Patent Application
20250146973
The present application relates to methods of sensing, including label-free methods of sensing.
Nanopores can be used to discriminate between analytes through the analysis of changes in conduction current profiles during translocation. However, the translocation times of analytes through some nanopore-based sensors are extremely fast, which limits the fidelity of electrical data that can be collected. Additionally, some nanopore-based sensing methods provide only certain types of data regarding analytes. There is a need for improved methods of sensing using nanopores, including methods that provide additional data regarding analytes and that can differentiate between additional types of analytes.
In one aspect, methods of sensing are described herein. In some embodiments, such a method comprises providing a sensor. The sensor comprises a first layer having at least one single nanohole structure or at least one dual nanohole structure, and a second layer having at least one nanopore. The single nanohole structure comprises only one nanohole. The dual nanohole structure comprises a first nanohole and a second nanohole connected by a gap. Additionally, the one nanohole (in the case of the single nanohole structure) or the gap (in the case of a dual nanohole structure) of the first layer is aligned with the nanopore of the second layer in a direction corresponding to a translocation direction across the first and second layers.
Sensors described herein having a dual nanohole structure can have any construction, structure, or property of a dual nanohole sensor described hereinbelow. Similarly, a sensor described herein having a single nanohole structure (rather than a dual nanohole structure) can likewise have any construction, structure, or property of a sensor described below.
Turning again to the method steps, a method described herein further comprises providing a test sample comprising an analyte and contacting the test sample with the first layer of the sensor. The method also comprises irradiating the single nanohole structure or the dual nanohole structure of the sensor with a beam of electromagnetic radiation and optically trapping the analyte in the single nanohole structure or in the dual nanohole structure and/or in the gap of the first layer of the sensor.
Moreover, in some cases, the method further comprises applying a first electric field across the nanopore to draw one or more of the analytes into the nanopore, wherein the first electric field comprises a direct current (DC) electric field. The method also comprises applying a second electric field across the nanopore after applying the first electric field, wherein the second electric field comprises a pulsed, modulated, or alternating current (AC) electric field.
In addition, in some implementations, a method described herein further comprises measuring one or more analyte properties or other properties potentially associated with optical trapping or translocation of the analyte through the nanopore of the sensor. For example, in some cases, a method described herein further comprises measuring a change in current and/or phase across the nanopore during application of the second electric field while the analyte is optically trapped and/or during one or more translocation events of the analyte through the nanopore. A method described herein can also (or alternatively) comprise measuring at least one kinetic parameter of the analyte within the nanopore after removing or turning off the second electric field. Further, in some embodiments, at least one kinetic parameter is measured while the analyte decelerates or comes to a stop while optically trapped.
In addition, in some cases, measuring change in current and/or phase further comprises determining a charge of a translocating analyte. Moreover, in some instances, measuring change in current and/or phase further comprises determining a dielectric constant of a translocating analyte.
Methods of sensing described herein, in some implementations, further comprise measuring a surface plasmon resonance of the single nanohole structure or the dual nanohole structure after optically trapping the analyte in the single nanohole structure or in the dual nanohole structure and/or in the gap of the first layer of the sensor. Additionally, in some such embodiments, measuring the surface plasmon resonance further comprises determining the mass of the optically trapped analyte.
Methods described herein can be used for sensing a variety of analytes. For example, in some cases, the analyte comprises complexed and/or non-complexed biomolecules. In other instances, the analyte comprises a nanoparticle such as an inorganic nanoparticle.
Additional features and embodiments are further described in the detailed description which follows.
Embodiments described herein can be understood more readily by reference to the following detailed description, examples, and figures. Devices and methods described herein, however, are not limited to the specific embodiments presented in the detailed description, examples, and figures. It should be recognized that these embodiments are merely illustrative of the principles of the present invention. Numerous modifications and adaptations will be readily apparent to those of skill in the art without departing from the spirit and scope of the invention.
All publications, patents and patent applications mentioned in this specification are herein incorporated in their entirety by reference into the specification, to the same extent as if each individual publication, patent, or patent application was specifically and individually indicated to be incorporated herein by reference. In addition, citation or identification of any reference in this application shall not be construed as an admission that such reference is available as prior art to the present invention. To the extent that section headings are used, they should not be construed as necessarily limiting.
In addition, all ranges disclosed herein are to be understood to encompass any and all subranges subsumed therein. For example, a stated range of “1.0 to 10.0” should be considered to include any and all subranges beginning with a minimum value of 1.0 or more and ending with a maximum value of 10.0 or less, e.g., 1.0 to 5.3, or 4.7 to 10.0, or 3.6 to 7.9.
All ranges disclosed herein are also to be considered to include the end points of the range, unless expressly stated otherwise. For example, a range of “between 5 and 10,” “from 5 to 10,” or “5-10” should generally be considered to include the end points 5 and 10.
Further, when the phrase “up to” is used in connection with an amount or quantity, it is to be understood that the amount is at least a detectable amount or quantity. For example, a material present in an amount “up to” a specified amount can be present from a detectable amount and up to and including the specified amount.
It is also to be understood that the article “a” or “an” refers to “at least one,” unless the context of a particular use requires otherwise.
Additionally, in any disclosed embodiment, the terms “substantially,” “approximately,” and “about” may be used interchangeably. These terms generally refer to an approximation corresponding to less than or equal to 10% variation of the stated quantity, such as plus or minus 10%, plus or minus 5%, or plus or minus 3%.
In one aspect, methods of sensing are described herein. In some embodiments, such a method comprises providing a sensor. The sensor comprises a first layer having at least one single nanohole structure or at least one dual nanohole structure, and a second layer having at least one nanopore. The single nanohole structure comprises only one nanohole. The dual nanohole structure comprises a first nanohole and a second nanohole connected by a gap. Additionally, the one nanohole (in the case of the single nanohole structure) or the gap (in the case of a dual nanohole structure) of the first layer is aligned with the nanopore of the second layer in a direction corresponding to a translocation direction across the first and second layers.
Sensors described herein having a dual nanohole structure can have any construction, structure, or property of a dual nanohole sensor described hereinbelow or described in United States Patent Application Publication No. 2020/0393456A1 by Alexandrakis et al. and directed to “Nanosensors and Methods of Making and Using Nanosensors,” the entirety of which patent application publication is hereby incorporated by reference (hereinafter referred to as “US 2020/0393456A1”). For example, in some cases, the sensor has the structure of
Similarly, a sensor described herein having a single nanohole structure (rather than a dual nanohole structure) can likewise have any construction, structure, or property of a sensor described in US 2020/0393456A1, with the exception that the first layer comprises a single nanohole structure in place of the dual nanohole structure. Additionally, the nanohole of the single nanohole structure of such a sensor described herein can have the size and/or shape of any nanohole (e.g., a first nanohole or a second nanhole) described in US 2020/0393456A1. Other sizes and shapes are also possible. It is further to be understood that a single nanohole sensor described herein can be formed using the general methods described in US 2020/0393456A1 (modified as needed for the formation of a single nanohole structure rather than a dual nanohole structure), or using other methods known to one of ordinary skill in the art.
Turning in more detail to features of sensors described herein, a sensor comprises, in some embodiments, a chip or a wafer. The chip or wafer, in some cases, is defined by an xy-plane comprising at least a first layer and a second layer. The first layer, in some embodiments, is essentially parallel to the second layer, which is in contrast to a perpendicular z-direction. The z-direction is a translocation direction that is perpendicular and extending through the xy-plane of the chip or wafer. In some embodiments, the translocation direction goes through the first layer and the second layer of the xy-plane. The translocation direction, in some embodiments, is unidirectional, wherein the first layer is penetrated before the second layer. For example, a translocation direction can correspond to a movement through the chip or wafer from a cis chamber to a trans chamber, wherein the cis chamber is in communication with and, in some instances, partially defined by the first layer, and the trans chamber is in communication with the second layer. In some embodiments, a chip or wafer of a sensor described herein can have a substantially rectangular or square shape. In some cases, a chip or wafer can have a length and/or width of about 5-50 mm, 5-40 mm, 5-30 mm, 10-30 mm, or 10-20 mm, or about 15 mm.
In some embodiments, the first layer is positioned above or superior to the second layer. In some instances, the first and second layer are immediately adjacent layers. In some such embodiments, the first and second layers may be joined or adhered to one another via direct layer/wafer bonding. Alternatively, in other cases, the first layer and the second layer are not immediately adjacent layers but are instead spaced apart by or adhered together with one or more adhesion layers. In some such embodiments, the adhesion layer bonds to both the first layer and to the second layer with a greater bonding strength than the first layer and the second layer would bond to one another in the absence of the adhesion layer. Any adhesion layer not inconsistent with the objectives of the present disclosure may be used in a sensor described herein. For example, in some cases, an adhesion layer is formed from a metal (e.g., an elemental metal or a mixture or alloy of different metals), which may be particularly useful for adhering or bonding a metal first layer to an electrically insulating material second layer described herein, or for adhering or bonding a gold first layer to a silicon nitride second layer as described herein. In some embodiments, a metal adhesion layer can comprise titanium (e.g., elemental titanium metal) or chromium (e.g., elemental chromium metal). Other materials may also be used to form an adhesion layer of a sensor described herein. Further, an adhesion layer can have any thickness or average thickness not inconsistent with the objectives of the present disclosure. For example, in some cases, an adhesion layer has a thickness of up to 50 nm, up to 20 nm, up to 10 nm, or up to 5 nm. In some instances, an adhesion layer described herein has a thickness of about 0.5-20 nm, 0.5-15 nm, 0.5-10 nm, 1-20 nm, 1-15 nm, 1-10 nm, or 1-5 nm. In some embodiments, an adhesion layer can have a thickness of about 0.5-5 nm, 0.5-4 nm, 0.5-3 nm, 0.5-2 nm, or 0.1-1 nm.
Additionally, in some embodiments, the first and/or second layer of a sensor described herein is formed from an inorganic material, such as a metal (which may be an elemental metal or mixture or alloy of metals) or an electrically insulating material, as described further herein below.
Moreover, the first layer, in some cases, functions as an optically sensing layer. In some embodiments, as described above, the first layer is formed from a metal. Any metal not inconsistent with the objectives of the present disclosure may be used. For example, the metal can be an elemental metal or a mixture or alloy of metals. In one embodiment, the first layer is formed from gold. In some instances, a first layer described herein is formed from a different metal. The first layer material is not necessarily particularly limited. In some cases, a specific material is chosen because of its electrical conductivity properties, its chemical inertness in biological systems, and/or its compatibility with device fabrication methods described herein. In another aspect, the first layer has an average thickness of up to 500 nm in the translocation direction. In some embodiments, for example, the first layer has an average thickness of about 5-110 nm, 10-120 nm, 20-130 nm, 30-140 nm, 60-200 nm, 70-300 nm, 80-400 nm, 90-500 nm, or about 50-150 nm in the translocation direction.
Additionally, the first layer, in some embodiments, comprises at least one dual or double nanohole structure, wherein the dual or double nanohole structure comprises a first nanohole and a second nanohole. The first nanohole, in one embodiment, is essentially the same as the second nanohole. In another aspect, the nanoholes of the first layer each have an average diameter in the direction perpendicular to the translocation direction of about 80-150 nm. In other aspects, the nanoholes of the first layer each have an average diameter in the direction perpendicular to the translocation direction of about 100-150 nm, 80-100 nm, 80-120 nm, 90-120 nm, 90-130, or 100-120 nm.
In some embodiments, the nanoholes can have a center-to-center separation distance of about 150 nm or less, and in some cases, the nanoholes can overlap. For example, the nanoholes, in some embodiments, can each have a perimeter drawn by a theoretical line, thereby creating two imaginary circle-like shapes. In some instances, the theoretical lines defining the perimeter shape of each nanohole intersect in one or two locations. When the lines touch or intersect, it is understood that the nanoholes touch or overlap, respectively. In other instances, the theoretical lines defining the perimeter of each nanohole may not touch or intersect. When the lines do not touch or intersect, it is understood that the nanoholes do not touch or overlap. In some cases, the nanoholes can have a center-to-center separation distance of about 50-150 nm, 75-150 nm, 80-140 nm, 80-130 nm, or 100-120 nm.
In another embodiment, the nanoholes can have sloped or tapered interior walls along the translocation direction. The sloped interior walls, in some instances, can have a grade of about 10-30%. For example, the nanoholes can have an interior wall with a downward slope in the translocation direction such that each nanohole is shaped like an inverted cone or a funnel. In some cases, the sloped or tapered walls can have a grade of about 10-20%, 15-20%, or 15-30%.
Moreover, it is to be understood that a “nanohole” described herein can have any shape not inconsistent with the objectives of the present disclosure, including any cross-sectional shape in the xy-plane (perpendicular to the translocation direction). In some embodiments, one or both nanoholes are generally round, circular, ovoid, or ellipsoidal (ignoring any “gap” between the nanoholes, as described above). In other instances, one or both nanoholes have a triangular or other polygonal cross-sectional shape in the xy-plane. The precise shape of a nanohole described herein is not particularly limited. It is further to be understood that the size and/or center-to-center separation of a pair of nanoholes described herein can be selected based on the cross-sectional shape of the nanohole and/or based on the biomolecule analyte to be optically trapped in the dual nanohole structure. In one exemplary embodiment, for instance, two equilateral triangular nanoholes may be used having side lengths of 50-150 nm, wherein vertices of the triangular nanoholes are joined or separated by the gap of the dual nanohole structure.
For clarity, it is to be understood that a sensor described herein can have a single nanohole structure rather than, or in place of, a dual nanohole structure. In such instances, the single nanohole can have the same size, shape, and other physical characteristics as one of the nanoholes of the dual nanohole structure described hereinabove. For example, in some cases, the single nanohole of the first layer of a sensor described herein has a diameter or average size in the direction(s) perpendicular to the translocation direction of about 80-150 nm. In other cases, the single nanohole of the first layer has a diameter or an average size in the direction(s) perpendicular to the translocation direction of about 100-150 nm, 80-100 nm, 80-120 nm, 90-120 nm, 90-130, or 100-120 nm. Additionally, as an another example, the single nanohole of a first layer described herein, in some instances, can have sloped or tapered interior walls along the translocation direction. The sloped interior walls, in some instances, can have a grade of about 10-30%. For example, the single nanohole can have an interior wall with a downward slope in the translocation direction such that the single nanohole is shaped like an inverted cone or a funnel. In some cases, the sloped or tapered walls can have a grade of about 10-20%, 15-20%, or 15-30%.
In some embodiments, the first layer can be non-continuous. As shown in
In some embodiments, a non-continuous first layer comprises a perimeter circumscribing the dual nanohole of the first layer, wherein the perimeter is an edge of the first layer. For example, the perimeter can define an island of the first layer that is separated from other areas of the first layer. In some embodiments, a geometric perimeter, such as circular, rectangular, or square perimeter, can be defined around the dual nanohole of the first layer. In some cases, one or more areas of the first layer outside the perimeter can be etched away or removed, such that the second layer can be visible through the first layer in the areas where it is etched away or removed, as described, e.g., in US 2020/0393456A1. In some cases, a first layer perimeter surrounding a dual nanohole structure can have a circumference measuring 1 μm-50 mm, 1 μm-40 mm, 1 μm-30 mm, 1 μm-20 mm, 1 μm-10 mm, or 1 μm-1 mm. In some embodiments, an island 803 in the first layer can have an area of about 1 μm2-100 mm2. In some cases, a hole and/or island 803, as described herein that is not the one or more dual nanoholes in the first layer can have an area of about 100 μm2-100 mm2, 100 μm2-80 mm2, 100 μm2-70 mm2. 100 μm2-60 mm2, 100 μm2-50 mm2, or 100 μm2-40 mm2, 100 μm2-30 mm2, 100 μm2-20 mm2, or 100 μm2-10 mm2, 100 μm2-10 mm2, 100 μm2-5 mm2, 100 μm2-1 mm2, or 100 μm2-0.5 mm2.
Not intending to be bound by theory, it is believed a non-continuous first layer having islands and/or holes, as described above, can reduce metal layer shielding of the externally applied electric field across the sensor and can increase the electrical conductivity of ionic solution added to the sensor. Consequently, it is believed that a non-continuous first layer can increase the throughput of analytes present in the ionic solution. Additionally, in some cases, such holes and/or selectively deposited areas of the first layer can act as alignment markers.
The second layer, in some embodiments, enables electrical sensing and thus functions as an electrical sensing layer. The second layer is formed from an electrically insulating material in some cases. For example, in one embodiment, the second layer is formed from a silicon nitride. Any silicon nitride not inconsistent with objectives of the present disclosure can be used. In some cases, silicon nitride comprises SixNy. In some cases, silicon nitride comprises Si3N4. In some instances, a second layer described herein is formed from a ceramic material. As described above, such a ceramic material can be electrically insulating. In some embodiments, a second layer described herein is formed from a metal oxide such as a transition metal oxide. In some cases, a second layer described herein is formed from a silicon oxide such as SiO2. Other electrically insulating materials may also be used. The electrically insulating material is not necessarily particularly limited. In some cases, a specific material is chosen because of its electrical conductivity properties, its chemical inertness in biological systems, and/or its compatibility with device fabrication methods described herein. Additionally, the second layer, in some embodiments, has an average thickness of up to 100 nm, or up to 70 nm in the translocation direction. For example, the second layer can have an average thickness of about 5-100 nm, 5-70 nm, 10-70 nm, 20-80 nm, 20-70 nm, 30-120 nm, 30-90 nm, 30-70 nm, 40-100 nm, 40-70 nm, or 50-100 nm in the translocation direction.
In addition, the second layer, in some embodiments, comprises at least one nanopore. The nanopore, in one case, has a diameter of at least 5 nm. In other cases, the nanopore has a diameter of about 2-20 nm, 5-25 nm, 15-35 nm, 20-40 nm, or 10-30 nm. In some embodiments, the nanopore is a solid-state nanopore. As understood by one of ordinary skill in the art, a solid-state nanopore is not a biological nanopore, as a solid-state nanopore comprises structural and functional differences that are distinguishable from a biological nanopore.
In some embodiments, the first nanohole and the second nanohole are connected by a gap. The gap as described herein is defined by a continuous hole or opening in the first layer connecting the first nanohole and the second nanohole. The gap, in some instances, is measurable in the x- and y-directions of the xy-plane of the chip. In some instances, the gap defines a line.
In some embodiments, the length and the width of the gap are measured in the xy-plane. In one embodiment, the width and/or length of the gap is defined by a distance between the points of intersection of the theoretical lines defining the perimeter of each nanohole. In some cases, the gap has a width and/or length of about 10-50 nm. In some embodiments, the gap has a width and/or length of about 20-50 nm, 20-40 nm, 30-50 nm, or 20-30 nm.
In some cases, the width and/or length of the gap is defined by the diameter of the nanopore. For example, in some embodiments, the width and/or length of the gap is within 10% of the diameter of the nanopore. Additionally, the gap, in some embodiments, is continuous with the nanopore in the translocation direction. The gap, in other embodiments, has a measurable width and/or length greater than the diameter of nanopore. In other embodiments, the width and/or length of the gap is less than 200% the diameter of the nanopore. In some embodiments the width and/or length of the gap is between 100% and 200% the diameter of the nanopore.
In some embodiments, the center of the gap, determined by its center point in the xy-plane, and the center of the nanopore, also determined by its center point the xy-plane, are aligned in the translocation direction. For example, in some embodiments, the center of the gap of the first layer is aligned with the center of the nanopore of the second layer in a direction corresponding to a translocation direction across the first and second layers, and the centers are spatially separated in the x- or y-direction of the xy-plane by less than 10 nm, or less than 5 nm.
For clarity, it is to be understood that a sensor described herein can have a single nanohole structure rather than, or in place of, a dual nanohole structure. In such instances, the center of the single nanohole, determined by its center point in the xy-plane, and the center of the nanopore, also determined by its center point the xy-plane, are aligned in the translocation direction. For example, in some embodiments, the center of the single nanohole of the first layer is aligned with the center of the nanopore of the second layer in a direction corresponding to a translocation direction across the first and second layers, and the centers are spatially separated in the x- or y-direction of the xy-plane by less than 10 nm, or less than 5 nm.
The sensor chip or wafer described herein, in some embodiments, further comprises an optional third layer. The presence of such a layer is preferred in some embodiments. The third layer, in some embodiments is positioned inferior or adjacent to the second layer, such that the second layer is positioned between the first and third layers. The third layer, in some embodiments, can act as an electrically insulating layer that is secondary or supplemental to the second layer, which is also an insulating layer. Thus, the third layer, in some embodiments, can be formed from an electrically insulating material. For example, in some cases, the third layer can comprise or be formed from silicon dioxide (SiO2). In some instances, a third layer described herein is formed from an electrically insulating material described hereinabove, such as a ceramic material or transition metal oxide. Other electrically insulating materials may also be used. The electrically insulating material of the third layer is not particularly limited. In some cases, a specific material is chosen because of its electrical conductivity properties, its chemical inertness in biological systems, and/or its compatibility with device fabrication methods described herein. Not intending to be bound by theory, it is believed that, in some instances, the third layer, which can act as an insulating layer, can contain leakage or prevent the passage of current through any additional layers present beyond the second layer of the chip (in the “downward” direction in
In some embodiments, the sensor chip or wafer described herein can comprise a fourth layer positioned inferior or adjacent to the third layer, such that the third layer is positioned between the second and fourth layers. Alternatively, the fourth layer can be positioned adjacent the second layer in the absence of a third layer. The fourth layer, in some embodiments, comprises or is formed from silicon. For example, in some cases, the fourth layer can be formed from pure silicon. Other semiconducting materials may also be used. In some embodiments, the fourth layer can act as a semiconducting layer. The fourth layer, in some instances, can have an average thickness of about 1-1000 μm, 10-1000 μm, 50-1000 μm, or 50-500.
In some embodiments, the sensor chip or wafer described herein can comprise one or more layers in addition to the third and fourth layers, such that the fourth layer is positioned between the second or third layer and the one or more additional layers. For example, in some instances, one or more additional layers comprising silicon, including pure silicon, silicon dioxide, and/or silicon nitride can be used. Such additional layers can have an average thickness of about 1-1000 μm, 100-1000 μm, 200-800 μm, 300-600 μm, or about 500 μm.
In some cases, the third layer, fourth layer, and/or additional layers can define a window or an opening “beneath” the first and second layers (e.g., “downward” in
One non-limiting example embodiment of a sensor or chip described herein is illustrated schematically in
Another non-limiting example embodiment of a sensor or chip described herein is illustrated schematically in
Moreover, it is to be understood that the example embodiments of
It is also possible, in some embodiments, for the sensor to include or be coupled to a layer, device, or structure for concentrating a test sample, prior to analyzing the test sample as described herein. For example, in some cases, an isotachophoretic layer, device, or structure is disposed on top of the sensor. Such a layer, device, or structure can be used to concentrate an analyte within a test sample using isotachophoresis (ITP) prior to contacting the test sample with other layers or components of the sensor (e.g., the first layer of the sensor as described above).
The ITP device (2200) further comprises a first input port (2210), a second input port (2212), a T-junction (2214), a vertical feed channel (2216), and an eluate collection channel (2218). As described further herein, the vertical feed channel (2216) and the eluate collection channel (2218) intersect at the intersection point (2208).
As illustrated in
As illustrated in
Not intending to be bound by theory, it is believed that a large reduction (e.g., 100 to 10000 times reduction, such as the 1000 times reduction of the embodiment of
Since analyte migration is proportional to current density, the voltage is lowered once the analyte/test sample enters zones 1 and 2 for proper resolution. To further improve separation efficiency, one or more electro-osmotic flow suppressors (e.g. poly(vinylpyrrolidone) or poly(ethylene glycol) species of different molecular weights) can be added to both the leading and terminal electrolyte of the ITP process. Additionally, if desired, non-detectable spacer ions can be used, as understood by one of ordinary skill in the art.
As illustrated in
The test sample and ITP electrolytes/buffers are driven across or through the flow microchannel using electrical potential with valves 1 and 2 open, and valve 3 closed from a power supply (e.g., XHR 600-1, Xantrex Technology Inc., Vancouver, Canada, not shown). The two buffers/electrolytes (one preceding and one lagging relative to the analyte plug) move with different speeds under the external voltage difference applied across the entire device. As a result, the preceding electrolyte, lagging electrolyte, and analyte-containing test sample mix, and as they travel through the sinuous ITP separation channel (2202), chemical species within the mixed fluid separate according to ITP principles (e.g., based on mass due to electrophoretic mobility differences). At steady state, molecular species, such as proteins, within the test sample will separate out into discrete zones because of the difference in their electrophoretic mobilities. The targeted analyte band should occur or be found at an intersection point (2220) of the ITP separation channel (2202) and the vertical feed channel (2216) to the nanopore sensor. After analyte fractionation occurs in the ITP separation channel (2202), valves 1 and 2 are closed, valve 3 is opened, and the analyte-containing fraction from the ITP separation channel is fluidically injected into the vertical feed channel (2216), which runs to the SANE sensor (e.g., the sensor (4000) of
In some implementations, an ITP microchannel structure or device described above is disposed over or in contact with the first layer of the sensor. Moreover, the ITP structure or device can be bonded or adhered to one or more other layers of the sensor. In some cases, for example, the ITP structure or device forms a unitary chip with the first layer and the second layer of the sensor.
Turning again to specific steps of methods of sensing described herein, a method described herein further comprises providing a test sample comprising an analyte and contacting the test sample with the first layer of the sensor. Particular test samples and analytes are described further hereinbelow, including the specific Examples. More generally, the test sample can be provided and contacted with the first layer of the sensor in any manner not inconsistent with the technical objectives of the present disclosure. In some embodiments, for example, the test sample is provided in a chamber positioned cis of a translocation direction of the sensor. For example, a cis chamber can be positioned adjacent and/or superior to a first layer of a chip of the sensor, as described above, such that placing or positioning the test sample in the cis chamber comprises contacting the test sample with the first layer of the sensor.
It is also possible, in some embodiments, to concentrate the test sample prior to contacting the test sample with the first layer of the sensor. In particular, in some cases, the concentration of the analyte within the test sample is increased. Such concentration of the test sample or of the analyte within the sample can be carried out in any manner not inconsistent with the objectives of the present disclosure. In some preferred embodiments, the test sample is concentrated using isotachophoresis (ITP). For example, in some implementations, the test sample is concentrated using an ITP microchannel structure, including an ITP microchannel structure described further hereinabove and hereinbelow.
Additionally, in some embodiments, the first layer of the sensor is an optically sensing layer. When the first layer of the sensor is the optically sensing layer, it is expected that the test sample is subjected or exposed to the optically sensing layer of the sensor before being subjected or exposed to other layers of the sensor.
Methods described herein also comprise irradiating the single nanohole structure or the dual nanohole structure of the sensor with a beam of electromagnetic radiation. Irradiating the nanohole structure (single or dual) can comprise irradiating with a laser beam or laser light (or other suitable electromagnetic radiation). In some cases, the laser beam can be polarized circularly and/or linearly prior to focusing on the nanohole structure. In some cases, linearly polarized light is preferred for impingement on the nanohole structure. Additionally, in some instances, the laser beam can be focused on the nanohole structure using one or more mirrors.
The wavelength of electromagnetic radiation used is not necessarily limited. In some embodiments, the laser beam or other electromagnetic radiation comprises visible light or has a wavelength (or average wavelength) centered in the visible region of the electromagnetic spectrum, such as between 450 nm and 750 nm, between 500 nm and 700 nm, or between 550 nm and 650 nm. In some cases, the laser beam or other electromagnetic radiation comprises infrared (IR) light or has a wavelength (or average wavelength) centered in the IR region of the electromagnetic spectrum. For example, in some instances, a laser beam described herein has a wavelength centered in the near-IR (NIR, 750 nm-1.4 μm), short-wavelength IR (SWIR, 1.4-3 μm), mid-wavelength IR (MWIR, 3-8 μm), or long-wavelength IR (LWIR, 8-15 μm). Moreover, in some embodiments, a laser beam or other electromagnetic radiation described herein has an emission profile/wavelength distribution overlapping with a wavelength at which water and/or biological tissue has an absorption minimum, such as a wavelength between about 700 nm and about 800 nm or between about 1.25 μm and about 1.35 μm.
In another aspect, the method of sensing comprises optically trapping the analyte in (1) the single nanohole structure or (2) in the dual nanohole structure and/or (3) in the gap of the first layer of the sensor. Optical trapping, in some embodiments, is a result of irradiating the nanohole structure of the first layer of the sensor with the beam of electromagnetic radiation. Thus, in methods of sensing described herein, a step of optically trapping the analyte is performed prior to and/or during a step of irradiating the nanohole structure. In some instances, the optical trapping lasts for at least 1 microsecond and less than 100 seconds. In some embodiments, optical trapping lasts for about 1 millisecond-60 sec, 1 millisecond-30 sec, 1 millisecond-10 sec, 1 millisecond-5 sec, 1 millisecond-5 sec, or 10 millisecond-5 sec.
In some embodiments, a step of optically trapping the analyte can comprise one or more optical trapping events. In some cases, an optical trapping event represents the optical trapping of a single analyte species, such as a single molecule. In other cases, an optical trapping event represents the optical trapping of more than one analyte species, such as more than one analyte molecule. For example, in some cases, an optical trapping event can include optical trapping of a first analyte molecule followed by optical trapping of a second analyte molecule, wherein the second analyte molecule is optically trapped with the first analyte molecule. Moreover, as one example for a particular type of analyte, it should be understood that optically trapping a first non-complexed biomolecule and a second non-complexed biomolecule is not the same as optically trapping a complexed biomolecule.
In some instances, optically trapping the analyte results in a measurable surface plasmon resonance. When a first analyte species (e.g., a first molecule) is optically trapped, a first surface plasmon resonance can be measured. When a second analyte species (e.g., a second molecule) is optically trapped with a first biomolecule, a second surface plasmon resonance can be measured. In some instances, a first plasmon resonance measured from the first optically trapped analyte species/molecule can be subtracted from the second plasmon resonance measurement to obtain information related to the second analyte species/molecule. In some embodiments, the surface plasmon resonance provides information about the mass of an optically trapped analyte such as an optically trapped biomolecule. Therefore, measuring the surface plasmon resonance, in some embodiments, comprises measuring the mass of the optically trapped analyte (e.g., biomolecule).
Thus, methods of sensing described herein, in some implementations, further comprise measuring a surface plasmon resonance of the single nanohole structure or the dual nanohole structure after optically trapping the analyte in the single nanohole structure or in the dual nanohole structure and/or in the gap of the first layer of the sensor. Additionally, in some such embodiments, measuring the surface plasmon resonance further comprises determining the mass of the optically trapped analyte.
Furthermore, in some embodiments, optically trapping comprises slowing or delaying the translocation of the analyte. Such a slowing or delaying can provide greater resolution to the sensing mechanisms of sensors described herein.
Moreover, in some cases, a method described herein further comprises applying a first electric field across the nanopore to draw one or more of the analytes into the nanopore, wherein the first electric field comprises a direct current (DC) electric field. The electric field can be applied across the nanopore in any manner not inconsistent with the technical objectives of the present disclosure. In some cases, the DC electric field is provided by placing patch clamp electrodes in the cis and trans chambers of the sensor. Thus, in some cases, applying an electric field across the nanopore can comprise applying an electric field from the cis chamber to the trans chamber of the sensor. In some embodiments, applying an electric field comprises applying a 10-1000 mV bias. In some cases, applying an electric field comprises applying a 10-500 mV, 50-50 0 mV, 100-500 mV, or about 250 mV bias.
Additionally, in some embodiments, the DC electric field is temporarily reversed. For example, in some cases, the DC electric field is applied from the trans chamber to the cis chamber. Temporary reversal of the electric field is sometimes performed to prevent clogging, or a build-up of biomolecules, at the nanopore and/or the nanohole structure. It is to be understood, however, that such reversal of the DC electric field is not the same as applying an alternating current (AC), pulsed, or modulated field, including as described hereinbelow. Instead, temporary reversal of the DC electric field is a separate step for reducing or preventing clogging of the nanopore.
Methods described herein also comprise applying a second electric field across the nanopore after applying the first electric field, wherein the second electric field comprises a pulsed, modulated, or alternating current (AC) electric field. This second field can be applied in any manner not inconsistent with the technical objectives of the present disclosure. In some embodiments, for example, the second electric field is provided by placing patch clamp electrodes in the cis and trans chambers of the sensor (and the same electrodes can be used for providing the second field as well as the first DC field; different electrodes may also be used). Additionally, as described above for the first (DC) electric field, the second electric field can be applied from the cis chamber to the trans chamber of the sensor. In some embodiments, applying the second electric field comprises applying a 10-1000 mV bias, a 10-500 mV bias, a 50-500 mV bias, a 100-500 mV bias, or about a 250 mV bias.
Additionally, the second field can have an AC, pulse, or modulation frequency of up to 1 GHz. For example, in some embodiments, an AC field of frequencies up to 1 GHz can be applied to Ag/AgCl electrodes by an external function generator and detected by Axopatch electronics. It is further to be understood that an electric field that is “pulsed” or “modulated” without necessarily being an “AC” field can be “pulsed” or “modulated” by virtue of the existence of one or more of the following: cycles or periods of “on” and “off” times; cycles or periods of “high” intensity and “low” intensity; or cycles or periods of “high” frequency and “low” frequency.” In such cases, “high” and “low” are relative terms, where the terms are relative to one another (i.e., relatively high compared to relatively low), and the differences between “high” and “low” frequency or intensity is at least 20% (based on the larger number as the denominator). It is further to be understood that “on” times refer to time periods in which the AC, pulsed, or modulated electric field is “on” or “applied,” and “off” times refer to time periods in which the AC, pulsed, or modulated electric field is “off” or “not applied.”
Moreover, it is to be understood that the second field described herein can be and preferably is applied ‘over’ the first field (the DC electric field). More particularly, in some preferred embodiments, the first field (the DC electric field) is applied continuously or substantially continuously throughout a sensing method described here. The second field (the AC, pulsed, or modulated field) is applied over or simultaneously with a ‘baseline’ provided by the first field (the DC field). In other words, in some preferred embodiments described herein, even when the external second field (the AC, pulsed, or modulated field) is not applied, a DC voltage is always applied throughout the entire method, and this DC voltage (typically up to +/−200 mV) provides the baseline on which the modulation (e.g., the AC modulation) ‘rides.’ Thus, in some preferred embodiments, the first field and the second field are both applied contemporaneously for a period of time or, in other words, the second field is applied after the first field is applied, but the first field continues to be applied even during application of the second field.
Further, it should also be noted that the AC, pulsed, or modulated waveforms of the second field may or may not be sinusoidal. In some instances, these waveforms are sinusoidal. Alternatively, in other cases, these waveforms can have any other bipolar form, such as provided by square waves or triangular waves.
As described herein, applying an electric field across the nanopore, in some embodiments, results in one or more translocation events of an analyte species (such as an analyte biomolecule). A translocation event, as described herein, comprises the entry and exit of an individual analyte species (e.g., a biomolecule analyte) through the nanopore. Moreover, in some instances, applying an electric field (first and/or second) generates a measurable current across the nanopore.
In addition, in some implementations, a method described herein further comprises measuring one or more analyte properties or other properties potentially associated with optical trapping or translocation of the analyte through the nanopore of the sensor. For example, in some cases, a method described herein further comprises measuring a change in current and/or phase across the nanopore during application of the second electric field while the analyte is optically trapped and/or during one or more translocation events of the analyte through the nanopore. As described further herein, such a measurement can, in some embodiments, provide information regarding the analyte that may not otherwise be known or detected. For example, in some cases, measuring change in current and/or phase further comprises determining a charge of a translocating analyte. Moreover, in some instances, measuring change in current and/or phase further comprises determining a dielectric constant of a translocating analyte.
In addition to the measurements of the foregoing paragraph, a method described herein can also (or alternatively) comprise measuring at least one kinetic parameter of the analyte within the nanopore after removing or turning off the second electric field. Further, in some embodiments, at least one kinetic parameter is measured while the analyte decelerates or comes to a stop while optically trapped.
Various kinetic parameters can be measured using a method described herein. For example, in some cases, the at least one kinetic parameter comprises one or more of the following well known parameters in the field of biochemistry: equilibrium dissociation constant (Kd), binding on-rate (kon), binding off-rate (koff), and bound fraction (i.e., the fraction of analyte that is in a bound or complexed state rather than an unbound or uncomplexed state). In other instances, the at least one kinetic parameter comprises analyte size (volume), analyte charge (effective charge on the outer surface of the analyte), or analyte conformation. It is to be understood that these kinetic parameters can apply to analytes that are single molecules, molecular complexes such as protein complexes, or nanoparticles used for drug and gene delivery. Moreover, the foregoing kinetic parameters can be measured in any manner not inconsistent with the technical objectives of the present disclosure. For instance, in some embodiments, the foregoing kinetic parameters can be measured as follows:
Equilibrium dissocation constant (Kd) (units of molar (M)) is the inverse of Ka, the association constant, in which Ka=kon/koff.
Binding on-rate (kon) is the on-rate constant measured in units of M−1 s−1.
Binding off-rate (koff) is the off-rate constant measured in units of s−1, indicative of the rate of analyte unbinding events per second.
Bound fraction is the fraction of bound protein events detected over all analyte events detected. It is equal to the number of events detected by the sensor as bound analyte divided by the total number of events detected by the sensor from bound and unbound analyte.
Analyte size (volume) is the volume of analyte represented by a sphere of equivalent volume and is measured in nm3.
Analyte charge (effective charge on the outer surface of the analyte) is the net charge surrounding the surface of the analyte, and it is expressed as units of single electron charge (e) or in Coulomb.
Analyte conformation is assessed by detecting shape changes of the analyte while trapped. If the analyte is not rigid, e.g. a protein, it can change shape dynamically while inside the optical trap of the sensor. The changes in shape (protein conformation) result in dynamic changes of both optical and electrical signals detected by the sensor. Different protein shapes scatter light, causing optical signal variability, and block the current conducting through the nanopore, inducing electrical signal variability, by different amounts.
As described herein, a method according to the present disclosure can use various parameters or measurements to detect and/or characterize an analyte of interest. For example, in some cases, a method described herein uses one or more of the following parameters or measurements to detect and/or characterize an analyte: optical data, current (e.g., across a nanopore), command voltage, conductance (e.g., ratio of current to command voltage), phase change, post-decay drive fits (e.g., intercept of regression, magnitude of oscillation, decay frequency, slope of linear drift component, decay phase, and/or decay coefficient), optical step change, trapping event counts, trapping current, and nanopore translocation current spikes.
In addition, methods described herein can be used for sensing a variety of analytes. For example, in some cases, the analyte comprises complexed and/or non-complexed biomolecules. Complexed and/or non-complexed biomolecules can include, but are not necessarily limited to, exosomes, endosomes, micelles, nucleotides, proteins, lipids, and/or carbohydrates. The biomolecules can be, in some instances, complexed with one or more secondary biomolecules. Exemplary secondary biomolecules may include, but are not limited to small molecules, nucleotides, oligonucleotides, aptamers, proteins, antibodies, lipids, and/or carbohydrates. The secondary biomolecules, in some embodiments, may be of similar origin as the complexed and/or non-complexed biomolecules. In another embodiment, the secondary biomolecules may be of different origin than the complexed and/or non-complexed biomolecules. For example, a secondary biomolecule may be derived from different species or other foreign organism. In other instances, the biomolecules can be complexed with non-biological molecules. Non-biological molecules may include, but are not limited to any kind of pharmaceutical, such as an antibody, a recombinant protein, a small molecule, or other synthetic product.
Moreover, in some embodiments, the test sample is a biological sample obtained from an animal or human subject, such as a human patient or animal patient in need of diagnosis (e.g., through detection or characterization of an analyte present in a sample taken from the human or animal patient). Some conditions, illnesses, or diseases may especially benefit from the sensing provided by methods described herein. For example, in some instances, cancer diagnosis and/or treatment can be improved using the sensors and methods of the present disclosure, as described further hereinbelow in the specific Examples. Thus, in some cases, an analyte described herein comprises a Peptide-presenting Major Histocompatibility Complex Class-I (pMHC) or pMHC component. In some instances, the analyte comprises a HLA-A2 pHMC or HLA-A2 pMHC component. In still other implementations, the analyte comprises a T-Cell Receptor-mimic (TCRm) antibody. Moreover, in some preferred embodiments, the analyte comprises a TCRm antibody against a HLA-A2 pHMC or against a HLA-A2 pHMC component.
In other instances, the analyte comprises a nanoparticle such as an inorganic nanoparticle. The inorganic nanoparticle can be a metal nanoparticle, such as a gold (Au), silver (Ag), platinum (Pt), or other nanoparticle. In other instances, the inorganic nanoparticle can be a ceramic or glass nanoparticle, such as a nanoparticle formed from silica (SiO2) or titania (TiO2).
Moreover, in some embodiments, a test sample described herein is provided in an ionic solution, such as a salt solution. Any ionic or salt solution not inconsistent with the objectives of the disclosure can be used, including NaCl, KCl, or CaCl2) solution.
In some embodiments, methods of sensing described herein provide for detecting or sensing analytes (such as biomolecules) at a milli-(10-3), micro-(10-6), nano-(10-9), pico-(10-12), femto-(10-15), or atto-(10-18) molar concentration of the analytes (e.g., biomolecules).
Some features and characteristics of sensors and methods of sensing according to the present disclosure are described in further detail in the specific Examples below.
Nanopores can be used to discriminate among analytes through the analysis of changes in conduction current profiles during translocation. Nanopore measurements can enable the discrimination between single molecule species in solution and can help achieve low-cost and label-free DNA sequencing. Other additional possible applications are expanding rapidly. However, the translocation times of analytes through a traditional nanopore are extremely fast, which limits the fidelity of electrical data that can be collected. Through the use of optical trapping enabled by the self-induced back-action (SIBA) effect, nanopores can be enhanced not only by slowing down the translocation of analytes but also by introducing new dimensionality to the collected data through the collection of optical data simultaneously with electrical data.
Herein, the present inventors describe a SIBA actuated nanopore electrophoresis (SANE) sensor, effectively a nanopore with plasmonic optical trapping, that has been shown to be capable of trapping individual nanoparticles, proteins and protein complexes and through the use of bimodal optical and electrical data, discriminating between analyte species. The present Example combines driving the SANE sensor with an AC voltage (or other modulated or pulsed electric field), which was previously inaccessible, mostly because of fast translocation times that are typically in the hundreds of us, which would necessitate a MHz driving frequency that exceeds the available frequency limit. The optical trap of the SANE sensor provides a trapping duration in the seconds range, which allows for frequencies as low as 1 Hz. An upper bound of possible AC measurement frequencies can be set by the amplifier hardwired filters (100 kHz) and the data acquisition sample rate (500 kHz).
Not intending to be bound by theory here (or elsewhere throughout this application), it is believed that by introducing modulation, the mobility of ions in the fluid around analytes depends on the analyte surface charge and the material properties in the case of nanoparticles, thus possibly enhancing the capacity of the sensor to distinguish between analyte species. This disclosure presents results of a new sensing method able to discriminate between 20-nm SiO2 and 20-nm Au nanoparticles using electrical measurements. By applying a DC command voltage with a superimposed AC frequency sweep, while keeping the nanopores optically trapped in the vicinity of the nanopore's entrance, SiO2 and Au nanoparticles were found have distinctly different electrical responses. This disclosure demonstrates the feasibility of performing these AC (or other pulsed or modulated) measurements with a plasmonic nanopore.
The setup used for the method, including a laser diode, optics to polarize the laser beam, the sensor setup, the AC- and DC-generating devices, and the data acquisition instruments, are herein provided in a graphical schematic (
For the AC measurements, the optical data are also sent to an Arduino Uno (15) (Adafruit Industries, New York, NY) programmed to perform edge detection, which is to be understood as that it looks for an optical step change indicative of a trapping event that is accompanied by a simultaneous current spike in the electrical data stream. The Arduino then sends a trigger to an Agilent 33250A Waveform/Function Generator (16) (Agilent Technologies, Santa Clara, CA) that generates a user-defined pulse at a selected frequency and amplitude. The function generator then sets the command voltage for the Axopatch 200B, which is operated in voltage clamp mode. The ratio of the measured current amplitude to command voltage, i.e. conductance, is then computed as a characteristic parameter of the nanoparticle response to the AC modulation. This is referred to herein as an AC burst event. Further, the fast-Fourier transform (FFT) analysis of the recorded AC signal responses enables the calculation of the phase change while the particle is driven by the AC burst. Additional data types are obtained from fitting post-drive decay data once the driving burst has stopped to an empirically-derived formula incorporating a damped oscillation term, as described below. The DC voltage is consistently on and kept at 100 mV (−ve cis to +ve cis).
A baseline AC response is established using a model cell reference block provided by the Axopatch 200B manufacturer for calibrating the system. A model cell bath that has an equivalent circuit of a 10 MΩ resistor in series with a 4 pF capacitor is used for impedance matching during these baseline measurements with the Axopatch 200B. A baseline is taken with the sensor for the AC measurements using 40 μL of 0.3 M KCl solution at 7.4 pH. Baseline measurements are performed at 110 mV command voltage with one of the following AC frequencies superimposed: 1, 2, 5, 10, 20, 50, 100, 1000, 2000, 5000, 10000, 20000, 50000 and 80000 Hz.
The amplitude of the waveform at each frequency is set to ensure there was a high signal-to-noise ratio but low enough to ensure the Axopatch 200B does not saturate while recording the current response. 1 Hz measurements are taken with a 10 Vp-p signal, and 1 kHz are generally collected at 50 mV p-p. The Axopatch 200B front-switched command voltage port is used to connect to the function generator. This port reduces all signals by a factor of 20. Each frequency is set to pulse 5 times with 10 cycles each to enable testing the reproducibility of the response. In addition, signal decays recorded at the end of each burst are analyzed to generate additional data types for the characterization of nanoparticles.
All experiments were performed with a single sensor to allow for easy comparison among different experiments. To generate the optical trap and nanopore parts of the sensor, Ne/He focused ion beam milling was performed. 20+4-nm SiO2 nanoparticles (MEL0010, NanoComposix, zeta potential=−40 mV) and 20±2-nm Au nanoparticles (C11-20-TM-DIH-50, Nanopartz, zeta potential=−15 mV) were assessed with this method. The SiO2 nanoparticles were used to assess the system's response to dielectric materials. The Au particles were used to show the response of the system to conductive materials, and their responses were expected to be different. The SiO2 nanoparticles were tested at the same frequencies as the baseline measurement for comparison with the empty trap response and model cell response. Post-decay analyses were also performed after driving the nanoparticles in the trap at a single frequency of 100 Hz. The Arduino was programmed to trigger an AC burst on both a positive and negative optical step change to ensure a trapping event AC burst was paired to a non-trapping burst event.
Once the data were collected, each pertinent AC burst event was noted for start and stop times and if it took place during a trapping event. The event parameters were imported into a MongoDB document database (MongoDB Inc, New York, NY) and then loaded into MATLAB (MathWorks, Natick, MA) to be processed first for a frequency response and then for a decay response. The axon binary file (.abf) generated by the pCLAMP software (Molecular Devices) was trimmed according to the event times, and a FFT was performed on the current response and command voltage of the .abf data, as depicted in
The data were processed for their post-drive decay profiles. The updated event parameters were loaded back into MATLAB in which the pCLAMP data was trimmed at the termination of the AC pulse by smoothing the command voltage and then taking the first derivative to determine where along the time-axis the signal returned to the DC baseline voltage. The event time was then extended by 3 ms. Using the “prepareCurveData( ) and “fit( ) functions in MATLAB with the “NonlinearLeastSquares” option, a fit of the decay profile was created using Equation 1:
The form of this equation was selected empirically as the sum of a damped oscillation, representing the particle's decaying oscillation, and a line with a negative slope, representing particle drift. The initial fitting parameters were set accordingly as follows: a1 and a2 were set to the minimum absolute value of the trimmed data segment, b/was set to the maximum absolute value of the trimmed data segment, c/was set empirically to 15000 for SiO2 nanoparticles and 10000 for Au nanoparticles and the empty trap, d1 was set to 0, and e1 was set empirically to 0.653. The R2 value of the fit was used to filter out poorly fitting data, with the threshold set at 0.9, and the remaining parameters were analyzed to see how the nanoparticle type and concentrations would affect the post-drive decay parameters.
A baseline measurement of the model cell was performed. This was then repeated with the SANE sensor loaded with 0.3 M KCl and an empty trap. The results of the baseline experiments are presented in
In order to avoid multi-particle interactions in the vicinity of the optical trap, ultra-low concentrations of SiO2 nanoparticles (1 fM) were used at the same frequencies used for the baseline and compared to the response of the empty trap corrected for model cell response. The largest change in phase was seen at 1 kHz (
In addition to doing a frequency scan, both SiO2 and Au nanoparticles were driven at 100 Hz to compared their relative phase shift (
Subsequently, the post-drive decay that was fitted to damped oscillation riding on a linear slope according to Eq. I was assessed. Any fit with an R2 value less than 0.9 was excluded from the analysis to mitigate the effect of noisier measurements on fit parameters. The 100 Hz pulse post-drive decay parameters of the fit to Eq. 1 are presented in
In this Example, a SANE sensor was used with an AC voltage to optically trap SiO2 and Au particles, and after an AC burst at 100 Hz, the post-drive decay data profiles were fit using a damped oscillation model. The function selected to fit the post-drive decay (Eq. 1) was selected empirically by the inspection of curves similar to the one shown in
The baseline analysis of
For the driven oscillation at 100 Hz, the highest conductance was seen with the Au nanoparticles, and the conductance increased with Au nanoparticle concentration. In contrast, the SiO2 nanoparticles showed a lower conductance than the Au nanoparticles but still higher than that of the empty sensor. The difference in conductance caused by Au and SiO2 nanoparticles can be explained by the fact that Au is a conductor that would oppose changes in the applied voltage bias, which would, in turn, cause ionic fluid motion around these particles. Additionally, SiO2 nanoparticles are dielectric, causing them to polarize in the presence of an AC field, which would augment ionic charge displacement relative to an empty sensor but less so than the conducting Au nanoparticle. Another factor to consider is that particles are likely to accumulate over the SANE sensor because they are driven there by the DC bias; however, they cannot easily tunnel through the optical trap. This results in a “traffic jam” of nanoparticles that may all feel the effects of the applied AC bias. Not intending to be bound by theory, it may be for this reason that the phase response for SiO2 nanoparticles in
Next, in respect to the SANE sensor-derived parameters from the post-drive signal decay profiles, the larger amplitude line intercept and the steeper linear slope superimposed onto the post-drive decay at 100 Hz (
The decay magnitude in
Additionally, the decay coefficient in
The optical performance of the system did not provide any insight other than there was no optical response to the AC modulation of analytes. However, this does provide some insight into the stiffness of the optical trap and the direction of motion the particles experience. The majority of optical transmission change seen by the SANE sensor is due to particles entering and leaving the trap. When a particle enters the trap, it serves as a dielectric lens and increases the intensity of light that is transmitted through the pore. If the AC modulation is causing a particle to oscillate in line with the nanopore without leaving the trap, it would likely not cause a noticeable change in the optical transmission. Neumeier et al. states that the time required for a particle to return to its favored state in the optical trap is on the order of pico-seconds. However, the particle must make its way through the trap in order for it to translocate and leave the trap. This could allow for the inline movement of the nanoparticle while in the trap. No oscillations occurred that conclusively resulted in a trapping or translocation event; therefore, the force preventing a nanoparticle from leaving the trap was likely greater than the force of the driven oscillation.
This Example presents data for the use of a SANE sensor for AC-, pulsed-, or modulation-driven plasmonic nanopore sensing. The disclosed results in this Example show that the AC method used with the SANE sensor could discriminate between Au and SiO2 nanoparticles of the same diameter. The model-deduced oscillation parameters during post-drive decay of the AC bursts both appeared to be concentration-dependent. At the lower concentrations used for each particle type, the difference in values between particle types for a given oscillation parameter became more pronounced. These types of AC measurements can be useful for the characterization of biological nanoparticles, such as liposomes, gene therapy vehicles, and drug delivery particles.
Peptide-presenting Major Histocompatibility Complex Class-I (pMHC) receptors being targeted by recombinant T-Cell Receptor-mimic (TCRm) antibodies can mediate the killing of specific cancer cells. High TCRm affinity combined with high avidity, meaning higher density binding to multiple ligand targets, can augment antitumor response significantly up to a limit set by autoimmunity. The cell copy number of pMHCs targeted by specific TCRms is an important determinant of avidity and therefore antitumor response. To advance the cancer immunotherapy field, technologies are needed to quantify both the number and heterogeneity of pMHC ligands in cells obtained from a patient tumor to select the antibodies with highest antitumor activity potential.
High-specificity tools that are needed to empirically assess whether a candidate TCRm antibody will target pMHCs in a patient's tumor are not readily available. Lower pMHC levels in cancer cells (10-100/cell) and expression heterogeneity in a tumor reduce the mean copy number. As a result, this limits the ability of commercial assays (enzyme-linked immunoassays (ELISAs), plasmon resonance techniques, antibody-dependent cellular cytotoxicity assays, complement dependent cytotoxicity assays, etc.) to detect many potential peptide targets. Since a needle biopsy collects >106 cells, back of the envelope calculations show that commercial assays require cancer cell expansion in culture to acquire enough protein to detect a given pMHC target. However, cell culture induces clonal bias, where many possible targets are eliminated after several passes.
In this Example, this disclosure presents the results of a new AC nanopore sensing method that can be used to differentiate the specific binding of an antigen and antibody from non-specific binding at ultra-low analyte concentrations, down to low attomolar (aM). This work is helpful in eliminating the need for cancer cell expansion in testing tumor pMHC heterogeneity.
Making pMHC's Commonly Overexpressed in GI, Breast, and Lung Cancers
In order to create pHMCs that are commonly overexpressed in gastrointestinal (GI), breast, and lung cancers, the recombinant peptides that are presented by HLA-A*0201 (HLA-A2) were synthesized, including cyclin-dependent kinase-2 (CdK-2; KIGEGTYGV), Systenin (SLMDHTIPEV), and TP53 (VVPCEPPEV).
FACS sorting results for yeast expression is presented herein in
The SANE sensor described in Example 1 was used herein. To test the sensitivity of the SANE sensor in detecting pMHCs and TCRms, titrations were performed for RAH, a pMHC peptide, and anti-RAH (TCRm) in homogeneous solutions as well as an equimolar heterogeneous solution of the mixture of the two (RAH-anti-RAH). Optical measurements, DC electrical measurements, and AC electrical measurements were taken.
The equimolar solution mixture was initially incubated at 100 nM and then titrated down to concentrations ranging from 1 aM to 10 fM, and the trapping of pMHC-TCRms was observed at all concentrations. These dilutions are several orders of magnitude beyond the preliminary data of 0.1 pM. These new results show the extreme sensitivity of this plasmonic nanosensor technology.
Subsequently, these DC optical-electrical data types were combined with the AC data types to generate multidimensional data sets for classifying and separating bound from unbound protein with higher accuracy. To expand the number of parameters used to identify trapped analytes, AC currents were introduced to the sensor to explore the frequency response of the sensor and the trapped proteins. A program was developed using Lab VIEW that delivers a pulse train that contains 10 cycles, each of 10, 20, 50, 100, 200, 500, 1000, 2000, 5000, 10000, 20000, 50000 and 100000 Hz (
Using the method for the SANE sensor described in Example 1, AC bursts were applied to solutions of 1 aM RAH, 1 aM anti-RAH, and an equimolar heterogeneous solution of the mixture of the two (1 aM RAH-anti-RAH). An example plot of the Ipatch current response (pA,
With reference once more to
Further,
To further display the differences in these parameters upon binding, the data were plotted in a 3D multimodal display.
This Example discloses results that show that the SANE sensor and AC method used with it are able to distinguish between bound and unbound antigen-antibody solutions at ultra-low analyte concentrations. Different dimensions, such as optical trapping time, decay coefficient, and decay magnitude, are able to differentiate between bound and unbound complexes. This will be useful for critical applications such as screening cancer patient tumors for pMHCs without the need for cell culture to generate enough protein material.
This Example describes an integrated isotachophoresis (ITP) platform that was developed to mount on top of the nanosensor described in Example 1 to concurrently increase the concentration of TCRm antibodies and target pMHCs while separating them from unbound proteins of different sizes and charge. For ˜104 cancer cells per assay, each cell expressing 10-100 copies of a specific pMHC, 1-10 nM of targeted complexes will exist in the concentrated analyte plug.
Before testing, to decontaminate the channel, the chip is cleaned with a 15% bleach solution, rinsed with deionized water, and vacuumed inside to remove any bleach residue. The channel is then flushed with LE solution several times before the TE reservoir is rinsed with deionized water to strongly dilute any LE solution residue before filling the reservoir with TE solution.
Because of the scarcity of the target protein samples (pMHC ligands and TCRms), low-cost samples comprising anionic dye-labeled Dextran conjugates weighing 40 kDa and 70 kDa were used to test the functionality of the experimental setup. Samples were diluted down to 0.25 ng/ml in TE solution before injection. Two sets of circular PDMS extensions were attached on top of the channels using plasma bonding to have better control over the volume added to the reservoirs and to help keep the electrodes stationary. Two sets of Pt electrodes were connected to the power supply (Sorensen XHR600-1.7 DC) and submerged into the cathode and anode reservoirs filled with TE and LE solutions, respectively. The loaded microchip was mounted above a 4× objective lens of an Olympus Confocal Microscope FV 3000. Constant voltage up to 600 V was applied by the power supply. TE solution containing HEPES and LE solution containing HCl were titrated to the same pH with Tris. HEPES was chosen because of its low electrophoretic mobility, and the chloride ion was selected for its high electrophoretic mobility. Tris was the counterion. The LE and TE solutions also contained 1% (w/v) polyvinylpyrrolidone to suppress the effect of electroosmotic flow.
Additional non-limiting, example embodiments are further described below.
Embodiment 1. A method of sensing comprising:
(1) providing a sensor comprising (a) a first layer having at least one single nanohole structure or at least one dual nanohole structure, and (b) a second layer having at least one nanopore, wherein the single nanohole structure comprises only one nanohole, wherein the dual nanohole structure comprises a first nanohole and a second nanohole connected by a gap, and wherein the one nanohole or the gap of the first layer is aligned with the nanopore of the second layer in a direction corresponding to a translocation direction across the first and second layers;
(2) providing a test sample comprising an analyte;
(3) contacting the test sample with the first layer of the sensor;
(4) irradiating the single nanohole structure or the dual nanohole structure of the first layer of the sensor with a beam of electromagnetic radiation;
(5) optically trapping the analyte in the single nanohole structure or in the dual nanohole structure and/or in the gap of the first layer of the sensor;
(6) applying a first electric field across the nanopore to draw one or more of the analytes into the nanopore, wherein the first electric field comprises a direct current (DC) electric field;
(7) applying a second electric field across the nanopore after applying the first electric field, wherein the second electric field comprises a pulsed, modulated, or alternating current (AC) electric field; and
(8) measuring one or more of: (a) change in current and/or phase across the nanopore during application of the second electric field while the analyte is optically trapped and/or during one or more translocation events of the analyte through the nanopore; or (b) at least one kinetic parameter of the analyte within the nanopore after removing or turning off the second electric field.
Embodiment 2. The method of Embodiment 1, wherein the at least one kinetic parameter is measured while the analyte decelerates or comes to a stop while optically trapped.
Embodiment 3. The method of Embodiment 1 or Embodiment 2, wherein the at least one kinetic parameter comprises one or more of the following: equilibrium dissociation constant (Kd), binding on-rate (kon), binding off-rate (koff), bound fraction (i.e., the fraction of analyte that is in a bound or complexed state rather than an unbound or uncomplexed state), analyte size (volume), analyte charge (effective charge on the outer surface of the analyte), and analyte conformation.
Embodiment 4. The method of any of the preceding Embodiments, wherein measuring change in current and/or phase further comprises determining a charge of a translocating analyte.
Embodiment 5. The method of any of the preceding Embodiments, wherein measuring change in current and/or phase further comprises determining a dielectric constant of a translocating analyte.
Embodiment 6. The method of any of the preceding Embodiments further comprising measuring a surface plasmon resonance of the single nanohole structure or the dual nanohole structure after optically trapping the analyte in the single nanohole structure or in the dual nanohole structure and/or in the gap of the first layer of the sensor.
Embodiment 7. The method of Embodiment 6, wherein measuring the surface plasmon resonance further comprises determining the mass of an optically trapped analyte.
Embodiment 8. The method of any of the preceding Embodiments, wherein the analyte comprises complexed and/or non-complexed biomolecules.
Embodiment 9. The method of any of the preceding Embodiments, wherein the test sample is a biological sample obtained from an animal or human subject, such as a human patient or animal patient in need of diagnosis.
Embodiment 10. The method of Embodiment 9, wherein the analyte comprises a pMHC or pMHC component.
Embodiment 11. The method of Embodiment 9, wherein the analyte comprises a HLA-A2 pHMC or HLA-A2 pMHC component.
Embodiment 12. The method of Embodiment 9, wherein the analyte comprises a TCRm antibody.
Embodiment 13. The method of Embodiment 9, where in the analyte comprises a TCRm antibody against a HLA-A2 pHMC or against a HLA-A2 pHMC component.
Embodiment 14. The method of any of Embodiments 1-7, wherein the analyte comprises an inorganic nanoparticle.
Embodiment 15. The method any of the preceding Embodiments, wherein the test sample is concentrated prior to contacting the test sample with the first layer of the sensor.
Embodiment 16. The method of Embodiment 15, wherein the test sample is concentrated using isotachophoresis (ITP).
Embodiment 17. The method of Embodiment 15, wherein the test sample is concentrated using an ITP microchannel structure.
Embodiment 18. The method of Embodiment 17, wherein the ITP microchannel structure is disposed over the first layer of the sensor.
Embodiment 19. The method of Embodiment 18, wherein the ITP microchannel structure forms a unitary chip with the first layer and the second layer of the sensor.
Various embodiments of the present invention have been described in fulfillment of the various objectives of the invention. It should be recognized that these embodiments are merely illustrative of the principles of the present invention. Numerous modifications and adaptations thereof will be readily apparent to those skilled in the art without departing from the spirit and scope of the invention.
This application claims priority to U.S. Provisional Patent Application No. 63/301,614 filed on Jan. 21, 2022, the entire contents of which are incorporated herein by reference.
This invention was made with government support under Grant No. 1R21CA240220-01A1 awarded by the National Cancer Institute and Grant No. T32 HL134613 awarded by the National Heart, Lung, and Blood Institute. The Government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2023/011259 | 1/20/2023 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63301614 | Jan 2022 | US |
