Patent Application
20140099726
A nanopore is a nano-scale conduit that forms naturally as a protein channel in a lipid membrane (a biological pore), or is engineered by drilling or etching the opening in a solid-state substrate (a solid-state pore). When such a nanopore is incorporated into a nanodevice comprising two chambers which are separated by the nanopore, a sensitive patch-clamp amplifier can be used to apply a trans-membrane voltage and measure ionic current through the pore.
Nanopores offer great promise for inexpensive whole genome DNA sequencing. In this respect, individual DNA molecules can be captured and driven through the pore by electrophoresis, with each capture event detected as a temporary shift in the ionic current. The sequence of a DNA molecule can then be inferred from patterns within the shifted ionic current record, or from some other auxiliary sensor in or near the nanopore, as DNA passes through the pore channel.
In principle, a nanopore sequencer can eliminate the needs for sample amplification, the use of enzymes and reagents used for catalytic function during the sequencing operation, and optics for detection of sequencing progress, some or all of which are required by the conventional sequencing-by-synthesis methods.
Electric nanopore sensors can be used to detect DNA in concentrations/volumes that are no greater than what is available from a blood or saliva sample. Additionally, nanopores promise to dramatically increase the read-length of sequenced DNA, from 450 bases to greater than 10,000 bases.
There are two principle obstacles to nanopore sequencing: (1) the lack of sensitivity sufficient to accurately determine the identity of each nucleotide in a nucleic acid for de novo sequencing (the lack of single-nucleotide sensitivity), and (2) the ability to regulate the delivery rate of each nucleotide unit through the nanopore during sensing. While many research groups are developing and improving nanopores to address obstacle 1, there is no method for addressing obstacle 2 that does not involve the use of enzymes or optics, both of which work only in specialized nanopore techniques and which incur higher complexity and cost compared to purely electrical methods.
In one embodiment, the present disclosure provides a device which comprises (a) a plurality of chambers, each chamber in communication with an adjacent chamber through at least one pore wherein the device contains at least two pores defined as a first pore and a second pore, (b) means to move at least a portion of the polymer out of the first pore and into the second pore and (c) at least one sensor capable of identifying individual components of the polymer during movement of the polymer through the first and second pores, provided that when only a single sensor is employed, the single sensor does not include two electrodes placed at both ends of a pore to measure an ionic current across the pore.
In one embodiment, the first and the second pores are about 1 nm to about 100 nm in diameter, and are about 10 nm to about 1000 nm apart from each other.
In one embodiment, the sensor is configured to identify the polymer by measuring a current, a voltage, pH, an optical feature or residence time associated with the polymer or one or more components of the polymer.
In one embodiment, the sensor is configured to form a tunnel gap allowing the polymer to pass through the tunnel gap when the polymer is loaded in both the first and the second pores.
In one embodiment, the sensor comprises a membrane having a hole forming the tunnel gap. In one embodiment, the hole is substantially round.
In one embodiment, the sensor comprises two ends forming a tunnel gap therebetween.
In one embodiment, the sensor comprises an end and a substantially flat surface forming a tunnel gap therebetween. In one embodiment, the sensor comprises two electrodes placed apart to form a tunnel gap therebetween.
In one embodiment, the sensor is placed within the first pore. In another embodiment, the sensor is placed in a chamber between the first and the second pores. In one embodiment, a device wherein the sensor is aligned with the first and second pores.
In one embodiment, the sensor comprises gold, platinum, graphene, or carbon.
In one embodiment, the tunnel gap is from about 1 nm to about 20 nm wide.
In one embodiment, the sensor comprises surface modification by a reagent. In one embodiment, the reagent is capable of forming a non-covalent bond with a nucleotide. In one embodiment, a device wherein the bond is a hydrogen bond.
In one embodiment, the reagent is selected from the group consisting of 4-mercaptobenzamide and 1-H-Imidazole-2-carboxamide.
In one embodiment, the first and second pores are substantially coaxial.
In one embodiment, the devices comprises a material selected from the group consisting of silicon, silicon nitride, silicon dioxide, graphene, carbon nanotubes, TiO2, HfO2, Al2O3, metallic layers, glass, biological nanopores, membranes with biological pore inserts, and combinations thereof.
In one embodiment, the first pore and the second pore are about 0.3 nm to about 100 nm in depth.
In one embodiment, the device comprises at least two electrodes for connecting to a power supply to generate a voltage across both of the first and second pores. In one embodiment, the voltages across the first and second pores are individually adjustable.
Also provided, in one embodiment, is a method for determining the sequence of a polynucleotide, which comprises:
(a) loading a sample comprising a polynucleotide,
(b) adjusting the voltages across the first and second pores so as to place the polynucleotide across both pores, wherein the polynucleotide moves across both pores at the same direction; and
(c) identifying a plurality of nucleotides of the polynucleotide using the sensor.
Provided as embodiments of this disclosure are drawings which illustrate by exemplification only, and not limitation, wherein:
Specifically,
In
Some or all of the figures are schematic representations for exemplification; hence, they do not necessarily depict the actual relative sizes or locations of the elements shown. The figures are presented for the purpose of illustrating one or more embodiments with the explicit understanding that they will not be used to limit the scope or the meaning of the claims that follow below.
Throughout this application, the text refers to various embodiments of the present nutrients, compositions, and methods. The various embodiments described are meant to provide a variety of illustrative examples and should not be construed as descriptions of alternative species. Rather it should be noted that the descriptions of various embodiments provided herein may be of overlapping scope. The embodiments discussed herein are merely illustrative and are not meant to limit the scope of the present invention.
Also throughout this disclosure, various publications, patents and published patent specifications are referenced by an identifying citation. The disclosures of these publications, patents and published patent specifications are hereby incorporated by reference into the present disclosure to more fully describe the state of the art to which this invention pertains.
As used in the specification and claims, the singular form “a”, “an” and “the” include plural references unless the context clearly dictates otherwise. For example, the term “an electrode” includes a plurality of electrodes, including mixtures thereof.
As used herein, the term “comprising” is intended to mean that the devices and methods include the recited components or steps, but not excluding others. “Consisting essentially of” when used to define devices and methods, shall mean excluding other components or steps of any essential significance to the combination. “Consisting of” shall mean excluding other components or steps. Embodiments defined by each of these transition terms are within the scope of this invention.
All numerical designations, e.g., distance, size, temperature, time, voltage and concentration, including ranges, are approximations which are varied (+) or (−) by increments of 0.1. It is to be understood, although not always explicitly stated that all numerical designations are preceded by the term “about”. It also is to be understood, although not always explicitly stated, that the components described herein are merely exemplary and that equivalents of such are known in the art.
One embodiment of the present disclosure provides a device that includes a plurality of chambers, each chamber in communication with an adjacent chamber through at least one pore. Among these pores, two pores, namely a first and a second pores, are placed to allow at least a portion of a polymer to move out of the first pore and into the second pore. Further, the device includes a sensor capable of identifying the polymer during movement. In one aspect, the identification entails identifying individual components of the polymer. Preferrably, when a single sensor is employed, the single sensor does not include two electrodes placed at both ends of a pore to measure an ionic current across the pore.
In one aspect, the device includes three chambers connected through two pores. Devices with more than three chambers can be readily designed to include one or more additional chambers on either side of a three-chamber device, or between any two of the three chambers. Likewise, more than two pores can be included in the device to connect the chambers.
In one aspect, there can be two or more pores between two adjacent chambers, to allow multiple polymers to move from one chamber to the next simultaneously. Such a multi-pore design can enhance throughput of polymer analysis in the device.
In some aspects, the device further includes means to enable movement of a polymer from one chamber to another. In one aspect, the movements results in loading the polymer across both the first pore and the second pore at the same time. In another aspect, the means further enables the movement of the polymer, through both pores, at the same direction.
For instance, in a three-chamber two-pore device (a “two-pore” device), each of the chambers can contain an electrode for connecting to a power supply so that a separate voltage can be established across each of the pores between the chambers. In accordance with one embodiment of the present disclosure, provided is a device comprising an upper chamber, a middle chamber and a lower chamber, wherein the upper chamber is in communication with the middle chamber through a first pore, and the middle chamber is in communication with the lower chamber through a second pore.
With reference to
Each of the pores (111 and 112) independently has a size that allows a small or large molecule or microorganism to pass. In one aspect, each pore is at least about 1 nm in diameter. Alternatively, each pore is at least about 2 nm, 3 nm, 4 nm, 5 nm, 6 nm, 7 nm, 8 nm, 9 nm, 10 nm, 11 nm, 12 nm, 13 nm, 14 nm, 15 nm, 16 nm, 17 nm, 18 nm, 19 nm, 20 nm, 25 nm, 30 nm, 35 nm, 40 nm, 45 nm, 50 ram, 60 nm, 70 nm, 80 nm, 90 nm or 100 nm in diameter.
In one aspect, the pore is no more than about 100 nm in diameter. Alternatively, the pore is no more than about 95 nm, 90 nm, 85 nm, 80 nm, 75 nm, 70 nm, 65 nm, 60 nm, 55 nm, 50 nm, 45 nm, 40 nm, 35 nm, 30 nm, 25 nm, 20 nm, 15 or 10 nm in diameter.
In one aspect, the pore has a diameter that is between about 1 nm and about 100 nm, or alternatively between about 2 nm and about 80 nm, or between about 3 nm and about 70 nm, or between about 4 nm and about 60 nm, or between about 5 nm and about 50 nm, or between about 10 nm and about 40 nm, or between about 15 nm and about 30 nm.
In some aspects, the pore has a substantially round shape. “Substantially round”, as used here, refers to a shape that is at least about 80 or 90% in the form of a cylinder. In some embodiments, the pore is square, rectangular, triangular, oval, or hexangular in shape.
Each of the pores (111 and 112) independently has a depth. In one aspect, each pore has a depth that is least about 0.3 nm. Alternatively, each pore has a depth that is at least about 0.6 nm, 1 nm, 2 nm, 3 nm, 4 nm, 5 nm, 6 nm, 7 nm, 8 nm, 9 nm, 10 nm, 11 nm, 12 nm, 13 nm, 14 nm, 15 nm, 16 nm, 17 nm, 18 nm, 19 nm, 20 nm, 25 nm, 30 nm, 35 nm, 40 nm, 45 nm, 50 nm, 60 nm, 70 nm, 80 nm, or 90 nm.
In one aspect, each pore has a depth that is no more than about 100 nm. Alternatively, the depth is no more than about 95 nm, 90 nm, 85 nm, 80 nm, 75 nm, 70 nm, 65 nm, 60 nm, 55 nm, 50 nm, 45 nm, 40 nm, 35 nm, 30 nm, 25 nm, 20 nm, 15 or 10 nm.
In one aspect, the pore has a depth that is between about 1 nm and about 100 nm, or alternatively between about 2 nm and about 80 nm, or between about 3 nm and about 70 nm, or between about 4 nm and about 60 nm, or between about 5 nm and about 50 nm, or between about 10 nm and about 40 nm, or between about 15 nm and about 30 nm.
In one aspect, the pores are spaced apart at a distance that is between about 10 nm and about 1000 nm. In one aspect, the distance is at least about 10 nm, or alternatively at least about 20 nm, 30 nm, 40 nm, 50 nm, 60 nm, 70 nm, 80 nm, 90 nm, 100 nm, 150 nm, 200 nm, 250 nm, or 300 nm. In another aspect, the distance is no more than about 1000 nm, 900 nm, 800 nm, 700 nm, 600 nm, 500 nm, 400 nm, 300 nm, 250 nm, 200 nm, 150 nm, or 100 nm. In yet another aspect, the distance is between about 20 nm and about 800 nm, between about 30 nm and about 700 nm, between about 40 nm and about 500 nm, or between about 50 nm and about 300 nm.
The two pores can be arranged in any position so long as they allow fluid communication between the chambers and have the prescribed size and distance between them. In one aspect, the pores are placed so that there is no blockage directly between them. Still, in one aspect, the pores are substantially coaxial, as illustrated in
In one aspect, the device, through the electrodes in the chambers, is connected to one or more power supply. In some aspects, the power supply is comprised of a voltage-clamp or a patch-clamp for supplying the voltage across each pore, which can also measure the current through each pore independently. In this respect, the power supply can set the middle chamber to a common ground for both voltage sources. In one aspect, the power supply is configured to provide a first voltage between the upper chamber (e.g., Chamber A in
In some aspects, the first voltage and the second voltage are independently adjustable. In one aspect, the middle chamber is adjusted to be ground relative to the two voltages (illustrated in
Adjustment of the voltages can be used to control the movement of charged particles in the chambers. For instance, when both voltages are set in the same direction, a properly charged particle can be moved from the upper chamber to the middle chamber and to the lower chamber, or the other way around, sequentially. Otherwise, a charged particle can be moved from either the upper or the lower chamber to the middle chamber and kept there.
The adjustment of the voltages in the device can be particularly useful for controlling the movement of a large molecule, such as a charged polymer, that is long enough to cross both of the pores at the same time. In such an aspect, the movement and the rate of movement of the molecule can be controlled by the relative magnitude and direction of the voltages, which will be further described below.
The device can contain materials suitable for holding liquid samples, in particular, biological samples, and/or materials suitable for nanofabrication. In one aspect, such materials include dielectric materials such as, but not limited to, silicon, silicon nitride, silicon dioxide, graphene, carbon nanotubes, TiO2, HfO2, Al2O3, or other metallic layers, or any combination of these materials. A single sheet of graphene forms a membrane ˜0.3 nm thick, and can be used as the pore-bearing membrane, for example.
Devices that are microfluidic and that house two-pore microfluidic chip implementations can be made by a variety of means and methods. For a microfluidic chip comprised of two parallel membranes, both membranes can be simultaneously drilled by a single beam to form two concentric pores, though using different beams on each side of the membranes is also possible in concert with any suitable alignment technique. In general terms, the housing ensures sealed separation of Chambers A-C. In one aspect, the housing would provide minimal access resistance between the voltage electrodes (two sources and one ground) and the nanopores, to ensure that each voltage is applied principally across each pore.
In one aspect,
More specifically, the pore-bearing membranes can be made with TEM (transmission electron microscopy) grids with 5-100 nm thick silicon, silicon nitride, or silicon dioxide windows. Spacers can be used to separate the membrane, using an insulator (SU-8, photoresist, PECVD oxide, ALD oxide, ALD alumina) or an evaporated metal (Ag, Au, Pt) material, and occupying a small volume within the otherwise aqueous portion of Chamber B between the membranes. The holder is seated in an aqueous bath that comprises the largest volumetric fraction of Chamber B. Chambers A and C are accessible by larger diameter channels (for low access resistance) that lead to the membrane seals.
A focused electron or ion beam can be used to drill pores through the membranes, naturally aligning them. The pores can also be sculpted (shrunk) to smaller sizes by applying the correct beam focus to each layer. Any single nanopore drilling method can also be used to drill the pair of pores in the two membranes, with consideration to the drill depth possible for a given method and the thickness of the membranes. Predrilling a micro-pore to a prescribed depth and then a nanopore through the remainder of the membranes is also possible, to further refine membrane thicknesses.
In another aspect, insertion of biological nanopores into solid-state nanopores to form a hybrid pore can be used in either or both nanopores in the two-pore method (Hall et al., Nat. Nanotech., 5(12):874-7, 2010). The biological pore can increase the sensitivity of the ionic current measurements, and are useful when only single-stranded polynucleotides are to be captured and controlled in the two-pore device, e.g., for sequencing.
By virtue of the voltages present at the pores of the device, charged molecules can be moved through the pores between chambers. Speed and direction of the movement can be controlled by the magnitude and direction of the voltages. Further, because each of the two voltages can be independently adjusted, the movement and speed of a charged molecule can be finely controlled in each chamber.
For instance, the device can be used to admix two positively charged molecules in a controlled manner. To this end, the first molecule is initially loaded in the upper chamber and the second in the lower chamber. A first voltage across the first port can induce movement of the first molecule into the middle chamber from the upper chamber. Likewise, a second voltage, in the opposite direction to the first voltage, can induce movement of the second molecule into the middle chamber from the lower chamber. Due to the opposite directions of the voltages, both molecules will be kept in the middle chamber so as to react with each other. Further, by adjusting the relative magnitudes of the voltages, the inflow speeds of each molecules can be fine tuned, leading to controlled reaction.
Another example concerns a charged polymer, such as a polynucleotide, having a length that is longer than the combined distance that includes the depth of both pores plus the distance between the two pores. For example, a 1000 bp dsDNA is ˜340 nm in length, and would be substantially longer than the 40 nm spanned by two 10 nm-length pores separated by 20 nm. In a first step, the polynucleotide is loaded into either the upper or the lower chamber. By virtue of its negative charge under a physiological condition (˜pH 7.4), the polynucleotide can be moved across a pore on which a voltage is applied. Therefore, in a second step, two voltages, in the same direction and at the same or similar magnitudes, are applied to the pores to induce movement of the polynucleotide across both pores sequentially.
At about time when the polynucleotide reaches the second pore, one or both of the voltages can be changed. Since the distance between the two pores is selected to be shorter than the length of the polynucleotide, when the polynucleotide reaches the second pore, it is also in the first pore. A prompt change of direction of the voltage at the first pore, therefore, will generate a force that pulls the polynucleotide away from the second pore (illustration in
If, at this point, the magnitude of the voltage-induced force at the first pore is less than that of the voltage-induced force at the second pore, then the polynucleotide will continue crossing both pores towards the second pore, but at a lower speed. In this respect, it is readily appreciated that the speed and direction of the movement of the polynucleotide can be controlled by the directions and magnitudes of both voltages. As will be further described below, such a fine control of movement has broad applications.
Accordingly, in one aspect, provided is a method for controlling the movement of a charged polymer through a pore. The method entails (a) loading a sample comprising a charged polymer in one of the upper chamber, middle chamber or lower chamber of the device of any of the above embodiments, wherein the device is connected to a power supply for providing a first voltage between the upper chamber and the middle chamber, and a second voltage between the middle chamber and the lower chamber; (b) setting an initial first voltage and an initial second voltage so that the polymer moves between the chambers, thereby locating the polymer across both the first and second pores; and (c) adjusting the first voltage and the second voltage so that both voltages generate force to pull the charged polymer away from the middle chamber (voltage-competition mode), wherein the two voltages are different in magnitude, under controlled conditions, so that the charged polymer moves across both pores in either direction and in a controlled manner.
For the purpose of establishing the voltage-competition mode in step (c), the relative force exerted by each voltage at each pore is to be determined for each two-pore device used, and this can be done with calibration experiments by observing the influence of different voltage values on the motion of the polynucleotide (motion can be measured by sensing location-known and detectable features in the polynucleotide, with examples of such features detailed later in this provisional document). If the forces are equivalent at each common voltage, for example, then using the same voltage value at each pore (with common polarity in upper and lower chambers relative to grounded middle chamber) creates a zero net motion in the absence of thermal agitation (the presence and influence of Brownian motion is discussed below). If the forces are not equivalent at each common voltage, then achieving equal forcing requires identification and use of a larger voltage at the pore that experiences a weaker force at the common voltage. Calibration for voltage-competition mode is required for each two-pore device, and would be required for specific charged polymers or molecules for which features that pass through each pore influence the force.
In one aspect, the sample containing the charged polymer is loaded into the upper chamber and the initial first voltage is set to pull the charged polymer from the upper chamber to the middle chamber and the initial second voltage is set to pull the polymer from the middle chamber to the lower chamber. Likewise, the sample can be initially loaded into the lower chamber.
In another aspect, the sample containing the charged polymer is loaded into the middle chamber and the initial first voltage is set to pull the charged polymer from the middle chamber to the upper chamber and the initial second voltage is set to pull the charged polymer from the middle chamber to the lower chamber.
The term “charged polymer” or “polymer” refers to a polymer that contains sufficient charged units at the pH of the solution that it can be pulled through a pore by electrostatic forces. In one embodiment, each unit of the charged polymer is charged at the pH selected. In another embodiment, the charged polymer is comprised of sufficient charged units to be pulled into and through the pores by electrostatic forces. For example, a peptide containing sufficient entities which can be charged at a selected pH (lysine, aspartic acid, glutamic acid, etc.) so as to be used in the devices and methods described herein is a charged polymer for purposes of this invention. Likewise, a copolymer comprising methacrylic acid and ethylene is a charged polymer for the purposes of this invention if there is sufficient charged carboxylate groups of the methacrylic acid residue to be used in the devices and methods described herein is a charged polymer for purposes of this invention. In one embodiment, the charged polymer is comprised one or more charged units at or close to one terminus of the polymer. In another embodiment, the charged polymer is comprised of one or more charged units at or close to both termini of the polymer.
In some aspects, the charged polymer is a polynucleotide or a polypeptide. In a particular aspect, the charged polymer is a polynucleotide. Non-limiting examples of polynucleotides include double-stranded DNA, single-stranded DNA, double-stranded RNA, single-stranded RNA, and DNA-RNA hybrids.
In one aspect, the adjusted first voltage and second voltage at step (c) are about 10 times to about 10,000 times as high, in magnitude, as the difference between the two voltages. For instance, the two voltages are 90 mV and 100 mV, respectively. The magnitude of the voltages (˜100 mV) is about 10 times of the difference between them, 10 mV. In some aspects, the magnitude of the voltages is at least about 15 times, 20 times, 25 times, 30 times, 35 times, 40 times, 50 times, 100 times, 150 times, 200 times, 250 times, 300 times, 400 times, 500 times, 1000 times, 2000 times, 3000 times, 4000 times, 5000 times, 6000 times, 7000 times, 8000 times or 9000 times as high as the difference between them. In some aspects, the magnitude of the voltages is no more than about 10000 times, 9000 times, 8000 times, 7000 times, 6000 times, 5000 times, 4000 times, 3000 times, 2000 times, 1000 times, 500 times, 400 times, 300 times, 200 times, or 100 times as high as the difference between them.
In one aspect, real-time or on-line adjustments to first voltage and second voltage at step (c) are performed by active control or feedback control using dedicated hardware and software, at clock rates up to hundreds of megahertz. Automated control of the first or second or both voltages is based on feedback of the first or second or both ionic current measurements.
The device of the present disclosure can be used to carry our analysis of molecules or particles that move or are kept within the device. In one aspect, the device further includes one or more sensors to carry out the analysis. It is contemplated that, when a single sensor in used, the single sensor preferably does not include a pair of electrodes placed at two sides of a pore to measure an ionic current across the pore when a molecule or particle, in particular a polymer, moves through the pore.
In one aspect, the device includes a sensor suitable for identifying a polymer, or individual components of a polymer. Non-limiting examples of individual components include monomer units and monomer units with modifications approximate to or on the monomer unit. When the polymer is a polynucleotide, an individual component can be one or more nucleotide units or the nucleotide units bound by a protein factor, without limitation.
The sensors used in the device can be any sensor suitable for identifying a molecule or particle, such as a polymer. For instance, a sensor can be configured to identify the polymer by measuring a current, a voltage, pH, an optical feature or residence time associated with the polymer or one or more individual components of the polymer.
In one embodiment, the sensor measures an optical feature of the polymer or a component (or unit) of the polymer. One example of such measurement includes identification by infrared (or ultraviolet) spectroscopy of a absorption band unique to a particular unit.
When residence time measurements are used, they will correlate the size of the unit to the specific unit based on the length of time it takes to pass through the sensing device.
Still further, the sensor can include an enzyme distal to that sensor which enzyme is capable of separating the terminal unit of the polymer from the penultimate unit thereby providing for a single molecular unit of the polymer. The single molecule, such as a single nucleotide or an amino acid, can then be detected with methods such as mass spectrometry. Methods for measuring such a single unit are known in the art and include those developed by Cal Tech (see, e.g., http://spectrumieee.org/tech-talk/at-work/test-and-measurement/a-scale-for-weighing-single-molecules). The results of that analysis can be compared to those of the sensing device to provide confirmation of the correct analysis.
Accordingly, the sensor can be placed at a location within the device that enables such measurement (
Likewise, the sensor, with electrodes 161 and 162, can be placed at an exit of one of the pores (
In yet another aspect, the sensor does not include electrodes but takes a form of a membrane or scaffold (191) having an opening (e.g., a hole) that allows passing of a polymer. For instance, one or more layers of graphene membrane with an opening, which can be just like a nanopore as the other nanopores, can act as the auxiliary sensor, which itself is positioned in the chamber between the two pores. In one aspect, the graphene member comprises a single sheet, double sheet, or more than two sheets.
In some embodiments, the sensor is configured to proximate a polymer when the polymer is loaded in both the first and the second pores. Therefore, when the polymer moves through both pores, the sensor is close enough to the polymer to measure the polymer. In this respect, the sensor can be aligned with the two pores, either in one of the pores, or between them, as illustrated in
In some embodiments, the sensor is configured to form a tunnel gap allowing a polymer to pass through the tunnel gap. In the device as illustrated in
It has been shown that individual nucleotides can be discriminated in a precision 0.8 nm tunneling gap.
In some embodiments, the sensor is functionalized with reagents that form distinct non-covalent bonds with each DNA base. In this respect, the gap can be larger and still allow effective measuring. For instance, a 2.5 nm gap can be as effective, when used with a functionalized, as a 0.8 nm gap. Tunnel sensing with a functionalized sensor is termed “recognition tunneling.” Using a Scanning Tunneling Microscope (STM) with recognition tunneling, a DNA base flanked by other bases in a short DNA oligomer can be identified.
Recognition tunneling can also provide a “universal reader” designed to hydrogen-bond in a unique orientation to each of the four DNA bases (A, C, G, T) and also to the base 5-methyl-cytosine (mC) which is naturally occurring due to epigenetic modifications.
A limitation with the conventional recognition tunneling is that it can detect only freely diffusing DNA that randomly binds in the gap, or that happens to be in the gap during microscope motion, with no method of explicit capture to the gap. Further, the collective drawbacks of the STM setup will go away when the recognition reagent, once optimized for sensitivity, is incorporated within an electrode tunneling gap in a nanopore channel.
Accordingly, in one embodiment, the sensor comprises surface modification by a reagent. In one aspect, the reagent is capable of forming a non-covalent bond with a nucleotide. In a particular aspect, the bond is a hydrogen bond. Non-limiting examples of the reagent include 4-mercaptobenzamide and 1-H-Imidazole-2-carboxamide.
A significant advantage of the methods in the present technology is that is that they can be engineered, in principle, to provide direct tracking of progress through homopolymeric regions (base repeats). Direct base repeat tracking is not possible with ionic current sensing. Tracking repeats is essential, for example, since deletions and insertions of specific mononucleotide repeats (7, 9 nt) within human mitochondrial DNA have been implicated in several types of cancer.
In ionic current sensing, there is no distinct signal-per-nucleotide of motion of homopolymeric ssDNA through the pore. It is contemplated that an ideal nanopore sequencing platform should utilize an auxiliary sensing method that can track per-nucleotide motion progress while also achieving single-nucleotide sensitivity. Transitions between neighboring nucleotides in oligomers can be observable with recognition tunneling, making it a candidate for sequencing that permits direct base-repeat tracking.
Therefore, the methods of the present technology can provide DNA delivery rate control for one or more recognition tunneling sites, each positioned in one or both of the nanopore channels, and voltage control can ensure that each nucleotide resides in each site for a sufficient duration for robust identification.
Sensors in the devices and methods of the present disclosure can comprise gold, platinum, graphene, or carbon, or other suitable materials. In a particular aspect, the sensor includes parts made of graphene. Graphene can act as a conductor and an insulator, thus tunneling currents through the graphene and across the nanopore can sequence the translocating DNA.
In some embodiments, the tunnel gap has a width that is from about 1 nm to about 20 nm. In one aspect, the width of the gap is at least about 1 nm, or alternatively at least about 1.5, 2, 2.5, 3, 3.5, 4, 4.5, 5, 6, 7, 8, 9, 10, 12 or 15 nm. In another aspect, the width of the gap is not greater than about 20 nm, or alternatively not greater than about 19, 18, 17, 16, 15, 14, 13, 12, 11, 10, 9, 8, 7, 6, 5, 4, 3, or 2 nm. In some aspects, the width is between about 1 nm and about 15 nm, between about 1 nm and about 10 nm, between about 2 nm and about 10 nm, between about 2.5 nm and about 10 nm, or between about 2.5 nm and about 5 nm.
A polymer, such as a DNA molecule, can be analyzed in the device of the present disclosure. In one aspect, the polymer is loaded into at least two pores in the device, as described above. Once the polymer is loaded, it is at a position suitable for detection by the sensor, by means of measuring a current, a voltage, pH, an optical feature or residence time associated with the polymer or components of the polymer.
For instance, a polynucleotide can be loaded into both pores by two voltages having the same direction. In this example, once the direction of the voltage applied at the first pore is inversed and the new voltage-induced force is slightly less, in magnitude, than the voltage-induced force applied at the second pore, the polynucleotide will continue moving in the same direction, but at a markedly lower speed. In this respect, the amplifier supplying voltage across the second pore also measures current passing through the second pore, and the ionic current then determines the identification of a nucleotide that is passing through the pore, as the passing of each different nucleotide would give rise to a different current signature (e.g., based on shifts in the ionic current amplitude). Identification of each nucleotide in the polynucleotide, accordingly, reveals the sequence of the polynucleotide.
In some aspects, repeated controlled delivery for re-sequencing a polynucleotide further improves the quality of sequencing. Each voltage is alternated as being larger, for controlled delivery in each direction. Also contemplated is that the two currents through the two pores can be correlated to improved accuracy. It is contemplated that Brownian motion may cause fluctuations in the motion of a molecule, affecting controlled delivery of the molecule. Such an effect, however, can be minimized or avoided by, e.g., during DNA sequencing, repeated controlled delivery of the DNA and averaging the sequencing measurements. Still further, it is contemplated that the impact of Brownian motion on the controlled motion of large molecules, such as polynucleotides and polypeptides, would be insignificant in particular when competing forces are pulling the larges molecules apart, generating tension within the molecule. It is contemplated that adhesion of the DNA to the pore walls, by surface charge modifications or chemistry to the pore surface, to create friction can also mitigate the influence of Brownian motion on the control performance of the two pore method.
Such a method provides a ready solution to problems that have not been solved in the prior art.
For instance, it is known that there are two competing obstacles to achieve the controlled delivery and accurate sensing required for nanopore sequencing. One is that a relatively high voltage is required, at the pore, to provide enough sequencing sensitivity. On the other hand, high voltages lead to fast passing of a polynucleotide through the pore, not allowing sufficient time for identification of each nucleotide.
More specifically, the nanopore sequencing platform requires that the rate of polynucleotide passage through the pore be controlled to 1 ms/nucleotide (nt) or slower, while still generating a sequence-sensitive current. This requires sufficiently high signal-to-noise for detecting current signatures (high voltage is better), but sufficiently slow motion of the molecule through the pore to ensure measurements are within the recording bandwidth (low voltage is better). In single pore implementations, polynucleotide speed is proportional to voltage, so higher voltage is better for sensing but worse for reducing polynucleotide speed: rates are 1 μs/nt and faster (>1000 times too fast) at voltages that promote polynucleotide capture. On the other hand, lower voltages reduce sensing performance, and also increase the relative contribution of rate fluctuations caused by Brownian motion that will undermine read accuracy.
Other than what is described herein, there are currently no methods for addressing these obstacles that do not involve the use of enzymes or optics, both of which work only in specialized nanopore techniques.
Several approaches have been proposed to address the problem associated with the lack of sensing capability, and under low voltages. One is to engineer biological nanopores to improve their sensitivity. Another is to use graphene membranes, which as a single sheet are thinner than the distance between nucleotides in ssDNA. Still another is the use of an auxiliary current measured in close proximity to the nanopore (e.g., tunneling currents).
Biological nanopores have been tested in the first approach. The a-hemolysin nanopore is the most commonly used biological pore in research. Studies have shown that α-hemolysin can resolve single nucleotide substitutions in homopolymers and abasic (1′,2′-dideoxy) residues within otherwise all-nucleobase DNA. However, single nucleotide sensitivity is not possible in heteromeric DNA with wild-type (WT) α-hemolysin, for which the ionic current is influenced by ˜10 nucleotides in the channel. Protein engineering of α-hemolysin has been used to improve its sensitivity for DNA analysis and sequencing. One such mutant pore uses α-hemolysin with a covalently attached molecular adapter (Clarke et al., Nat. Nanotech, 4(4):265-70, 2009) that is capable of discriminating the four nucleoside 5′-monophosphate molecules with high accuracy. However, this mutant pore does not appear to have sensitivity for sequencing intact heteromeric ssDNA that passes through the pore.
Another exemplary biological pore is MspA, which has a funnel-like shape that focuses the sensitivity of the ionic current to the bottom of the channel. Moreover, achieving rate reduction of DNA through MspA and α-hemolysin can be achieved by using enzymes. As shown in
Presently, there is no nanopore for which ionic current sensing can provide single-nucleotide sensitivity for nucleic acid sequencing. Still, improvements to the sensitivity of biological pores and solid-state pores (graphene) are active and ongoing research fields. One issue is that ionic current sensing does not permit direct tracking of progress through homolymeric regions (base repeats), since there is no distinct signal-per-nucleotide of motion of homopolymeric ssDNA through the pore. Tracking repeats is essential, for example, since deletions and insertions of specific mononucleotide repeats (7, 9 nt) within human mitochondrial DNA have been implicated in several types of cancer (Sanchez-Cespedes, et al., Cancer Research, 61(19):7015-7019, 2001). While enzymes on biopores can reduce the rate of translocation, there is lack of control over the dwell time of each nucleotide. On the other hand, using a constant delivery rate with two-pore control, non-deterministic pauses are eliminated, and accurate estimation of repeat lengths can be made. Re-reading the repeat section many times can also improve the estimation errors and identify error bounds, and this can be done without having to reverse the polymerization chemistry caused by enzymes.
A recent study showed that, with a single nanopore, reduced rates cannot be achieved by merely reducing the voltage (Lu et al., Biophysical Journal, 101(1):70-9, 2011). Instead, as voltage is reduced, the rates of a single-stranded DNA (ssDNA) become more random (including backtracking), since Brownian motion becomes an increasing contributor to velocity fluctuations. The study also shows that high voltage force is required to suppress Brownian-motion induced velocity fluctuations that will otherwise confound sequencing measurements, even when using an idealized single-nucleotide-sensitive nanopore sensor.
The sequencing method provided in the present disclosure, based on a two-pore device, provides a ready solution to these problems and additional advantages over the existing methods. In concert with one or two pores that have sufficient sensitivity for sequencing, at high or low voltage, the two pore control solves the sequencing rate control problem of single nanopore implementations. Such pores can include biological pores housed in solid-state substrates, biological pores in membranes formed across solid-state substrates, or solid-state pores (e.g., in graphene, silicon, or other substrates). In one aspect, an enzyme such as phi29 on a biological pore such as MspA can be used at one or both pores, with high voltages used to generate large signals for sequencing and a low differential voltage that generates a force on each enzyme that is sufficient to hold the enzymes in position atop each pore and permit polymerization-catalyzed DNA motion, but not large enough to stall or dissociate the enzymes. Such a configuration can improve the methods in Cherf et al., Nat. Biotech., 30(4):344-8, 2012 and Manrao et al., Nature Biotechnology, 30:349-53, 2012, by significantly boosting the measurement signal, and permitting two pores to read one stand of DNA at the same time.
In addition, the method of the present disclosure can generate sufficiently high voltage at the pore to ensure detection sensitivity at the pore using ionic current sensing. It is plausible that high voltage would suppress Brownian motion enough to ensure constant rates through each pore, and configuration of the DNA outside each pore will affect the energetics of motion of DNA in either direction. Additionally, the voltage competition used in the method (
The method can be used to identify the composition of monomers in a charged polymer. In one aspect, the monomer unit is a nucleotide when the polymer is a single stranded DNA or RNA. In another aspect, the monomer unit can be a nucleotide pair, when the polymer is double stranded.
In one aspect, the method can be used to identify a modification to the polymer, such as a molecule bound to a monomer, in particular when the bound molecule is charged. The bound molecule does not have to be charged, however, as even a neutral molecule can change the ionic current by virtue of its size.
In another aspect, the modification comprises the binding of a molecule to the polymer. For instance, for a DNA molecule, the bound molecule can be a DNA-binding protein, such as RecA, NF-κB and p53. In yet another aspect, the modification is a particle that binds to a particular monomer or fragment. For instance, quantum dots or fluorescent labels bound to a particular DNA site for the purpose of genotyping or DNA mapping can be detected by the device. Accordingly, the device of the present disclosure provides an inexpensive way for genotyping and DNA mapping as well, without limitation.
In one aspect, the polymer is attached to a solid support, such as a bead, at one end of the polymer.
Also provided, in one embodiment, is a method for determining the sequence of a polynucleotide, comprising: (a) loading a sample comprising a polynucleotide in the upper chamber of the device of any of the above embodiments, wherein the device is connected to a power supply for providing a first voltage between the upper chamber and the middle chamber, and a second voltage between the middle chamber and the lower chamber, wherein the polynucleotide is optionally attached to a solid support at one end of the polynucleotide; (b) setting an initial first voltage and an initial second voltage so that the polynucleotide moves from the upper chamber to the middle chamber and from the middle chamber to the lower chamber, thereby locating the polymer across both the first and second pores; (c) adjusting the first voltage and the second voltage so that both voltages generate force to pull the polynucleotide away from the middle chamber, wherein the two voltages are different in magnitude, under controlled conditions, so that the polynucleotide moves across both pores in one direction and in a controlled manner; and (d) identifying each nucleotide of the polynucleotide that passes through one of the pores, by measuring an ionic current across the pore when the nucleotide passes that pore.
The present technology is further defined by reference to the following examples. It will be apparent to those skilled in the art that many modifications, both to threads and methods, may be practiced without departing from the scope of the current invention.
This example shows that capture of DNA into each pore in a two-pore device is readily detected as shift in each independent ionic pore current measured.
This example demonstrates dual-pore capture using dsDNA with and without a bead attached to one end. Experiments with bead-tethered ssDNA can also be explored.
Upon capture and stalling of the DNA, the pore voltage nearest the bead (V1,
When a bead is used, the bead has a proper size that prevents the bead from passing either or both of the pores. Methods that ensure a 1 to 1 bead-DNA ratio have been developed in the art. For example, monovalent streptavidin-coated Quantum dots (QDs; QD655, Invitrogen) conjugated to biotinylated DNA duplexes (or ssDNA) can provide beads in the 20-30 nm diameter range, with larger beads (30-100 nm) possible by using gold particles or latex. The influence of bead on hydrodynamics and charge, as it relates to capture rate, can be considered in designing the experiments.
Without beads, dsDNA passes through a pore at ˜0.1 ms/kbp. DNA of lengths 500 bp and 4 kbp, and λ-phage dsDNA molecules (˜48 kbp) can be used. DNA samples can be delivered from chamber A into both pores, using a common voltage polarity for each pore to promote capture from chamber A and passage through chamber B into chamber C (
Using nanopore diameters 10 nm and larger minimizes the interaction (e.g., friction and sticking) between dsDNA and the nanopore walls. For larger pores, although dsDNA can be captured in an unfolded and folded configurations, the single-file (unfolded) configuration is more likely at higher voltages, and with shorter (kbp) dsDNA. For an inter-pore distance of 500 nm or less, it is contemplated that the probability of dual-pore capture, following capture at the first pore (between chambers A and B) is very high, for voltages of at least 200 mV in 1 M KCl.
The radial distance within which voltage influence dominates thermal diffusion, and leads to capture with high likelihood, has been estimated to be at least 900 nm (larger than the inter-pore distance) for a range of pore sizes (6-15 nm diameter), voltages (120-500 mV), and with dsDNA at least 4 kbp in length (Gershow and Golovchenko, Nature Nanotechnology, 2:775-779, 2007). These findings support a high likelihood of prompt dual-pore capture of dsDNA, following single (first) pore capture of the dsDNA.
The capture and control of DNA through the two pores can benefit from active control hardware and real-time algorithms. The inventors have developed active control hardware/software for DNA control. See, for example, Gyarfas et al, Biophys. J., 100:1509-16, 2011); Wilson et al., ACS Nano., 3(4):995-1003, 2009; and Benner et al., Nat. Nanotech., 2(11):718-24, 2007. A useful software is the LabVIEW software (Version 8, National Instruments, Austin, Tex.), implemented on an FPGA (field-programmable gate array) system (PCI-7831 R, National Instruments)). The referenced FPGA can control up to 4 amplifiers simultaneously. Further, the Axon Digidata 1440A Data Acquisition System used to digitize and log data onto a PC has 16 input channels, enough to record voltage and current for up to 8 amplifiers in parallel. Other real-time operating system in concert with hardware/software for real-time control and measurement could also be used for controlling the amplifiers, and digitizing and logging the data.
The inventors have also developed a low-noise voltage-clamp amplifier termed the “Nanoclamp,” (Kim et al., IEEE Trans. On Biom. Circ. And Syst. In press, May 2012; Kim et al., Elec. Lette., 47(15):844-6, July, 2011; and Kim et al., Proceedings of the IEEE International SoC Design Conference (ISOCC), November, 2010) to functionalize and optimize the use of one or more nanopores in small-footprint and multi-channel devices. Any other commercial bench-top voltage-clamp or patch-clamp amplifier, or integrated voltage-clamp or patch-clamp amplifier could be used for two pore control and measurement.
For a variety of solid-state pore materials and diameters, 0.1-10 kbp takes ˜1 ms to translocate. With a FPGA-controlled amplifier, one can detect capture and initiate competing voltage control within 0.020 ms, much faster than the 1 ms total passage time of 1 kbp DNA; thus, triggering the control method before DNA escapes (with no bead attachment) also has high likelihood. As demonstration of control, the time to, and direction of, exit of the molecule from the pores can be demonstrated as a function of the magnitude of and difference between the competing voltages (
Force uncertainty induced by random transverse DNA motion is likely minimal. Additionally, the voltage force causes an electroosmotic flow (EOF) in the opposite direction of DNA motion, causing the DNA to move slower than it would in the absence of the induced counterion flow. Since different radial positions of the molecule can give rise to different EOF fields in the nanopore, one issue is whether the effective charge density and therefore the net driving force vary enough during fluctuations in DNA radial position to induce speed fluctuations. It is believed that the effective charge density of DNA in 1M KCl is stable for a distance of 1 nm or more between the pore wall and the DNA.
Additionally, SiN nanopores have a negative surface charge that intrinsically repels DNA. Thus, although the molecule will undergo radial position fluctuations, by using SiN pores with diameter greater than a few nanometers, it is likely that each constant voltage value will result in a constant effective force at each of the two pores, and thus a constant velocity in the direction of larger force when using two competing voltages in the two-pore setup. Treatment of other pore material surfaces can produce comparable effects to that of SiN.
Velocity uncertainty induced by random translational DNA motion that is caused by Brownian motion may be reduced by increasing the competing voltages. Experiments can be carried out to determine whether such reduction will occur. A single-nanopore study (Lu, et al., Biophysical Journal, 101(1):70-79, 2011) supports that increasing the competing forces can reduce uncertainty caused by Brownian motion. The study analyzed the velocity fluctuations caused by Brownian motion, which occur on fast (nanosecond) time scales, and the sequencing errors that result from such fluctuations. Assuming a hypothetical and idealized single-nucleotide sensor (noise-free detection at >22 MHz bandwidth), Brownian motion alone results in 75% read error. The relevant parameter for predicting the error is kBT/F*(0.34 nm), which is the ratio of thermal energy to the work done to translocate the DNA the distance a between nucleotides (0.34 nm). In the ratio, force F=Vλ is the voltage V driving DNA with charge density λ (0.2 e−/bp for dsDNA). For the present control method, increasing the voltage 50× results in 5% read error, with higher voltage further improving errors. With a single pore, however, since mean velocity
To maintain the 22 MHz bandwidth, a 50× increase in force with a single nanopore would have to be paired with a 50× increase in solution viscosity to maintain the same
This example shows that the two-pore device can be used to map the binding of a DNA-binding protein to dsDNA, and for proteins that have or do not bind to specific sequences.
As demonstrated in Example 1, DNA samples can be captured from Chamber A. RecA protein catalyses an ATP-dependent DNA strand-exchange reaction that pairs broken DNA with complementary regions of undamaged DNA. Using a poorly hydrolyzable ATP analogue ATP γS, RecA filaments bound to dsDNA are very stable in high salt (e.g., 1M KCl) when first assembled in physiological salt. As an alternative to ATPγS, which is slowly hydrolyzed, this example can also use ADP-AlF4 (aluminum tetrafluoride), which does not turnover at all, and causes RecA to bind more tightly to the DNA.
Detection of RecA filaments bound to λ-DNA through 20-30 nm nanopores has been demonstrated (Kowalczyk et al., Nano Lett., 10(1):324-8, 2010; Smeets et al., Nano Lett., 9(9):3089-95, 2009; and Hall et al. Nano Lett., 9(12):4441-5, 2009], but filaments <20 bp (6 or fewer RecA proteins) in length cannot be resolved using a single nanopore, due to the coupling between translocation rate and measurement SNR.
Initial experiments of this example use bead-bound and unbound λ-DNA that has been exposed to varying concentrations of RecA, to generate DNA that is nearly uncoated, partially coated, and fully coated. Real-time monitoring of each pore current can be used to gauge progress of the controlled delivery, and will be correlated for location mapping of the filaments. Repeated measurements of each DNA will improve accuracy of RecA mapping.
The added charge and bulk, and stability in high salt, when RecA is bound to DNA make it an ideal candidate to attempt detection and location mapping during controlled delivery with the proposed instrument.
Control of RecA-bound DNA can also be attempted without a bead attached to arrest translocation. As with dsDNA experiments in Example 1, active voltage control can be used to promptly initiate competing voltage control before the DNA exits the nanopores. As charged species that bind to DNA affect the mobility of DNA in an electric field, by altering the net charge and stiffness of the DNA, motion control tuning experiments can examine the influence of RecA binding to dsDNA on the force balance used to control the motion of the dsDNA.
This example can demonstrate that the shortest observed filament lengths, at low RecA concentrations, can be measured at high SNR and at sufficiently slow and controlled rates, so that any RecA protein bound in isolation can be detected if present.
The two-pore device therefore provides a completely new single-molecule instrument for basic research, as one could examine the capability to detect binding of additional proteins to the RecA-DNA filament, which would increase the filament width and thus be detected by a decrease in observed current. For example, proteins that bind to the RecA-DNA filament include LexA and bacteriophage lambda repressors, which use RecA to sense the status of the cell and switch on or off downstream regulatory events.
Calibration experiments would involve detecting proteins that bind to specific sequences (locations) on the DNA, so that protein-induced shifts in the current would then permit estimation of the speed and rate control performance of the DNA through the pores. Example proteins that bind to specific sites on dsDNA include Lac repressor (binds to a 21 bp segment), phage lambda repressor (which has multiple operator sites on λ-DNA), and other proteins.
This example demonstrates the production of up to 10 kb ssDNA with doubled-stranded segments of varying lengths.
In a first step, 10 kbp dsDNA can be created by long range PCR. One end of the strand is biotinylated for bead attachment, and the strands are separated by chemical denaturing. The unbeaded 10 kb ssDNA then serves as the measured strand in two-pore experiments. Complementary single-stranded segments with desired sizes can be created by PCR followed by bead capturing and strand separation.
ssDNA segments of varying lengths and at multiple sites within the measured 10 kb ssDNA can be used, starting with a set of 100 nt segments. Ionic current through a single solid-state pore was used to differentiate dsDNA from ssDNA homopolymers, and purine and pyrimidine homopolymers in (Skinner et al., Nano Lett., 9(8):2953-60, January 2009). Thus, likelihood of differentiating single from double stranded segments in DNA is high at sufficiently high voltage using the two-pore device. Mapping ssDNA vs. dsDNA segments enables nanopore sequencing using the hybridization-assisted method (though this method as proposed relies on a costly hybridization-assisted process), and can be used reveal both location and identity of target DNA sequences over long distances (targeted sequencing). One can also explore the use of Single Strand DNA Binding (SSB) proteins, as beads that will further amplify the ssDNA vs. dsDNA differences in ionic current by binding to the ssDNA and creating a larger impedance than dsDNA.
This example demonstrates the capture and control of a long ssDNA and the detection and localization of a RecA filament bound to the ssDNA. Additionally, it shows that the two-pore device can detect purine vs. pyrimidine homopolymeric segments within the ssDNA.
Stochastic detangling of 7 kb ssDNA through a 10 nm pore in a 20 nm SiN membrane can be carried out as shown in Stefan et al., Nano Lett., 10:1414-20, 2010. While the single-nanopore method in Stefan et al. 2010 unravels the ssDNA by the mechanical contact force between the tangled ssDNA and the pore/membrane surface, it is contemplated that the dual-pore competing voltage setup can electrophoretically force ssDNA to detangle near and in between the pores at sufficiently high competing voltages, by the action of each voltage force on the DNA backbone nearest each pore.
Detangling and subsequently precision control of the rate of ssDNA through the two pore setup is important for eventual sequencing of long ssDNA molecules. At sufficiently high voltage (˜400 mV), it is possible to discriminate purine and pyrimidine homopolymeric segments within ss-DNA (Skinner et al., Nano Lett., 9(8):2953-60, January 2009), which is valuable for diagnostic applications and possibly cancer research.
This example also explores the use of RecA, or perhaps other Single Strand DNA Binding (SSB) proteins, as detectable “speed-bumps” that are differentiable from the ssDNA ionic current by binding to the ssDNA and creating a larger impedance. These speed bumps will allow direct quantification of the controlled ssDNA speeds that are possible, which in turn will demonstrate that the required 1 ms/nt is achievable. Since RecA is not required to bind to specific trinucleotide sequence sites, but binds preferentially to TGG-repeating sequences and also tends to bind where RecA filaments are already formed, calibration experiments will require the use of other ssDNA binding molecules that do bind to specific known sequence locations. Having known binding sites that are detectable as they pass through each pore is required to determine the speed of the molecule as a function of the competing voltage values. A non-limiting example is to use duplex strands (or bead-tethered duplex strands) that hybridize to one or more known sites, from which the shifts in current could be used to detect passage of each duplex through each pore, and then estimate the passing strand speed for the chosen voltage values. Subsequently, RecA filaments can be formed and detected on such molecules, keeping the duplex feature(s) as benchmark detection points relative to which RecA filaments can be detected and their position inferred.
Methods for determining genetic haplotypes and DNA mapping by incorporating fluorescent labels into dsDNA (Xiao, et al., U.S. Pat. No. 7,771,944 B2, 2010) can also use the two pore device, since the bead labels (e.g., quantum dots, or any fluorescent label) is bulkier and will produce shifts in the current just as binding proteins on dsDNA would. Moreover, the two-pore method is simpler and much less expensive than using high resolution imaging methods (i.e., total internal reflection fluorescence microscopy) to detect and map the label positions. It is also noted that any velocity fluctuations caused by Brownian motion during controlled delivery are much less deleterious for detecting larger features (proteins, duplex segments, bead attachments) than for detecting smaller features.
It is to be understood that while the invention has been described in conjunction with the above embodiments, that the foregoing description and examples are intended to illustrate and not limit the scope of the invention. Other aspects, advantages and modifications within the scope of the invention will be apparent to those skilled in the art to which the invention pertains.
