly onto sensing chips for 6 h at 10 uM concentration in a 1M NaCl buffer solution under a controlled atmosphere, followed by extensive washing. The various surface modification steps were followed by x-ray photoelectron spectroscopy and the presence of the expected elements and peak shifts confirmed the transformation of the sensing surface. The bridge PCR amplification itself is done by thermal cycling the sensor chip in a standard PCR tube along with the appropriate reagents and a 900 bp template previously prepared. The result of the attachment chemistry is a chain of covalent bonds securely immobilizing 900 bp DNA molecules to the chip surface. The length of 900 bp template was selected since its length is a close match the fabricated pore length. The solid-phase amplified DNA molecules are linearized by a restriction enzyme. The dense presence of immobilized nucleic acid from solid-phase amplification is verified by fluorescence microscopy with appropriately excited SYBR Gold nucleic acids dye. FIG. 7 shows the microscopy images the charge sensor chip with and without the immobilized DNA molecules.
In order to verify that the result of fluorescence response is from successful bridge amplification and not from nonspecific binding, several control experiments were carried out. In each experiment, a component in the surface chemistry (aminosilane, crosslinker and thiolated oligo) was omitted prior to thermal cycling that nominally would results in PCR amplification. FIG. 7 shows the result. The absence of any of the component resulted in a low level of fluorescence signifying that the solid-phase PCR amplification is only successful when all of three surface chemistry components are in place. Slightly elevated fluorescence brightness in the case where aminosilane was deposited and thiolated oligonucleotide primers were incubated without the presence of the crosslinker, is speculated to be low levels of nonspecific adhesion of negatively charged oligos to the positively charged amine group of the silanes.
FIG. 7 shows an example optical microscopy study of solid-phase PCR amplified DNA on the sensor surface, consistent with various aspects of the present disclosure: a) bright field microscopy image of a sensor with solid-phase PCR amplified DNA of ca. 900 bases (the light gray area is the SiN_x/SiO₂membrane. The dark porous area is the Pt sense electrode with the pore array. The 1 μm pores are defined photolithography; b) fluorescence microscopy image of the chip at the location shown in a). SYBR Gold stained DNA fluoresces signifying an abundance of DNA molecules, a successful solid-phase PCR amplification; and c) a series of control experiments, where a component in the surface chemistry has been omitted to enhance fluorescent brightness of with various missing surface chemistry.
With nucleic acids immobilized, a passive sensor chip, consistent with various aspects of the present disclosure, can be operated in the following fashion: i) a positive potential is applied to the cathode, ii) application of the potential creates an electric field near the pore in such way that the immobilized DNA molecules are drawn into the pore and iii) the DNA molecules' presence in the pore under external electric field leaves an electrical signal onto the platinum sense electrode whose potential is recorded for analysis. With DNA covalently immobilized on the top surface, the sensor's ability to distinguish charge is tested by observing the signal difference between a negative control experiment where there is no surface chemistry done to the sensor chips, chips with single-stranded (SS) DNA attached with 900e− per molecule and chips with DS DNA attached with 1800e− per molecule of charge.
FIG. 8 shows an example demonstration of a charge sensor, consistent with various aspects of the present disclosure: a) the schematic of the charge sensor setup (the charge sensor is attached to a unity gain voltage amplifier reading the response of the sensor electrode during the experiments. Linearized DNA is immobilized on the SiO₂surface). When positive potential is applied to the cathode, DNA molecules are pulled into the pore, placed near the sense electrode altering its potential. The experiments are done in low concentration salt (100 μM KCl) to enhance the electrical detection; and b) the sensor potential monitored during experiments (each curve represents an experiment with a sensor chip. Green lines represent the negative control experiments where there were no DNA attached. Black lines represent those chips that have had SS DNA attached to them. Red lines represent those chips that have had DS DNA attached. While some chip-to-chip variation is observed, overall the sensor is able to differentiate the net charge difference between the three cases. With a 100 μV difference observed between ca. 900 bases, we see a 110 μV/base.
Example applications for nucleic acids sensors include sequencing and microarrays.
Many of the sequencing technologies are based on the sequencing by synthesis (SBS). The majority of the methods are based on polony sequencing. The SBS reaction appears as follows:
$\begin{matrix} D N A (n) + dNTP \overset{Polymerase}{} D N A (n + 1) + H^{+} + {PP}_{i} & (Eq . 1) \end{matrix}$
Where DNA(n) is a DNA molecule with n bases, dNTP is the deoxynucleotide triphosphate and PP_iis the pyrophosphate. Thus there are three items can be detected by varieties of sensors for SBS. The addition of the base itself, the proton released during synthesis and the pyrophosphate released during synthesis. A polony consists of 1000+ identical copies of a DNA molecule to be sequenced. The multiplexed signal given off by the identical individual DNA molecules being synthesized in a polony in parallel enhances the integrity of calling (reading) a base.
Based on the solid-phase PCR amplification (bridge amplification), a sequencing platform that uses optics to detect the addition of fluorescently modified base (the increase of DNA(n) to DNA(n+1)). While the modification is necessary for DNA that does not naturally fluoresce, such modification can disrupt the polymerase enzyme's natural functioning to result in higher erroneous incorporation, which statistically occur in parts of the polony. Once such erroneous incorporations occur, the molecule no longer produces the right signal and contributes to read error of that entire polony. When sufficient number of DNA molecules in a polony has been corrupted (i.e. is off phase), the polony loses the ability to accurately call a base. This limits the read length to be 100˜300 bases. Further, the laser optical sources are bulky; and the cameras acquiring the images of sequencing results are slow and produce large data files. The recent developments in optical detection have been limited to incremental improvements in performance signifying its mature developmental status.
Pyrosequencing detects the release of the pyrophosphate, a byproduct of the synthesis reaction (see Eq. 1). It is an optics-based technology where a series of reactions are done in microfluidically confined reaction chambers to fluorescently observe the presence of the pyrophosphate. Challenges can arise from difficulty of scaling the signal transduction from the reaction wells to the sensor, for which bundles of fiber-optic cables were used. However the technology does have advantages, the pyrosequencing synthesis reaction does not require modified reagents. The result is a resilience to phasing error with read length being 1000+, an order of magnitude larger than the more modern techniques that surpassed pyrosequencing in popularity.
Solid state pH sensors have also been successfully used to detect the H⁺ ion released from polonies after base incorporation. Because the sensor is based on solid-state devices sequencing technology, it is dramatically faster than that of the optical sensors. However, since it is the pH that is sensed, each reagent's pH must be carefully calibrated and the reaction chamber cannot be strongly buffered. This results in a delicate initialization process, that is time consuming and prone to failure. The local pH change in and around a sensor also diffuse away transiently and the synthesis result cannot be accessed multiple times, needing a fixed window upon which data must be gathered. Further, since pH-based sensors detect a byproduct of a specific molecular biological event in DNA synthesis, SBS can be accomplished.
pH sensors operate on the logarithmic nature of pH. Solid state sensors used for pH detection can be based on the ion Sensitive Field-Effect Transistor (ISFET) technology. The ISFETs have a linear output response to change in pH, ca. 50 mV/pH. The pH depends logarithmically on the synthesis of a nucleotide. Miniaturization of device dimensions is a frequently used method of cost reduction in semiconductor microfabrication. To ensure that homopolymers of various lengths are distinguishable, the pH sequencing method can require a high number of clonal DNA near a sensor.
Long ranged interaction (>100 nm) can be exploited for both sensing and actuating charged biomolecules including nucleic acids. Various sensors, consistent with the present disclosure, detect the charge in the phosphate backbone. In these embodiments, electrical solid-state sensors enable a fast read operation in both sequencing and DNA microarray. For sequencing application, unlike sensing pH, sensing the charge in the phosphate backbone will result in a linear response to the number of bases incorporated, thus not suffering from poor homopolymer performance. It is also permanently fixed and can be accessed multiple times for error reduction.
Reading a charge in a phosphate backbone that is inherent in the DNA molecules themselves, can greatly simplify the chemistry. Accordingly, charge sensors, consistent with various aspects of the present disclosure, may not require modified reagents (nucleotide, polymerase) or additional reagents for detection (e.g. ATP sulfurylase, luciferase). Thus, the nanofluidic charge sensors offer simple replacement of current sensing methods while maintaining the various advantages of the sensing methods. Additionally, the sequencing platform based on the sensor described herein can have the long read length of pyrosequencing, speed of solid-state sequencing and the robustness traditionally associated with optical sensing. Thus, by using a manufacturable solid state sensor that is capable of directly detecting changes in the inherent charge of a DNA molecule, we can circumvent many issues plaguing the current and emerging next-generation sequencing platforms. The solid state integrated charge sensor can function independent of pH, read the base incorporation events quickly, have efficient data storage and also have a less expensive scaling cost with better homopolymer resolution.
The embodiments and specific applications discussed herein (and in the Appendices of the underlying provisional application) may be implemented in connection with one or more of the above-described aspects, embodiments and implementations, as well as with those shown in the appended figures. This description and the various embodiments form part of this patent document and are fully incorporated herein by reference.
Further, the following Appendices are hereby fully incorporated by reference for their general and specific teachings: Appendix A, Appendix B, and Appendix C. Consistent with embodiments of the present disclosure, the Appendices, entitled “Appendix A”, “Appendix B”, and “Appendix C”, describe and show examples of sensor devices and methods of detecting charged analytes using a sensor device, in accordance with the present disclosure.
As illustrated, various modules and/or other circuit-based building blocks may be implemented to carry out one or more of the operations and activities described herein or in the Appendices, and/or shown in the block-diagram-type figures. In such contexts, these modules and/or building blocks represent circuits that carry out one or more of these or related operations/activities. For example, in certain of the embodiments discussed above and in the Appendices, one or more modules and/blocks are discrete logic circuits or programmable logic circuits configured and arranged for implementing these operations/activities, as in the circuit modules/blocks shown above and in the Appendices. In certain embodiments, the programmable circuit is one or more computer circuits programmed to execute a set (or sets) of instructions (and/or configuration data). The instructions (and/or configuration data) can be in the form of firmware or software stored in and accessible from a memory (circuit). As an example, first and second modules/blocks include a combination of a CPU hardware-based circuit and a set of instructions in the form of firmware, where the first module/block includes a first CPU hardware circuit with one set of instructions and the second module/block includes a second CPU hardware circuit with another set of instructions.
Various embodiments described above, and discussed in the Appendices may be implemented together and/or in other manners. One or more of the items depicted in the present disclosure and in the Appendices of the provisional application can also be implemented separately or in a more integrated manner, or removed and/or rendered as inoperable in certain cases, as is useful in accordance with particular applications. In view of the description herein, those skilled in the art will recognize that many changes may be made thereto without departing from the spirit and scope of the present disclosure.
Accordingly, in view of the description herein, those skilled in the art will recognize that many changes may be made thereto without departing from the spirit and scope of the present disclosure.

Claims

What is claimed is:

1. A sensor for detecting a charged analyte, the sensor comprising:

a fluidic chamber having electrically opposing portions with a membrane between, the membrane providing a pore suitable for the passage of an electrolyte between the electrically opposing portions of the fluidic chamber, and having at least one charged analyte tethered in proximity to the pore;

a first circuit configured to apply an electric field capable of passing the electrolyte through the pore and pulling the at least one charged analyte into the pore; and

a second circuit configured to measure a signal indicative of the charge of the at least one charged analyte upon at least one charged analyte being pulled into the pore.

2. The sensor of claim 1, wherein the at least one charged analyte is tethered, in proximity to the pore, concurrently with the pulling of the at least one charged analyte into the pore.

3. The sensor of claim 1, wherein the electrically opposing portions include a top portion and a bottom portion separated by the membrane, and the at least one charged analyte is tethered in proximity to the pore caused by interaction between the at least one charged analyte and the electrically opposing portions of the fluidic chamber.

4. The sensor of claim 1, wherein the second circuit comprises a sensing electrode for measuring the signal, wherein the sensing electrode is located at a distance away from the at least one charged analyte.

5. The sensor of claim 2, wherein the distance is at least 2-times a Debye length associated with the at least one charged analyte.

6. The sensor of claim 2, wherein the Debye length is calculated using the Debye-Hückel equation: Λ_D˜√{square root over (∈kT/C₀)}, wherein _D=Debye length, =electric constant, k=Boltzman constant, T=temperature, and C₀=ionic concentration.

7. The sensor of claim 1, wherein the pore has a diameter that is sized for the charged analyte, wherein both the diameter of the pore and the diameter of the charged analyte are on the order of more than several nanometers.

8. The sensor of claim 1, wherein the electric field has a strength in terms of million of Volts per meter (V/m) sufficient to suppress electrical-charge shielding within a geometry defined by a portion of the membrane defining the pore.

9. The sensor of claim 1, wherein the at least one charged analyte has an electrical double layer (EDL) surrounding it and the electric field is capable of de-screening the EDL.

10. The sensor of claim 1, wherein the membrane and walls of the pore have an EDL surrounding them and the electric field is capable of de-screening the EDL.

11. The sensor of claim 1, wherein the electric field is capable of generating a non-equilibrium transport condition.

12. The sensor of claim 1, wherein the membrane is electrically insulating.

13. The sensor of claim 1, wherein the first circuit comprises a first electrode in one of the electrically opposing portions of the fluidic chamber and a second electrode in another of the electrically opposing portions of the fluidic chamber.

14. The sensor of claim 1, wherein the second circuit comprises an electrode embedded in the membrane in proximity to the pore.

15. The sensor of claim 1, wherein the signal is linearly proportional to the charge of the at least one charged analyte.

16. The sensor of claim 1, wherein the at least one charged analyte is a nucleic acid molecule.

17. The sensor of claim 1, wherein the at least one charged analyte is tethered in proximity to the pore by a molecular structure.

18. The sensor of claim 1, wherein the diameter of the pore corresponds to the at least one charged analyte to pass through the pore.

19. A method for detecting a charged analyte, the method comprising:

providing a fluidic chamber having electrically opposing portions with a membrane between, one of the electrically opposing portions having an electrolyte, the membrane providing a pore suitable for the passage of the electrolyte between the electrically opposing portions of the fluidic chamber, and having at least one charged analyte tethered in proximity to the pore;

applying an electric field to pass the electrolyte through the pore and pull the at least one charged analyte into the pore; and

measuring a signal indicative of the charge of the at least one charged analyte upon the at least one charged analyte being pulled into the pore.

20. The method of claim 19, wherein in response to the applied electric field, the at least one charged analyte is pulled to a position in proximity to a periphery of the pore.

21. A kit for detecting a charged analyte, the kit comprising:

at least one charged analyte; and

a sensor including:

a fluidic chamber having electrically opposing portions with a membrane between, the membrane providing a pore suitable for the passage of an electrolyte between the electrically opposing portions of the fluidic chamber, and having the at least one charged analyte tethered in proximity to the pore;

22. The kit of claim 21, wherein the second circuit comprises an amplifier capable of amplifying the signal.

23. The kit of claim 22, wherein the amplifier is within about 5000 μm from the pore.

24. The kit of claim 21, wherein the at least one charged analyte is tethered in proximity to the pore by, and sufficiently near the pore for, immobilization caused by interaction between the at least one charged analyte and the electrically opposing portions of the fluidic chamber.