JP2022511068A - Difference sensor measurement method and equipment - Google Patents

Abstract

This specification describes semiconductor-based sensor devices and methods for detecting biological material. [Selection diagram] FIG. 1A

Description

Biosensors typically include a device that integrates a binding molecule (also referred to herein as a "detector molecule") with a signal transducer to detect the presence of a particular target (eg, an analyte). It can provide a signal to recognize. Biosensors can be based on a variety of physical principles, but generally rely on the binding of bound molecules to the target analyte. A specific bond or reaction between the bound molecule and the analyte can be measured by introducing a signal, which is transmitted. Biosensors can be configured to address macromolecular recognition of various types of human cells, viruses, pathogens, and the like. Therefore, these devices have a wide range of diagnostic usefulness such as human health, food safety, drug response, and personalized medicine applications.

One of the major issues in biosensor development is signal noise. For example, some sensors are connected to an external measuring device such as an ohmmeter, voltmeter, or ammeter for the purpose of detecting changes in resistance. In such cases, very long electrodes, cables, and other circuits are used because the sensor is in a large fluid and the connection must be made outside the fluid. As a result, electrodes and other components can experience very high contact resistance when coupled to the sensor. All of these factors cause significant noise and background in the measurement. Other sources of noise include inductive loops in the circuit, noise generated by the fluid in the region where the electrodes are in contact with the fluid, thermal noise, and those associated with other sources. This noise can be largely dependent on the overall length of the circuit and the quality of all contacts and can be much greater than the change in resistance due to the coupling of the analyte. In many cases, the noise and background are the same on each part of the surface of the sensor, as they are due to large external fluctuations. Therefore, a method of removing such systematic noise from measurements is highly desirable for improved biosensors.

This specification describes semiconductor-based sensor devices and methods for detecting biological material.

On one side, the device is provided. The device comprises a first semiconductor-based sensor with a first insulating layer functionalized with a detector molecule. The device further comprises a second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor. The second semiconductor-based sensor includes a second insulating layer. The first insulating layer and the second insulating layer are the electricity of the first semiconductor-based sensor when the first semiconductor-based sensor and the second semiconductor-based sensor are exposed to a fluid containing an analyte. The change in physical characteristics is configured to be greater than the corresponding change in electrical characteristics of the second semiconductor-based sensor.

On the other side, the device is provided. The device comprises a first semiconductor-based sensor with a first insulating layer functionalized with a detector molecule. The device further comprises a second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor. The second semiconductor-based sensor includes a second insulating layer. The thickness of the first insulating layer is at least twice the thickness of the second insulating layer.

On the other side, the method is described. The method comprises exposing the device to a fluid containing the analyte. The device comprises a first semiconductor-based sensor with a first insulating layer functionalized with a detector molecule. The device further comprises a second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor. The second semiconductor-based sensor includes a second insulating layer. The method is at least partially to determine the electrical characteristics of the difference based on the changes in the electrical characteristics of the first semiconductor-based sensor and the changes in the electrical characteristics of the second semiconductor-based sensor. Includes measuring said changes in the electrical properties of the first semiconductor-based sensor and corresponding changes in the electrical properties of the second semiconductor-based sensor.

In some embodiments, the second insulating layer is functionalized with the detector molecule.

In some cases, the change in electrical properties of the first semiconductor-based sensor is at least twice as large as the corresponding change in electrical properties of the second semiconductor-based sensor.

In some embodiments, when the first semiconductor-based sensor and the first semiconductor-based sensor are exposed to fluid and the analyte binds to the detector molecule of the first semiconductor-based sensor. The capacitance between the formed layers of charge is such that the second semiconductor and the second semiconductor-based sensor are exposed to the fluid and the analyte is the detector of the second semiconductor-based sensor. Greater than the capacitance between the layers of charge formed when binding to the molecule.

In some cases, the thickness of the second insulating layer is at least twice the thickness of the first insulating layer and at least 50 times the thickness of the first insulating layer.

In some embodiments, the thickness of the first insulating layer is 2 nm or more and 5 nm or less. In some cases, the thickness of the second insulating layer is 10 nm or more and 100 nm or less.

In some embodiments, the dielectric constant of the first insulating layer is greater than the dielectric constant of the second insulating layer.

In some embodiments, the first insulating layer and / or the second insulating layer comprises an oxide.

In some embodiments, the detector molecule comprises an antibody, a DNA fragment, and / or an RNA fragment. In some cases, the detector molecule comprises an enzyme. In certain embodiments, the detector molecule comprises glucose oxidase. In certain embodiments, the detector molecule comprises urease.

In some embodiments, the analyte comprises a biomolecule. In some cases, the analyte comprises monosaccharides and / or polysaccharides. In some examples, the analyte comprises glucose. In some embodiments, the analyte comprises urea.

In some embodiments, the electrical property is an electrical property selected from the group consisting of voltage, resistance, conductance, and current.

In some cases, the first semiconductor-based sensor and / or the second semiconductor-based sensor comprises one or more nanowires.

In some embodiments, the first semiconductor-based sensor is at least partially covered by the first insulating layer and / or the second semiconductor-based sensor is covered by the second insulating layer. At least partially covered.

In some embodiments, the first semiconductor-based sensor and / or the second semiconductor-based sensor comprises a field effect transistor.

In some cases, the device may, at least in part, obtain differential electrical characteristics based on changes in the electrical characteristics of the first semiconductor-based sensor and the electrical characteristics of the second semiconductor-based sensor. It is configured to output.

In some embodiments, the method comprises detecting the presence of the analyte, at least in part, based on the electrical properties of the difference.

In some embodiments, the electrical properties of the difference are determined by subtracting the changes in the electrical properties of the second semiconductor-based sensor from the changes in the electrical properties of the first semiconductor-based sensor. To.

Other aspects and embodiments will become apparent from the following detailed description of the various non-limiting embodiments of the invention, when considered in conjunction with the accompanying figures. If the specification and the document incorporated by reference contain conflicting and / or contradictory disclosures, the specification will prevail.

Non-limiting embodiments of the invention are exemplified by reference to the accompanying figures, but these figures are schematic and are not intended to be drawn to scale. In the figure, each of the same or substantially identical components shown is usually represented by a single number. For clarity, not all components are labeled on all figures and all components of each embodiment of the invention are not required to be illustrated for those skilled in the art to understand the invention. Nor is it shown.
FIG. 1A shows a schematic diagram of an exemplary semiconductor sensor in the absence of binding of the analyte to the detector molecule, according to an embodiment. FIG. 1B shows a schematic diagram of an exemplary semiconductor sensor in the presence of an analyte bound to a detector molecule, according to an embodiment. FIG. 2A shows a schematic diagram of a device including a first semiconductor sensor and a second semiconductor sensor according to an embodiment. FIG. 2B shows a schematic diagram of a device including a first semiconductor sensor and a second semiconductor sensor according to an embodiment. FIG. 3 shows a schematic diagram of an apparatus including a first semiconductor sensor including a first insulating layer and a second semiconductor sensor including a second insulating layer according to an embodiment. FIG. 4A shows an equivalent circuit diagram of two-point measurement. FIG. 4B shows an equivalent circuit diagram of a device including a first semiconductor sensor and a second semiconductor sensor according to an embodiment. FIG. 5A shows a schematic diagram of a non-limiting embodiment of a device comprising a first semiconductor sensor and a second semiconductor sensor. FIG. 5B shows a schematic diagram of a non-limiting embodiment of a device comprising a first semiconductor sensor and a second semiconductor sensor. FIG. 6A illustrates an exemplary process step for creating an apparatus according to an embodiment. FIG. 6B illustrates an exemplary process step for creating an apparatus according to an embodiment. FIG. 6C illustrates an exemplary process step for creating an apparatus according to an embodiment. FIG. 6D illustrates an exemplary process step for creating an apparatus according to an embodiment. FIG. 6E illustrates an exemplary process step for creating an apparatus according to an embodiment.

As used herein, devices and methods for detecting biological substances (eg, analytes) are described. The device may use an active detector sensor to detect the analyte in the fluid in contact with the device. The analyte may bind to the functionalized detector molecule on the surface of the sensor. Such coupling can result in a measurable signal to detect the presence of the analyte. As further described below, the device utilizes a reference (eg, inactive) sensor that does not convey a signal (or conveys a minimal signal compared to the signal of the detector sensor) in addition to the active detector sensor. You may. The reference sensor may be located in the immediate vicinity of the detector sensor. Changes in the electrical characteristics (eg, resistance) of the detector sensor compared to the reference sensor may be measured. In this way, the reference sensor can provide filtered baseline measurements from the signal generated by the active detector sensor (eg, to remove noise that may be present in the sensor environment). (Uses difference measurement technology), enabling more sensitive detection of analytes. Sensors are used to detect the presence of specific molecules that may be useful in food safety, drug response, personalized medicine, cancer detection, disease validation, and other medical and biological applications. can do.

1A-1B show an active semiconductor-based detector sensor 10 according to an embodiment. The detector molecule 12 is functionalized on the surface 14 of the sensor. As described below, the detector molecule is selected according to its ability to bind to the analyte to be detected. FIG. 1B shows the fluid 16 in contact with the sensor. As shown, the fluid comprises Analyte 18 and another species 20. In this embodiment, the analyte binds to the detector molecule. The coupling of the analyte causes a measurable change in the physical properties (eg, electrical properties) of the semiconductor. For example, in an exemplary embodiment, as shown in FIGS. 1A and 1B, the measured resistance change ΔR indicates the presence of an analyte. In this way, the sensor can be used to detect the concentration of an analyte in a fluid sample.

It should be understood that changes in physical properties different from resistance may be measured. For example, in some embodiments, changes in conductance (or ΔG) may be measured, and in some embodiments, changes in conductivity may be measured. In some embodiments, structural changes during binding of the detector molecules cause measurable changes. In certain embodiments, the change is due to electrical gating by the analyte. In some embodiments, the change is due to a change in surface plasmon resonance. In some embodiments, changes in characteristics (eg, conductance, resistance) may generally be detected electrically by applying a current to the sensor and measuring the change in voltage. In some embodiments, changes in properties (eg conductance, resistance) may generally be detected electrically by applying an alternating electric field.

2A-2B show an active semiconductor-based detector sensor 10 and a semiconductor-based detector reference (eg, inactive) sensor 11. The reference sensor is placed in close proximity to the detector sensor so that it is exposed to similar environments.

As shown in FIG. 2A, in some embodiments, both the reference sensor and the detector sensor include a detector molecule capable of binding to an analyte (not shown). In such an embodiment, the binding of the detector molecule to the analyte results in measurable changes in the characteristics of the detector's semiconductor (eg, electrical characteristics), but the same as the characteristics of the reference sensor's semiconductor (eg, electrical characteristics). Or does not result in any (arbitrary) measurable change.

As shown in FIG. 2B, in some embodiments, the detector sensor comprises a detector molecule capable of binding to an analyte (not shown) and the reference sensor does not include a detector molecule. In such an embodiment, the binding of the detector molecule to the analyte results in a measurable change in the properties of the detector sensor's semiconductor (eg, electrical properties), but the analyte is not bound to the surface of the reference sensor. Therefore, there is no measurable change in the semiconductor characteristics of the reference sensor.

In both embodiments, changes in the electrical characteristics (eg, resistance) of the detector sensor compared to the reference sensor can be measured. For example, changes in the electrical characteristics of a semiconductor-based detector sensor are at least twice, in some cases at least five times, and in some cases at least five times, as compared to the corresponding changes in electrical characteristics of a second semiconductor-based reference sensor. It is at least 10 times larger, and in some cases at least 100 times larger. The reference sensor provides filtered baseline measurements from the signal generated by the active detector sensor (eg, using differential measurement techniques to remove noise that may be present in the sensor environment) and analysis. Allows for more sensitive detection of objects. Suitable differential measurement techniques are described in co-owned US patent application serial numbers 14 / 510,178, filed October 9, 2014, which are incorporated herein by reference in their entirety.

FIG. 3 shows an active semiconductor-based detector sensor 10 and a semiconductor-based detector reference (eg, inactive) sensor 11. Both the detector sensor and the reference sensor include an insulating layer 22. As described above, the insulating layer of the detector sensor is thinner than the insulating layer of the reference sensor.

The insulating layer of the reference sensor is thick enough so that when the analyte (not shown) binds to the detector molecule, there is little or no measurable change in the semiconductor properties of the reference sensor. Can be selected as. For example, the thickness of the insulating layer of the reference sensor may be 10 nm or more, 25 nm or more, 50 nm or more, and in some embodiments, it may be 100 nm or less, 50 nm or less, or 10 nm or less. It should be understood that both the upper limit and the lower limit may be appropriate (for example, 10 nm or more and 100 nm or less).

In such an embodiment, the insulating layer of the detector sensor is sufficient to cause measurable changes in the semiconductor properties of the detector sensor when the analyte (not shown) binds to the detector molecule. Can be selected to be thinner. For example, the thickness of the insulating layer of the detector sensor may be 2 nm or more and 2 nm or more and 4 nm or more, and in some embodiments, it may be 10 nm or less, 5 nm or less, or 4 nm or less. .. It should be understood that both the above upper limit and the lower limit can be appropriate (for example, 2 nm or more and 5 nm or less).

The thickness of the insulating layer of the reference sensor may be 2 times or more, 10 times or more, or 25 times or more the thickness of the insulating layer of the detector sensor. The thickness of the insulating layer of the detector sensor may be 50 times or less, 25 times or less, or 10 times or less the thickness of the insulating layer of the reference sensor.

Any suitable insulating layer can be used, including an oxide film (eg, aluminum oxide, silicon oxide) or a nitride film.

Other techniques may also be utilized to manufacture reference sensors and detector sensors. For example, the reference sensor may include a high insulation layer and the detector sensor may include a low insulation layer of similar thickness in order to generate a signal difference. In some embodiments, the reference sensor is gated with an external voltage to further control its resistance.

In some embodiments, the devices described herein may include an active sensor region, a reference (eg, inactive) sensor region, a voltage input electrode, a voltage measuring electrode, and a voltmeter. As an example, each sensor area is here composed of functionalized nanowires and pads on which electrodes are connected. The pad provides a continuous path from the electrode to the sensor.

In some embodiments, the two sensors are arranged parallel to each other. In some embodiments, the sensor is vertical. In one embodiment, the sensors are arranged linearly. Other geometric shapes that allow difference measurements can also be used.

In the difference technique, a constant voltage Vin is applied to one side of both sensor regions, and as a result, the voltage difference that occurs at the other end of the sensor is measured. In some embodiments, lock-in amplifiers are employed to supply and measure voltage. In some embodiments, due to the non-ohmic nature of semiconductor sensors (eg, nanowires), additional source-drain voltages may be applied to increase sensitivity. Universal noise and constant background signals will be present at the same level on the electrodes of both sensors and will be systematically subtracted, leaving only the resistance change of the active sensor to the reference sensor.

With reference to FIG. 4, the sensor can be thought of as two parallel resistances. The input voltage Vin is applied to both sensors and the resulting voltage difference between the sensors is measured. _R0 is the bare resistance of the sensor, and ΔR is the change in resistance due to the coupling of the analyte. Here, for the purpose of demonstration, it is assumed that R ₀ is the same for the active sensor and the inactive sensor. This analysis can be easily extended if the bare resistance can be different or if the resistance of the inactive sensor can be gate controlled. The noise may appear as an additional δR _i for each temporally random sensor i, and the temporally random noise δV_E, i may appear for each individual post-sensor electrode labeled by i. Since much of this noise generated by thermal fluctuations, fluid flow, and other large sources is the same in each part of the circuit, δR ₁ = δR ₂ , δV ₁ = δV ₂ at all times. In some measurements, V ₁ and V ₂ are the opposite of the definition here.

The measured voltage can be solved by a simple circuit analysis. For very small resistance changes expected in the present invention, ΔV = V ₁ − V ₂ = Vin * _δR / R ₀ .

The small δRs considered here are for demonstration purposes. Even larger resistance changes where V ₁ -V ₂ become non-linear at δR can be measured with significantly reduced noise.

Any suitable semiconductor-based sensor device can be used for the sensor. For example, suitable equipment is described in co-owned US literature. There are patent application serial numbers 14 / 510,178 filed on October 9, 2014 and co-owned international application numbers PCT / US2015 / 041527 published as international publication number WO2016 / 089453, all of which are: The whole is incorporated herein by reference.

Sensors include semiconductor materials. Suitable semiconductor materials from which nanosensors can be made include, but are not limited to, silicon, germanium, III-V semiconductors and the like. In some embodiments, the sensor comprises silicon. The sensor may include one or more layers formed above or below the semiconductor material, such as an additional semiconductor material layer or insulating layer (eg, an oxide such as silicon oxide).

In some embodiments, the semiconductor is patterned into nanowires (s). The sensor may be, in some embodiments, a field effect transistor sensor (eg, a nanosensor). In general, a field effect transistor (FET) uses an electric field to control a conductive channel, thus controlling the conductivity of charge carriers in the channel. By applying an electric field to the gate to change the size and shape of the conductive channel, the flow of charge carriers between the source and drain can be adjusted. In an exemplary biosensor configuration, the FET comprises a nanosensor (eg, nanowire) channel between the source and drain terminals. The surface of a nanosensor (eg, nanowire) can be biofunctionalized so that a biomolecular binding phenomenon can create an electric field similar to the controlled electric field applied to a conventional FET (FIG. 1). In some devices that use the FET principle, a designated and physically separated sensor surface can be formed by precision manufacturing. A FET sensor can be connected to an electronic circuit to monitor a particular conductance on the sensor surface. In some embodiments, many independent electronic circuits may be queried in massively parallel fashion in operation. FET biosensors can be adapted for the measurement of biomolecules that interact with such sensor surfaces (FIGS. 1A and 1B).

In some embodiments, the sensor comprises a nanoscale silicon-based FET device. Many of these devices exhibit the sensitivity, reliability, robustness, and sensor flexibility required for many multiple diagnostic microarrays. In some cases, nanoscale devices can be developed and / or mounted on conventional top-down silicon. In some cases, developing and implementing nanoscale equipment on top-down silicon to improve the reliability and robustness of top-down silicon semiconductor manufacturing processes and test them both in the clinic and in central laboratories. Error rate can be reduced. This will, in some cases, increase the effectiveness of each patient's visit to the laboratory or clinic, reduce diagnostic costs, and enable early diagnosis, treatment, and monitoring.

In certain embodiments, the nanosensor is a silicon nanochannel field effect transistor (FET) biosensor. Such sensors can be used to detect sensitive and / or label-free analytes. Such sensors can have excellent electrical properties and small dimensions. In certain embodiments, silicon nanochannels are ideally suitable for ultrasensitivity. In some cases, the high surface area-to-volume ratios of these systems allow the detection of single molecules. In some cases, such FET sensors (eg, biosensors) offer the advantages of high speed, low cost, high yield manufacturing without sacrificing the sensitivity typical of traditional optical techniques in diagnostics. do. Top-down manufacturing methods can be used to take advantage of complementary metal oxide semiconductor (CMOS) technology to achieve a richly multiplexed sensor array. Examples of nanochannel-based sensor systems are described, for example, in International Patent Publication WO2008 / 063901A1 by Yuchen et al. And International Patent Publication WO2009 / 124111A1 by Mohanti et al., Each of which is by reference for all purposes. The whole is built in.

In some embodiments, the semiconductor-based sensor can be part of a bias and measurement circuit. In some embodiments, the bias and measurement circuit operates by applying a bias voltage across the nanosensor (eg, nanochannel) in the circuit. The bias voltage can be selected to be sufficiently negative to achieve the desired dependence of the differential conductance of the detection element on the surface potential of the sensor (eg, the surface potential of the nanochannel). In one embodiment, this dependency has a steep slope region with a high amplification factor that is substantially greater than the reference amplification factor indicated by the zero bias state detector, thereby resulting in a relatively high signal vs. noise. Achieve the ratio. The bias and measurement circuit, in some embodiments, measures the differential conductance of the detection element and converts the measured differential conductance into a signal indicating the presence or activity of the analyte. In certain embodiments, the measured derivative conductance can be converted into a signal indicating the presence or activity of the analyte by using a look-up table or alternative conversion mechanism that reflects the previous calibration operation. In some embodiments, the applied gate voltage can be used to control the sensitivity of the sensor. According to certain embodiments, the bias gate voltage and the reference gate voltage can be used independently to control the sensitivity.

As mentioned above, at least a portion (eg, surface) of the sensor is functionalized with a detector molecule. The detector molecule may include, for example, an antibody, enzyme, protein, peptide, small molecule, nucleic acid, aptamer, acceptor molecule, polymer, and / or supermolecular structure. The detector molecule may be designed to be particle specific. That is, only one particular analyte binds to a particular detector molecule. In some embodiments, the detector molecule is an antibody. In certain embodiments, the detector molecule is a DNA or RNA fragment or comprises a DNA or RNA fragment. In some embodiments, the detector molecule is glucose oxidase or comprises glucose oxidase. In some cases, the detector molecule is or contains urease.

Analysts include, for example, proteins, small molecules, nucleic acids, peptides, antibodies, aptamers, biomarkers, genes, virus particles, supermolecular structures, polymers, receptor molecules, living cells, and / or living cell clusters. You may be. In some cases, the analyte comprises a biomolecule. In some examples, the analyte may be a protein biomarker or a genetic biomarker. In certain embodiments, the analyte is a monosaccharide and / or a polysaccharide, or comprises a monosaccharide and / or a polysaccharide. For example, in some cases, the analyte is or contains glucose. One of the non-limiting examples of small molecule analytes is urea.

It should be understood that the methods described herein are not limited to a particular detector molecule or analyte.

In some embodiments, the analyte and the detector molecule may be attached via a chemical interaction such as a chemical bond. The chemical bond may be a covalent bond or a non-covalent bond. In some cases, the chemical bond is a non-covalent bond such as a hydrogen bond, an ionic bond, a coordinate bond, and / or a van der Waals interaction. One or more of the species and / or substances (eg, analytes, detector molecules) may contain functional groups capable of forming such bonds. Covalent and non-covalent bonds between components can be formed by any type of reaction, as is known to those of skill in the art, using appropriate functional groups to undergo such reactions. Please understand that. Chemical interactions suitable for use in the various embodiments described herein can be readily selected by one of ordinary skill in the art based on the description herein.

In some embodiments, the association between the analyte and the detector species may occur via a biological binding phenomenon (ie, between a pair of complementary biological molecules). .. For example, the analyte or detector molecule may include a physical substance such as biotin that specifically binds to a complementary physical substance such as avidin or streptavidin on another species or substance. Other examples of biological molecules that can form biological bonds between pairs of biological molecules include, but are not limited to, proteins, nucleic acids, glycoproteins, carbohydrates, hormones, and the like. Non-limiting examples include antibody / peptide pairs, antibody / antigen pairs, enzyme / substrate pairs, enzyme / inhibitor pairs, enzyme / cofactor pairs, protein / substrate pairs, nucleic acids / nucleic acids. Pairs, protein / nucleic acid pairs, peptide / peptide pairs, protein / protein pairs, small molecule / protein pairs, acceptor / hormone pairs, acceptor / effector pairs, ligand / cell acceptor pairs, biotin Includes, but is not limited to, a / avidin pair, a biotin / streptavidin pair, a drug / target pair, a small molecule / peptide pair, a small molecule / protein pair, a small molecule / enzyme pair, and the like. Biological interactions between species and / or substances (s) used in the embodiments described herein are described herein for these functions, such biological interactions. Can be readily selected by one of ordinary skill in the art, based on the present specification and knowledge in the art regarding the examples of, and simple techniques for identifying suitable biological interactions.

In certain embodiments, the analyte and the detector species may be associated with each other via physical interaction. For example, in some embodiments, the analyte (eg, supramolecular structure) may be physically entangled with at least a portion of the detector species (eg, macromolecules).

In certain embodiments, the analyte and detector species may be associated with each other via a linking moiety (eg, other biological or chemical species that brings the analyte and detector species close together). For example, the shortest distance between the analyte and the detector species that are related to each other may be longer than the Debye length. In some examples, the shortest distance may be about 100 nanometers or less, about 50 nanometers or less, about 25 nanometers or less, about 10 nanometers or less, or about 1 nanometer or less.

5A and 5B show an embodiment of a differential resistance biosensor integrated on a microchip. An enlarged view of the sensor region shown in FIG. 5B, comprising an active sensor (semiconductor nanowires connected to a semiconductor pad), an inactive sensor deactivated by non-functionalization, and a voltage output electrode 4. In some embodiments, the electrodes are coated with an insulating layer that includes, but is not limited to, AL ₂ O ₃ , SiO ₂ , HfO ₂ , or Si ₃ N ₄ . FIG. 5 shows a complete microchip containing a bonding pad that connects to an external circuit. In this embodiment, the sensors are designed parallel to each other using nanowires parallel to the electrodes. Some embodiments will have different bonding pad configurations and different electrode configurations.

This is a non-limiting example of the process by which a representative sensor is made, including an exemplary geometry diagram. The sensor is typically made by overlaying a thin silicon layer, typically about 100 nm, on top of a silicon dioxide layer, typically about 200 nm thick.

In some embodiments, ions are injected into the silicon in the area where the connection pad is made. The ions can be metals or other dopants, as described above.

Next, nanowires and electrode mounting pads are created by electron beam lithography and reactive ion etching. This removes silicon everywhere except the sensor area. In some embodiments, the definition of nanowires and pads is divided into two steps, with nanowires and pads defined separately. In some embodiments, nanowires are defined by electron beam lithography and pads are defined by photolithography. In some embodiments, the pads and nanowires are all defined by photolithography. Our invention generally covers all methods of making nanowires and pads.

After that, a pattern of metal electrodes having a four-point measurement shape is created by photolithography, and the metal electrodes are deposited. The electrode contains a metal such as Au, Cu, Ag, Al, and may be an alloy or a metal multilayer. Often, an adhesive layer of Ti, Ta, or other metal is used. The electrodes are about 10-20 microns wide near the sensor and wider as they are farther away. Typically, the thickness of the metal is in the range of 100 nm.

The electrodes are then coated with a thick insulating barrier, typically about 100 nm, typically an oxide such as Al ₂ O ₃ , SiO ₂ , HfO ₂ , or Si ₃ N ₄ . The thick barrier also covers one of the sensors, deactivating the sensor.

The sensor and electrodes are then coated with a thin insulating barrier (typically 10 nm) on top made of similar oxides.

Finally, metal pads are deposited to allow connection to external measuring instruments. The pad is typically a highly conductive noble metal such as Au, Ag, Cu with a thickness of 1-2 microns. These usually include a thin (10 nm thick) Ti-like adhesive layer. The final metal deposition process may also include the deposition of additional electrodes to further improve the sensor. A schematic diagram of the completed sensor ready to be integrated into the final circuit is shown in FIG.

Claims (45)

A first semiconductor-based sensor with a first insulating layer functionalized with a detector molecule,
A second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor, comprising the second semiconductor-based sensor with a second insulating layer.
The first insulating layer and the second insulating layer are the electricity of the first semiconductor-based sensor when the first semiconductor-based sensor and the second semiconductor-based sensor are exposed to a fluid containing an analyte. A device configured such that the change in physical characteristics is greater than the corresponding change in electrical characteristics of the second semiconductor-based sensor. A first semiconductor-based sensor with a first insulating layer functionalized with a detector molecule,
A second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor, the second semiconductor-based sensor including the second insulating layer, and the thickness of the first insulating layer is the second. A device that is at least twice the thickness of the insulating layer. The device according to claim 1 or 2, wherein the second insulating layer is functionalized by the detector molecule. The apparatus according to claim 1, wherein the change in the electrical characteristics of the first semiconductor-based sensor is at least twice as large as the change in the corresponding electrical characteristics of the second semiconductor-based sensor. The first semiconductor-based sensor and a layer of charge formed when the first semiconductor-based sensor is exposed to fluid and the analyte binds to the detector molecule of the first semiconductor-based sensor. The capacitance between the second semiconductor and the second semiconductor-based sensor is formed when the second semiconductor-based sensor is exposed to the fluid and the analyte binds to the detector molecule of the second semiconductor-based sensor. The device according to any one of claims 1 to 4, which is larger than the capacitance between the layers of charge. The thickness of the second insulating layer is at least twice the thickness of the first insulating layer and at least 50 times the thickness of the first insulating layer according to any one of claims 1 to 5. The device described. The apparatus according to any one of claims 1 to 6, wherein the thickness of the first insulating layer is 2 nm or more and 5 nm or less. The apparatus according to any one of claims 1 to 7, wherein the thickness of the second insulating layer is 10 nm or more and 100 nm or less. The apparatus according to any one of claims 1 to 8, wherein the dielectric constant of the first insulating layer is larger than the dielectric constant of the second insulating layer. The apparatus according to any one of claims 1 to 9, wherein the first insulating layer and / or the second insulating layer contains an oxide. The device according to any one of claims 1 to 10, wherein the detector molecule comprises an antibody, a DNA fragment, and / or an RNA fragment. The device according to any one of claims 1 to 11, wherein the detector molecule contains an enzyme. The device according to any one of claims 1 to 12, wherein the detector molecule contains glucose oxidase. The device according to any one of claims 1 to 13, wherein the detector molecule contains urease. The apparatus according to any one of claims 1 to 14, wherein the analyte comprises a biomolecule. The apparatus according to any one of claims 1 to 15, wherein the analyte comprises a monosaccharide and / or a polysaccharide. The apparatus according to any one of claims 1 to 16, wherein the analyte comprises glucose. The apparatus according to any one of claims 1 to 17, wherein the analyte comprises urea. The device according to any one of claims 1 to 18, wherein the electrical characteristic is an electrical characteristic selected from the group consisting of voltage, resistance, conductance, and current. The device according to any one of claims 1 to 19, wherein the first semiconductor-based sensor and / or the second semiconductor-based sensor includes one or more nanowires. The first semiconductor-based sensor is at least partially covered by the first insulating layer and / or the second semiconductor-based sensor is at least partially covered by the second insulating layer. The device according to any one of claims 1 to 20. The device according to any one of claims 1 to 21, wherein the first semiconductor-based sensor and / or the second semiconductor-based sensor includes a field-effect transistor. The device is to output differential electrical characteristics, at least in part, based on changes in the electrical characteristics of the first semiconductor-based sensor and changes in the electrical characteristics of the second semiconductor-based sensor. The apparatus according to any one of claims 1 to 22, which is configured. A first semiconductor-based sensor having a first insulating layer functionalized with a detector molecule, and a second semiconductor-based sensor that telecommunicationss with the first semiconductor-based sensor, said to have a second insulating layer. Exposing the device with the second semiconductor-based sensor to the fluid containing the analyte,
The first to determine the electrical characteristics of the difference, at least in part, based on the changes in the electrical characteristics of the first semiconductor-based sensor and the changes in the electrical characteristics of the second semiconductor-based sensor. A method comprising measuring said change in said electrical property of a semiconductor-based sensor and said change in said corresponding electrical property of said second semiconductor-based sensor. 24. The method of claim 24, comprising detecting the presence of the analyte, at least in part, based on the electrical properties of the difference. 24. 25. The method according to any one of claims 24 to 26, wherein the second insulating layer is functionalized by the detector molecule. A layer of charge formed when the first semiconductor-based sensor and the first semiconductor-based sensor are exposed to the fluid and the analyte binds to the detector molecule of the first semiconductor-based sensor. The capacitance between the second semiconductor and the second semiconductor-based sensor is when the second semiconductor-based sensor is exposed to the fluid and the analysis material binds to the detector molecule of the second semiconductor-based sensor. The method of any one of claims 24-27, which is greater than the capacitance between the layers of charge formed. According to any one of claims 24 to 28, the thickness of the second insulating layer is at least twice the thickness of the first insulating layer and at least 50 times the thickness of the first insulating layer. The method described. The method according to any one of claims 24 to 29, wherein the thickness of the first insulating layer is 2 nm or more and 5 nm or less. The method according to any one of claims 24 to 30, wherein the thickness of the second insulating layer is 10 nm or more and 100 nm or less. The method according to any one of claims 24 to 31, wherein the dielectric constant of the first insulating layer is larger than the dielectric constant of the second insulating layer. The method according to any one of claims 24 to 32, wherein the first insulating layer and / or the second insulating layer contains an oxide. The method according to any one of claims 24 to 33, wherein the detector molecule comprises an antibody, a DNA fragment, and / or an RNA fragment. The method according to any one of claims 24 to 34, wherein the detector molecule comprises an enzyme. The method according to any one of claims 24 to 35, wherein the detector molecule comprises glucose oxidase. The method according to any one of claims 24 to 36, wherein the detector molecule comprises urease. The method according to any one of claims 24 to 37, wherein the analyte comprises a biomolecule. The method according to any one of claims 24 to 38, wherein the analyte comprises a monosaccharide and / or a polysaccharide. The method according to any one of claims 24 to 39, wherein the analyte comprises glucose. The method according to any one of claims 24 to 40, wherein the analyte comprises urea. The method according to any one of claims 24 to 41, wherein the electrical characteristic is an electrical characteristic selected from the group consisting of voltage, resistance, conductance, and current. The method according to any one of claims 24 to 42, wherein the first semiconductor-based sensor and / or the second semiconductor-based sensor comprises one or more nanowires. The method according to any one of claims 24 to 43, wherein the first semiconductor-based sensor and / or the second semiconductor-based sensor includes a field effect transistor. The first semiconductor-based sensor is at least partially covered by the first insulating layer and / or the second semiconductor-based sensor is at least partially covered by the second insulating layer. The method according to any one of claims 24 to 44.

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