JP2009520947A - Structure and manufacturing method of photonic crystal biosensor - Google Patents

Abstract

  The present invention provides a sensor configuration and a method of making a sensor.

Description

Government Involvement This invention was made with government support under grant number BES04-27657 awarded by the National Science Foundation. The government has certain rights in the invention.

  Label-free optical sensors based on surface structured photonic crystals have recently been demonstrated as a sensitive method for performing a wide variety of biochemical and cell-based assays. See, for example, Cunningham et al., Label-Free Assays on the BIND System. See Journal of Biomolecular Screening, 2004. 49: 481-490. These sensors reflect only a narrow band of wavelengths when illuminated at normal incidence with white light, and a positive shift in this reflected peak wavelength value (PWV) was detected. The adsorption of the substance on the sensor surface is shown. See, for example, Cunningham et al., Colorimetric resonant reflection as a direct bioassay assay technique. See Sensors and Actuators B, 2002.81: 316-328. By localizing incident photons separately at the resonant wavelength, a high light field is generated at the sensor surface that extends a short distance in the test sample, much like an evanescent field. The high degree of spatial confinement of resonant photons in the device structure provides the ability to perform strong interactions between this structure and adsorbed biological material, as well as high resolution imaging of protein and cell adhesion. . See, for example, Li et al., A new method for label-free imaging of biomolecular interactions. See Sensors and Actuators B, 2004.99: 6-13.

Previously, photonic crystal optical biosensors were manufactured from a continuous sheet of plastic film using a process in which periodic surface structures are replicated from a silicon master wafer using a UV curable polymer material. See, for example, Cunningham, et al., A plastic colorimetric resonant optical biosensor for multiple parallel detection of label-free biochemical interactions. See Sensors and Actuators B, 2002.85: 219-226. This patterned polymer can be subsequently coated with a high refractive index TiO ₂ layer that is generally thinner than the height of the surface structure. Such devices have been demonstrated for a wide variety of biochemical and cell-based assays with density detection sensitivity resolution of less than 0.1 pg / mm ² and a large dynamic range that allows detection of single cells. Yes. See, e.g., Lin et al., A label-free biosensor-based cell attachment assignment for char surface molecules. See Sensors and Actuators B, Accepted April 2005. In general, optimizing device sensitivity requires increasing the interaction between the electromagnetic field strength distribution and the molecules deposited on the photonic crystal surface. Therefore, the selection of the optical material and the design of the surface structure topology is based on the electromagnetic field profile from the inner region of the photonic crystal (which cannot interact with the adsorbed material here) to the region adjacent to the photonic crystal containing the liquid test sample. The purpose is to stretch.

  There is a need in the art for methods to increase the sensitivity of these and other types of sensors and reduce the cost of their manufacture.

SUMMARY OF THE INVENTION One embodiment of the present invention comprises a sensor comprising a nanoporous material having a low refractive index, supported on a bottom surface by a substrate, and coated on a top surface with a high dielectric constant dielectric coating. provide. A high dielectric constant dielectric coating or a high dielectric constant dielectric coating in combination with a nanoporous material forms a sub-wavelength grating structure. When the sensor is illuminated, a resonant grating effect occurs in the reflected radiation spectrum, and the depth and spacing of the subwavelength spacing grating structure is smaller than the wavelength of the resonant grating effect. When the sensor is illuminated with a broad band of optical wavelengths, a narrow band of optical wavelengths can be reflected from the sensor. The refractive index of the nanoporous material can be about 1.1 to about 2.2. In another embodiment, the nanoporous material may have a refractive index of about 1.1 to about 1.5. The spacing of the sub-wavelength spacing grating structure may be about 50 nm to about 1,500 nm, and the depth of the sub-wavelength spacing grating structure may be about 50 nm to about 900 nm. The nanoporous material may be porous silica xerogel, porous airgel, porous hydrogen silsesquioxane, B-stage polymer, porous methylsilsesquioxane, porous poly (arylene ether) or a combination thereof. May be. The substrate may include glass, plastic, or epoxy. The dielectric coating may have a refractive index of about 1.8 to about 3.0. The dielectric coating may include tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or combinations thereof. The refractive index of the substrate may be about 1.4 to about 1.6. The dielectric coating may have a thickness of about 30 nm to about 700 nm, and the nanoporous material may have a thickness of about 10 nm to about 5,000 nm. The dielectric coating may have a cover layer on its top surface. The sensor may further include one or more specific binding substances immobilized on the high dielectric constant dielectric coating. The sensor may further include one or more specific binding substances immobilized on the cover layer. The one or more specific binding substances may not include a detection label. The one or more specific binding substances may be bound to their binding partners. The one or more specific binding substances and the binding partner may not include a detection label. The one or more specific binding substances may be arranged in an array on the high dielectric constant dielectric coating. The one or more specific binding substances are arranged in an array on the cover layer.

  Another embodiment of the present invention provides a sensor comprising a waveguide structure formed by a waveguide film covering a substrate, the waveguide film having a refractive index higher than that of the substrate and diffracting A grating is included in the waveguide structure, and the diffraction grating is composed of a nanoporous material having a low dielectric constant. The nanoporous material may have a refractive index of about 1.1 to about 1.5. In another embodiment of the present invention, the refractive index of the nanoporous material may be about 1.1 to about 2.2. The nanoporous material may be porous silica xerogel, porous airgel, porous hydrogen silsesquioxane, B-stage polymer, porous methylsilsesquioxane, porous poly (arylene ether) or a combination thereof. May be. The substrate may include glass, epoxy or plastic. The waveguide film includes tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or a combination thereof. The waveguide film includes a polymer. The sensor further includes one or more specific binding substances immobilized on the waveguide film. The one or more specific binding substances may not include a detection label. The one or more specific binding substances may be bound to their binding partners. The one or more specific binding substances and the binding partner may not include a detection label. The one or more specific binding substances may be arranged in an array on the high refractive index dielectric coating.

  The use of a very low refractive index material for the surface structure in the sensor substantially increases the detection sensitivity. Thus, substances can be measured in test samples at concentrations, molecular weights or binding affinities that are 2-4 times lower than previously possible.

  One embodiment of the present invention is used to detect organic or inorganic substances such as proteins, DNA, small molecules, viruses, cells and bacteria, especially without the need for labels such as fluorescent or radioactive labels. Provide a sensor that can The photonic crystal sensor of the present invention reflects only a very narrow band of wavelengths or a single wavelength when illuminated with a broad wavelength light source (eg, white light or LED). This reflected color shifts to longer wavelengths in response to the binding of the substance to the sensor surface. The photonic crystal structure of the present invention provides a sensitivity that is 2 to 4 times higher than the previously described structure. An important difference in the sensor structure to obtain higher sensitivity is to replace the polymer sub-wavelength grating structure with a nanoporous low refractive index material.

  A method for manufacturing a sensor structure that enables low cost manufacturing is also disclosed. The sensor structure of the present invention has a higher sensitivity than conventional structures because it uses a nanoporous low refractive index material instead of a polymer sub-wavelength grating structure. If the refractive index of the sensor structure immediately below (or including) the subwavelength grating structure is lower than the refractive index of any liquid used in the sample, the photonic crystal's electromagnetic field will interact more strongly with the test sample. In effect, a structure is obtained in which the reflected wavelength is more strongly tuned by a predetermined amount of adsorbed biological material. This system can, for example, detect a single cell bound to its surface.

  The principles of the present invention can be applied to any biosensor that incorporates evanescent wave based biosensors and optical waveguides, for example. See, for example, US Pat. No. 4,815,843; US Pat. No. 5,071,248; US Pat. No. 5,738,825.

  This sensor is inter alia used for pharmaceutical research (eg high-throughput screening, secondary screening, quality control, cytotoxicity, clinical trial evaluation), life science research (eg proteomics, protein interaction analysis, DNA-protein interaction). Analysis, enzyme-substrate interaction analysis, cell-protein interaction analysis), diagnostic tests (eg protein presence, cell identification) and environmental detection (bacteria and spore detection and identification) . Previous patent applications and publications show that the surface of a photonic crystal biosensor, in combination with a high-resolution imager, uses many nanoliters of sample material in parallel on a single surface for many biochemical A method that can be used as a platform for conducting the assay is described. For example, see the following U.S. Patent Publication Numbers: 2002/0168295; 2002/0127565; 2004/0132172; 2004/0151626; 2003/0027328; 2003/0027327; 2003/017581; 2003/0068657; 2003/0059855; 2003 2003/0160766; 2003/0026891; 2003/0026891; 2003/0032039; 2003/0017580; 2003/0077660; 2004/0132214.

Photonic Crystal Sensor Use the photonic crystal sensor of the present invention to create sharp optical resonant reflections at specific wavelengths that can be used to track molecular interactions such as biological materials with high sensitivity Can do.

  The photonic crystal sensor includes a surface with a subwavelength structure. A subwavelength structured surface is a type of diffractive optical element that can mimic the effect of a thin layer coating. See, for example, Peng & Morris, “Resonant Scattering from Two-Dimensional Gratings” Opt. Soc. Am. A, Vol. 13, No. 5, 993, May 1996; Magnusson, & Wang, “New principal for optical filters”, Appl. Phys. Lett. , 61, 9, p. 1022, August, 1992; Peng & Morris, "Experimental demonstration of resonant anomalies in diffractive from two-dimensional gratations, p. 49, Op. thing. The grating of the photonic crystal sensor of the present invention has a small grating spacing compared to the wavelength of incident light, so that no diffraction orders other than this reflected and transmitted zeroth order are allowed. The photonic crystal sensor may include a grating, which is composed of or coated with a high dielectric constant dielectric material sandwiched between a substrate layer and a cover layer that fills the grooves of the grating. The A cover layer is not used if necessary. This grating structure selectively couples to light of a narrow band of wavelengths. This sensitive coupling condition can cause a resonant grating effect on the reflected radiation spectrum, thereby producing a narrow band of reflected or transmitted wavelengths. The depth and spacing of this grating is shorter than the wavelength of the resonant grating effect.

  The reflected or transmitted color of the photonic crystal sensor structure can be modified by the addition of molecules such as specific binding substances or binding partners, or both, to the top surface of the cover layer or lattice surface. This added molecule increases the optical path length of the incident radiation through the sensor structure, thereby modifying the wavelength at which maximum reflectance or transmission occurs.

  In one embodiment, the sensor is designed to reflect a single wavelength or a narrow band of wavelengths when illuminated with white light. As molecules bind to the sensor surface, this reflected wavelength (color) shifts due to a change in the optical path of the light coupled to the grating. By immobilizing molecules, eg, specific binding agents, on the sensor surface, complementary binding partner molecules can be detected without the use of any kind of fluorescent probe or particle label. This detection technique can be performed on the sensor surface immersed in liquid or dried.

  When illuminating the photonic crystal sensor with parallel white light, only a narrow band of wavelengths is reflected, or a single band of wavelengths is reflected. This narrow band of wavelengths is described as the wavelength “peak”. This “peak wavelength value” (PWV) changes as molecules are deposited on or removed from the sensor surface. The reader illuminates parallel white light at discrete locations on the sensor surface and collects the parallel reflected light. This collected light is collected in a wavelength spectrometer and the PWV is determined.

  FIG. 1 shows the structure of the photonic crystal sensor of the present invention. The sensor includes a substrate, a patterned, low-k, nanoporous material and a substantially uniform high refractive index coating. The surface of the low-k nanoporous material is patterned into a sub-wavelength spacing lattice structure, and a high refractive index material is deposited thereon.

In general, the low-k dielectric materials of the present invention have a dielectric constant k of about 1.1 to about 3.9. Examples of low-k dielectric materials include, for example: fluorosilicate glass (about 3.2 to about 3.9); polyimide (about 3.1 to about 3); hydrogen silsesquioxane ( HSQ) (about 2.9 to about 3.2); diamond-like carbon (about 2.7 to about 3.4); black diamond (SiCOH) (about 2.7 to about 3.3); Parylene-N ( About 2.7); B-staged polymers (CYCLOTENE ^™ and SiLK ^™ ) (about 2.6 to about 2.7); fluorinated polyimide (about 2.5 to about 2.9); methyl silsesquioxane ( MSQ) (about 2.6 to about 2.8); poly (arylene ether) (PAE) (about 2.6 to about 2.8); fluorinated DLC (about 2.4 to about 2.8); Parylene-F (about 2.4 to about 2.5); PTFE (about 1.9); Ka xerogels and aerogels (about 1.1 to about 2.2); porous Hidorogen silsesquioxane (HSQ) (about 1.7 to about 2.2); porous SiLK ^TM (B-stage polymer) ( About 1.5 to about 2.0); porous methylene silsesquioxane (MSQ) (about 1.8 to about 2.2); porous poly (arylene ether) (PAE) (about 1.8). To about 2.2).

The low-k nanoporous material is an inorganic, porous, oxide-like low dielectric material with an index of refraction n of about 1.1 to about 2.2, and preferably about 1.1 to about 1.5. is there. Low-k nanoporous materials include, for example, porous silica xerogels and aerogels (about 1.1 to about 2.2); porous HSQ (about 1.7 to about 2.2); porous SiLK ^™ (stage B Polymer) (about 1.5 to about 2.0); porous MSQ (about 1.8 to about 2.2); porous PAE (about 1.8 to about 2.2). In one embodiment of the invention, the nanoporous material is NANOGLASS®, which is porous SiO ₂ . The porosity is formed of SiO ₂ , which reduces the dielectric constant to as low as about 3.9 to 1.9.

  Examples of the material having a high refractive index suitable for the present invention include tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or combinations thereof. The high k dielectric material has a refractive index of about 1.8 to about 3.0. The refractive index n describes the optical characteristics of the medium and is defined as the ratio of the speed of light in free space to the speed of light in this medium. The substrate may comprise, for example, glass, plastic or epoxy.

In one embodiment of the invention, the sensor is defined by the following parameters:

In another embodiment of the invention, the sensor structure comprises the following substance:
And is defined by the following parameters:

The simulation of the above embodiment was performed using GSolver (Grating Solver Development Co., Allen, TX) and FDTD Solutions (Lumeral Solutions, Inc., Vancouver, BC, Canada). The results shown in FIG. 2A are expected to improve bulk sensitivity to a factor of more than twice that of the conventional design. Bulk sensitivity is determined by the bulk shift factor and is defined and calculated as follows for this embodiment.

  A bulk shift factor of about 150 is obtained for both simulation and experimental data for designs that do not incorporate nanoporous materials.

  Since this proposed device works by evanescent field interaction with materials very close to the sensor surface, it is beneficial to consider the refractive index shift not only for the entire bulk medium but also for the thin layer above the sensor. . FIG. 2B shows a GSOLVER simulation, in which a 20 nm thick “biological coating” is cast by a layer having a refractive index of 1.40. Individual biological molecules or fractions of a single layer of biological molecules do not have a defined refractive index, but this biological layer is a uniform film of defined thickness for illustrative purposes. As molded.

SEM images of the patterned NANOGLASS® structure shown in FIG. 3 provide evidence of a successful imprinting process. After depositing TiO ₂ , the sensitivity of the completed sensor was confirmed by examining the resulting peak wavelength (PWV) shift captured with a reader spectrometer using deionized water and isopropyl alcohol. It was. By applying to Equation 1 using the experimental data from FIG. 4, the bulk shift factor is:
Which matches within about 5% of the value demonstrated through simulation.

  The cross-sectional profile of the sub-wavelength grating may include any periodically repeating function, for example, “square-wave”. The grid may be composed of repetitive pattern shapes such as continuous parallel lines, squares, circles, ellipses, triangles, trapezoids, sine waves, oval, rectangles and hexagons.

  The sensor may comprise a one-dimensional linear grid surface structure, i.e. a series of parallel lines or grooves. A two-dimensional grating is characterized in that the two transverse directions intersecting the plane of the sensor surface are both subwavelengths, but the cross section of a one-dimensional grating has a subwavelength only in one transverse direction and is longitudinal. The direction may be longer than the wavelength of the resonant grating effect.

  These include, for example, triangular or v-shaped, u-shaped, reverse v-shaped or u-shaped, sinusoidal, trapezoidal, stepped, and square. The grating may also be changed in height to sine.

  Another sensor structure consisting of a set of concentric rings may be used. In this structure, the difference between the inner and outer diameters of each concentric circle is equal to about half the lattice spacing. Each successive circle has an inner diameter that is approximately one lattice spacing greater than the inner diameter of the previous circle. This concentric pattern extends to cover a single sensor location—eg, a microarray spot or a microtiter plate well. Each individual microarray spot or microtiter plate well has another concentric pattern centered therein. All polarization directions of such a structure have the same cross-sectional profile. The lattice spacing of the concentric structure is shorter than the wavelength of the resonant reflected light.

  The sensor of the present invention may comprise a cover layer on the surface of the grating opposite the substrate layer. If a cover layer is present, one or more specific binding substances are immobilized on the surface of the cover layer opposite the lattice. Preferably, the cover layer comprises a material having a lower reflectivity than the material comprising the grating. The cover layer may be made of, for example, glass (including spin-on glass (SOG)), epoxy, or plastic.

  Resonant reflection can also be obtained without a planarizing cover layer covering the grating. Without a planarizing cover layer, the surrounding medium (eg air or water) fills the grid. Thus, the molecule is immobilized to the sensor not only on the top surface, but on all surfaces of the lattice exposed to this molecule.

  The present invention provides a resonant reflection structure and a transmission filter structure. For a resonant reflection structure, the light output is measured as an illumination light beam on the same side of the structure. For a transmission filter structure, the light output is measured as an illumination beam on the opposite side of the structure. The reflected signal and the transmitted signal are complementary. That is, if the wavelength is strongly reflected, it is weakly transmitted. Assuming no energy is absorbed by the structure itself, this reflected + transmitted energy is constant at a given wavelength. This resonant reflection structure and transmission filter are designed to obtain highly efficient reflection at specific wavelengths. Thus, the reflective filter “passes” the narrow band of wavelengths, while the transmission filter “cuts” the narrow band of wavelengths from the incident light.

  In one embodiment of the invention, an optical device is provided. The optical device comprises a structure similar to any sensor of the present invention; however, the optical device does not comprise one or more binding substances immobilized on the grating. The optical device can be used as a narrow band optical filter.

  An evanescent wave based sensor may comprise a waveguide film supported by a substrate; between the waveguide films (and optionally as part of the substrate) is a diffraction grating. See, for example, U.S. Pat. No. 4,815,843. A low-k dielectric material, such as a low-k nanoporous material, may be used for the diffraction grating, or a low-k nanoporous material may be used in combination with the substrate. The waveguide comprises a waveguide film and a substrate. The waveguide film may be, for example, tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or combinations thereof, or a polymer such as polystyrene or polycarbonate. The diffraction grating is present at the waveguide film and substrate interface or in the volume of the waveguide film. The diffraction grating includes a low-k material, such as a low-k nanoporous material. The refractive index of this waveguide film is higher than that of adjacent media (ie, substrate and test sample). This substrate may be, for example, plastic, glass or epoxy. Certain binding substances may be immobilized on the surface of the waveguide film and a test sample may be added to the surface. Laser light propagates in the waveguide film by total internal reflection. Changes in the refractive index of the waveguide film caused by molecular bonding to the waveguide film can be detected by observing changes in the angle of the emitted and out-coupled light.

Sensor Creation The sensor of the present invention can be produced using a plastic rubber mold for embossing a lattice structure in a nanoporous material, which is in an uncured, deformable state. It is. Unlike non-plastic solid templates, plastic rubber templates can escape the solvent vapor generated by the nanoporous material curing process. Many plastic templates can be manufactured at low cost from a single silicon wafer “master” template, and a single plastic template can be used multiple times to create many structured nanoporous subwavelength grating structures. It can be manufactured at low cost.

  The sensor can be manufactured inexpensively with a large surface area and may also be incorporated, for example, in a single-use standard disposable liquid handling manner, such as a microplate, microarray slide or microfluidic chip.

  The process path for producing a photonic crystal incorporating a nanoporous layer is outlined in FIG. Patterned “master” wafers are typically designed with silicon or glass that includes features that accurately correspond to those that are subsequently imprinted into the porous film (see FIG. 5A). The master is then used as a mold and liquid elastomer is poured into it as shown in FIG. 5B. After curing, the newly formed negative rubber “daughter” mold is carefully peeled from the master. After application of the porous film to the desired substrate, the daughter mold is set on the uncured film, for example, as shown in FIG. 5D. With the mold in place, the porous film is partially cured, fully cured, or not cured. A gas permeable rubber mold allows for solvent evaporation during this curing process. Once the film is able to retain a rigid shape, the daughter mold is peeled off to completely cure the remaining structure. The completed device illustrated in FIG. 5F is obtained by depositing a thin high refractive index material that evenly spans the patterned surface of the porous film.

  Another approach for the production of a photonic crystal biosensor incorporating a nanoporous layer is illustrated in FIG. In this structure, a layer of nanoporous material is cured on the substrate. Next, a high dielectric constant material is uniformly formed on the porous layer. The high dielectric constant material has a dielectric constant k that is about 5% higher than that of the nanoporous material. In one embodiment of the invention, the high dielectric constant material has a k greater than about 3.5. The deposited high-k material is then patterned by e-beam or DUV lithography and subsequently etched to obtain the desired features. This sensor design is not very cost-effective because it requires a high-resolution lithographic process for each device, but obtains a sensitivity enhancement similar to that seen with the above-described sensors manufactured by imprinting. Promising.

  An evanescent wave based biosensor can also be made using the same process described herein.

Specific binding substance and binding partner One or more specific binding substances may be immobilized on the lattice or, if present, the cover layer, for example by physical adsorption or chemical binding. Specific binding substances are, for example, organic molecules such as nucleic acids, polypeptides, antigens, polyclonal antibodies, monoclonal antibodies, single chain antibodies (scFv), F (ab) fragments, F (ab ′) ₂ fragments, Fv fragments Small organic molecules, cells, viruses, bacteria, polymers, peptide solutions, single or double stranded DNA solutions, RNA solutions, solutions containing compounds from combinatorial chemical libraries, or biological samples; Alternatively, it may be an inorganic molecule. Biological samples are, for example, blood, plasma, serum, gastrointestinal secretions, tissue or tumor homogenate, synovial fluid, feces, saliva, sputum, cyst fluid, amniotic fluid, cerebrospinal fluid, ascites, lung lavage fluid, semen, lymph fluid May be tears or prostate fluid.

  Preferably, one or more specific binding substances are arranged in a microarray having distinct locations on the sensor. One or more specific binding substances can bind to their specific binding partners. A microarray of specific binding substances is one on the surface of the sensor of the present invention so that each of the surfaces has many different locations with different specific binding substances or different amounts of specific binding substances. Contains one or more specific binding substances. For example, the array may have 1, 10, 100, 1,000, 10,000, or 100,000 identifiable positions. Such a sensor surface is called a microarray. This is because one or more specific binding substances are typically arranged in a regular grid pattern in xy coordinates. However, the microarrays of the present invention may include one or more specific binding substances arranged in any type of regular or irregular pattern. For example, distinct locations can define a microarray having one or more specific binding agent spots. The microarray spots may be about 50 to about 500 microns in diameter. The microarray spot may also be about 150 to about 200 microns in diameter.

  The microarray on the sensor of the present invention may be created, for example, by placing a droplet of one or more specific binding substances on the surface of the xy grid on the grid, or on the cover layer surface. When the sensor is exposed to a test sample containing one or more binding partners, the binding partner is preferably attracted to a discrete location on the microarray containing a specific binding agent that has a high affinity for the binding partner. . Some of this distinct location collects binding partners on its surface, but not other locations.

The specific binding substance specifically binds to a binding partner that is contacted with the surface of the sensor of the present invention. A specific binding substance specifically binds to its binding partner, but does not substantially bind to other binding partners contacted with the sensor surface. For example, if the specific binding agent is an antibody and its binding partner is a specific antigen, the antibody specifically binds to a specific antigen but does not substantially bind to other antigens. Binding partners include, for example, nucleic acids, polypeptides, antigens, polyclonal antibodies, monoclonal antibodies, single chain antibodies (scFv), F (ab) fragments, F (ab ′) ₂ fragments, Fv fragments, small organic molecules, cells, viruses It may be a bacterium, polymer, peptide solution, single-stranded or double-stranded DNA solution, RNA solution, solution containing a compound from a combinatorial chemical library, inorganic molecule, or biological sample.

  One example of a microarray of the invention is a nucleic acid microarray, where each distinct location in the array contains a different nucleic acid molecule. In this embodiment, the spots in the nucleic acid microarray detect complementary chemical bonds with opposite strand nucleic acids in the test sample.

  While microtiter plates are the most common method used for biochemical assays, microarrays maximize the number of biochemical interactions that can be measured at one time while minimizing the volume of expensive reagents. It is often used as a means for this. By applying a specific binding substance using a microarray spotter on the sensor of the present invention, a specific binding substance density of 10,000 specific bindings per square inch can be obtained. By focusing the illumination beam to interrogate a single microarray position, the sensor can be used as a label-free microarray reading system.

  Although the specific binding agent or binding partner need not include a detectable label, the detectable label can be used to detect a specific binding agent or binding partner on the surface of the sensor. If the specific binding agents and binding partners of the invention do not have a detection label, they can enhance assay sensitivity, immobilize specific binding partners on the biosensor surface, specific binding agents for those binding partners Other types of labels and markers may further be provided for enhanced binding or hybridization and other purposes.

Immobilized molecules of one or more specific binding substances can be immobilized on the sensor so that they are not washed away by the rinsing procedure and binding to molecules in the test sample is hindered by the sensor surface. Absent. Several different types of surface chemistry strategies have been applied to covalent attachment of molecules to glass for use in, for example, various types of microarrays and sensors. These same methods can be easily adapted to the sensor of the present invention.

  One or more types of molecules can be attached to the sensor surface by physical adsorption (ie, without the use of a chemical linker) or by chemical bonding (ie, by use of a chemical linker). Chemical binding can result in a stronger binding of the molecule to the sensor surface, providing a defined orientation and higher order structure of the molecule bound to the surface.

  Other types of chemical bonds include, for example, amine activation, aldehyde activation, and nickel activation. These materials can be used to attach several different types of chemical linkers to the sensor surface. The amine surface can be attached to several types of linker molecules, but the aldehyde surface can be used to attach proteins directly without additional linkers. A nickel surface may be used to bind molecules with an incorporated histidine (“his”) tag. Detection of “histidine tagged” molecules on nickel activated surfaces is well known in the art (Whitesides, Anal. Chem. 68, 490, (1996)).

  Immobilization of specific binding materials to plastics, epoxies or high refractive index materials can be performed essentially as described for immobilization to glass. However, the acid wash step may be omitted if such treatment can damage the substance to which this specific binding substance is immobilized.

Liquid-containing container The sensor of the present invention may comprise an inner surface, for example, a bottom surface of a liquid-containing container. The liquid-containing container may be, for example, a microtiter plate well, a test tube, a petri dish or a microfluidic channel. One embodiment of the present invention is a sensor that is incorporated into any type of microtiter plate. For example, the sensor may be incorporated into the bottom surface of the microtiter plate by providing a reaction vessel wall on the resonant reflective surface so that each reaction “spot” is separated into a separate test sample. Can be exposed. Thus, each individual microtiter plate well can function as a separate reaction vessel. Thus, separate chemical reactions can occur in adjacent wells without internal mixing of the reaction solutions, and chemically separate test solutions can be added to individual wells.

  The most common assay format for pharmaceutical high-throughput screening laboratories, molecular biology research laboratories and diagnostic assay laboratories is the microtiter plate. The plate is a standard size plastic cartridge and may contain 96, 384 or 1536 individual reaction vessels arranged in a grid. Because of the standard mechanical configuration of these plates, liquid dispensing, robotic plate processing, and detection systems are designed to work in conjunction with this general scheme. The sensor of the present invention may be incorporated into the bottom of a standard microtiter plate. Since the sensor surface can be manufactured in a large area and the reading system does not make physical contact with the sensor surface, the focus resolution of the illumination optics and the xy scanning illumination / detection probe across the sensor surface Any number of individual sensor areas can be defined that are limited only by the stage.

  Sensors may also be incorporated into other disposable laboratory assay formats such as microarray slides, flow cells and cell culture plates. For compatibility with existing microarray handling equipment such as spotters and incubation chambers, it is desirable to incorporate the sensor into a common laboratory format.

Methods of using the sensors of the present invention can be used , for example, to study one or multiple molecules / molecular interactions in parallel; for example, one or more specific binding agents, their Binding to each binding partner can be detected without the use of a label by applying one or more binding partners to a sensor that has one or more specific bindings immobilized on its surface. . The sensor is illuminated with light, and the maximum reflected light wavelength or the minimum transmitted light wavelength is detected from the sensor. When one or more specific binding substances bind to their respective binding partners, the wavelength of the reflected light is such that the one or more specific binding substances are not bound to their respective binding partners. Compare and shift. When the sensor is coated with an array of discrete locations containing one or more specific binding substances, the maximum wavelength of reflected light or the minimum wavelength of transmitted light is detected from each discrete location of the sensor.

  In one embodiment of the invention, various specific binding substances, such as antibodies, may be immobilized in an array format on the sensor of the invention. The sensor is then contacted with a test sample of interest that includes a binding partner, eg, a protein. Only proteins that specifically bind to the antibody immobilized on the sensor remain bound to the sensor. Such an approach is essentially a large version of an enzyme-linked immunoassay; however, the use of enzymes or fluorescent labels is not necessary.

  The activity of the enzyme can be detected by detecting the wavelength of reflected light from a sensor to which one or more specific binding substances are immobilized, and adding one or more enzymes to the sensor. Clean the sensor and irradiate it with light. The wavelength of the reflected light is detected from this sensor. When one or more enzymes alter one or more specific binding substances of the sensor due to enzymatic activity, the wavelength of the reflected light is shifted.

  In addition, a test sample, eg, a cell lysate containing a binding partner, can be added to the sensor of the invention, followed by washing to remove unbound material. The binding partner bound to the sensor can then be eluted from the sensor and identified, for example, by a mass spectrometer. If necessary, a phage DNA display library may be applied to the sensor of the present invention, followed by washing to remove unbound material. Individual phage particles bound to the sensor can be isolated and then the inserts in these phage particles can be sequenced to determine the type of binding partner.

  For the above applications, and in particular for proteomics applications, the ability to selectively bind substances such as binding partners from the test sample onto the sensor of the present invention, followed by the separation of this sensor for further analysis. The ability to selectively remove binding material from the position is advantageous. The sensor of the present invention can also detect and quantify the amount of binding partner from the sample that is bound to a distinct location in the array of sensors by measuring the shift in wavelength of the reflected light. In addition, the wavelength shift at one distinct sensor location is compared to the positive and negative controls at the other distinct sensor location to determine the amount of binding partner bound to the distinct location of the sensor array. May be.

  In one embodiment of the invention, the interaction between the first molecule and the second test molecule can be detected. A sensor as described above is used except that there is no specific binding substance immobilized on its surface. Thus, the sensor includes a one or two dimensional grating, a substrate layer that supports the one or two dimensional grating, and optionally a cover layer. As described above, when the sensor is illuminated, a resonant grating effect is produced on the reflected radiation spectrum, and the depth and spacing of the grating is smaller than the wavelength of the resonant grating effect.

In order to detect the interaction between the first molecule and the second test molecule, a mixture of the first and second molecules is applied to separate locations on the sensor. The discrete location may be a spot or well on the sensor or a large area on the sensor. The mixture of the first molecule and the third control molecule may be added to a separate location on the sensor. This sensor may be the same sensor as described above or may be a second sensor. If this sensor is the same sensor, a second distinct location may be used for the mixture of the first molecule and the third control molecule. Alternatively, the same separate sensor location may be used after washing the first and second molecules from the sensor. The third control molecule does not interact with the first molecule and is approximately the same size as the first molecule. Measure the shift in wavelength of the reflected light from a separate location of the sensor (s). The shift in the wavelength of reflected light from separate locations with the first molecule and the second test molecule is greater than the shift in wavelength of reflected light from separate locations with the first molecule and the third control molecule. The first test molecule interacts with the second test molecule. The interaction may be, for example, hybridization of nucleic acid molecules, specific binding of antibodies and antibody fragments to antigens, and binding of polypeptides. The first molecule, the second test molecule, or the third control molecule can be, for example, a nucleic acid, polypeptide, antigen, polyclonal antibody, monoclonal antibody, single chain antibody (scFv), F (ab) fragment, F (ab ') ₂ fragments, Fv fragments, small organic molecules, cells, viruses and bacteria.

  All patents, patent applications and other scientific or technical documents referred to anywhere herein are incorporated by reference in their entirety. The methods and compositions described herein are representative of presently preferred embodiments, are exemplary, and are not intended as limitations on the scope of the invention. Variations and other uses herein will be apparent to those skilled in the art and are encompassed within the scope of the invention. The invention as described diagrammatically as appropriate herein may be practiced without any element (s) or limitation (s) not specifically disclosed herein. Thus, for example, in each case herein, the terms “comprising,” “essentially comprising,” and “consisting of” are interchanged with any of the other two terms. May be used. The terms and expressions used are used as descriptive terms, not limitations, and exclude any equivalents or parts thereof that are features shown and described in the use of such terms and expressions. However, it will be appreciated that various modifications are possible within the scope of the claimed invention. Thus, while the invention has been disclosed in detail by way of embodiments and optional features, modifications and variations of the concepts disclosed herein will be defined in the detailed description and appended claims. In particular, it should be understood that it is intended to be within the scope of the present invention.

  Further, if a feature or aspect of the invention is described with respect to a Markush group or other alternative classification, one of ordinary skill in the art will also recognize that the invention also includes any individual member or member subgroup of the Markush group or other group With respect to those described herein.

Example 1
Resonance wavelength and bulk refractive index sensitivity of one-dimensional surface photonic crystal biosensors using computer- coupled coupled wave analysis (RCWA) and time-difference time domain (FDTD) simulations Predicted. This device incorporates a low refractive index (n = 1.17) nanoporous dielectric surface structure in place of the previously reported polymer (n = 1.39) surface structure. Use a soft contact embossing method to create a surface-structured low refractive index porous film on a glass substrate that has the same depth and spacing as the previous polymer structure, allowing sensitivity to be compared in parallel I made it. The sensitivity of the porous glass biosensor was compared to the nanoporous polymer biosensor by a method that characterizes the sensitivity to bulk refractive index and surface adsorbents. Finally, protein binding assay comparisons were performed to demonstrate sensor stability and ability to functionalize the device for selective detection.

  The polymer and porous glass sensor was molded and simulated using two software packages. First, a 2-D diffraction grating analysis tool (GSOLVER) using the RCWA algorithm provides a quick and simple method for initial sensor modeling. Second, FDTD (Lumeral) provides a much more diverse and powerful tool that can be used for any optical device illuminated by any light source, at any temporary or spatial location, at any electrical field component. Can be calculated. See, for example, Kunz & Luebbers, The Finite Difference Time Domain Method for Electromagnetics. 1993, Boca Raton: CRC Press. FDTD was used to validate RCWA results and gain deeper insight into the effect of modifying sensor structure.

Both RCWA and FDTD simulations showed that replacing the patterned UV-cured polymer of conventional devices with a low refractive index material resulted in a 2-fold increase in bulk shift coefficient. The resonant wavelength of the porous glass sensor immersed in DI H ₂ O was predicted to be 844.3 nm by RCWA and a full width at half maximum (FWHM) of about 2 nm, as shown in FIG. Simulations predict that the bulk shift factor can be further improved with minor modifications to the sensor shape.

Bulk sensitivity tests using DI H ₂ O and IPA were performed with 23 porous glass sensors and 11 polymer sensors. The average PWV shift was 13.6 ± 2.4 nm and 5.1 ± 1.5 nm for the porous glass and polymer sensors, respectively. The bulk shift coefficient (ΔPWV / Δn) of the porous glass sensor is measured to be 2.7 ± 1.2 times greater than that of the polymer device. Measurement of porous glass devices in DI H ₂ O yields an average PWV of 829.5 ± 16.5 nm and FWHM of 3.5 ± 2.5 nm. One of the measured spectra is shown in FIG. 9, where the response is normalized as a perfect reflection to account for any device loss. The lower reflection efficiency and wider FWHM measured from the replicated device are believed to be due to small but measurable material loss and imperfections observed in this replicated structure. The large variation in measured spectral features is due, at least in part, to using several slightly different (but nominally identical) master patterns and lack of automation of the replication process.

Example 2
Sensor manufactured sol-gel derived low-index nanoporous silica thin films (see, eg, US Pat. No. 6,395,651) are used in conventional designs for UV. Incorporated into sensor instead of cured epoxy. Since low refractive index materials are cured by heat rather than UV exposure, new manufacturing processes need to be developed. While it is desirable to retain a low cost imprinting method, it has been apparent that plastic substrates cannot retain the high temperature requirements for porous glass annealing. One possible approach to sol-gel glass imprinting has been to use polydimethylsiloxane (PDMS) molds and glass substrates. See, for example, Parashar et al., Nano-replication of diffractive optical elements in sol-derived glasses. See Microelectronic Engineering, 2003.67-8: 710-719.

The sub-wavelength grating structure of the low-k biosensor was fabricated using a combination of lithography, molding and imprinting processes. First, a Sylgard 184 PDMS (Dow Corning) daughter mold is replicated from a silicon master wafer patterned with a positive image of the desired surface structure in the completed sensor. In order to facilitate the release of the PDMS mold from the silicon wafer, the wafer was surface treated with a release layer of dimethyldichlorosilane (Repel Silane, Amersham Biosciences). See, for example, Beck et al., Improving stamps for 10 nm level wafer scale nanoimprint lithography. See, Microelectronic Engineering, 2002.61-2: 441-448. The PDMS replica is then used to imprint a thin film of uncured NANOGLASS® (Honeywell Elec. Mat.), A low refractive index sol-gel glass that is spun-on to a glass substrate. After the low index dielectric material is cured, the plastic PDMS mold is removed and the sol-gel glass is fully cured by further baking. The sensor structure is completed by depositing 175 nm TiO ₂ on the patterned surface. Subsequent surface treatment with dimethyldichlorosilane enhances biomaterial adsorption and promotes sensor stability. A diagram showing a cross section of the device is shown in FIG.

The polymer structure is similar to that described in conventional publications. See, for example, Cunningham, et al., A plastic colorimetric resonant optical biosensor for multiple parallel detection of label-free biochemical interactions. See Sensors and Actuators B, 2002.85: 219-226. Both structures use 550 nm spacing and 170 nm imprint depth, while polyester / polymer and low index porous glass devices use 120 nm and 165 nm TiO ₂ coatings, respectively. The two devices are referred to as “polymer” and “porous glass” sensors throughout the remainder of this example. The polymer device is provided as an array of aligned sensors and attached to a bottomless 96-well standard microtiter plate (SRU Biosystems). The porous glass device is manufactured on a microscope glass slide of 75 mm × 25 mm × 1 mm. Adhesive rubber wells (Research International Corp.) are bonded to the glass surface to obtain liquid containment for 5-6 sensors on each slide.

Deionized water (DI H ₂ O, n = 1.333) and isopropyl alcohol (IPA, n = 1.378) were used to determine the bulk shift factor for each sensor. First, DI H ₂ O was pipetted on the surface of the sensor, and PWV was measured. The configuration of the reader has been reported previously. See, for example, Cunningham et al., Colorimetric resonant reflection as a direct bioassay assay technique. See Sensors and Actuators B, 2002.81: 316-328. A light source with a wide range of wavelengths is coupled to the optical fiber to illuminate a region of about 2 mm diameter on the surface of the photonic crystal from below the substrate with normal incidence. The reflected light is collected by a second optical fiber, which is then focused on the illumination fiber and measured by a spectrometer. The automated motion stage allows reflection data to be collected in parallel at timed intervals from many wells to obtain motion information.

Next, the surface was well dried and the previous step was repeated for IPA. The bulk shift coefficient between DI H ₂ O and IPA was then calculated as the change in PWV divided by the change in bulk refractive index.

Example 3
PPL Biomaterial Adhesion Test Poly (Lys, Phe) (PPL) prepared in a 0.5 mg / ml solution using 0.01 M phosphate buffered saline (PBS; pH = 7.4) applied to the sensor surface Sigma-Aldrich; MW = 35,400 Da) was characterized by sensitivity to surface adsorbents by detection of monolayer films. The bio-adhesion test started with pipetting PBS into the test well and was performed at a sampling interval of 1 minute. After 10 minutes, the buffer was replaced with PPL solution and allowed to stabilize for 30 minutes. The wells were then washed 3 times and filled with PBS, and data was acquired during the last 30 minutes.

  PPL was deposited on 5 porous glasses and 9 polymer sensors. FIG. 10 compares the kinetic plots of each device, which shows an approximately 4-fold increase in surface sensitivity for the porous glass sensor. The first stage establishes a baseline, the second corresponds to rapid surface adsorption and PPL saturation, and the last third curve is weakly or non-specifically bound by rinsing with PBS buffer The stability of the monolayer after exclusion of molecules is shown. The PWV shift generated during PPL immobilization on the porous glass sensor saturates more slowly than that measured using the polymer device. Clearly, this porous glass sensor surface is significantly less conductive for protein monolayer adsorption. Further optimization of the surface chemistry should reduce this effect. Nevertheless, porous glass sensors exhibit excellent stability after unbound molecules are washed away.

Example 4
To characterize the discrimination sensitivity as a function of distance from the multilayer polymer test sensor surface, a series of polymer electrolyte monolayers were deposited on the sensor. By alternately using a positively charged polymer layer and a negatively charged polymer layer while continuously monitoring on the detection device, a uniform self-limited polymer stack can be formed on the sensor. See, for example, Cunningham et al., Enhancing the surface sensitive of chromatic resonant optical biosensors. See Sensors and Actuators B, 2002.87 (2): 365-370. Three different polyelectrolytes were deposited on the sensor surface: anionic poly (sodium 4-styrenesulfonate) (PSS; MW = 70,000 Da), cationic poly (ethyleneimine) (PEI; MW = 60,000 Da), and cationic poly (allylamine hydrochloride) (PAH; MW = 70,000 Da). The polyelectrolyte was purchased from Sigma-Aldrich. A 0.9M NaCl buffer solution (Sigma-Aldrich) was prepared with deionized water. This polyelectrolyte was dissolved in a buffer solution at a concentration of 5 mg / ml. At a sampling interval of 1 minute, the sensitivity of the multilayer surface was characterized in 5 minute steps. First, NaCl buffer was pipetted into the sensor wells. Next, the buffer solution was taken out and the PEI solution was added. The wells were then washed 3 times and filled with buffer. The above two steps were repeated with PSS and PAH until seven PSS-PAH layers were deposited on a single PEI layer.

  The 14 alternating layers of PSS and PAH described above each produce a measurable shift in the detected PWV. Because they are adsorbed on the surface. FIG. 11 shows the spatial profile of PWV shift versus polymer thickness, where each PWV shift was measured in buffer after the washing step. Each monolayer of polyelectrolyte is about 4.4 nm thick and has a refractive index of 1.49. For example, Picart et al., Determination of structural parameters characterizing thin films by optical methods. See Journal of Chemical Physics, 2001.115 (2): 1086-1094. Porous glass sensors exhibit an average surface sensitivity of about 1.5 times that of polymer sensors. However, each of the first two layers (about 9 nm) deposited on the porous glass device results in a PWV shift that is twice the size of each of the remaining layers, while this effect is Note that it is not observed on the device.

Example 5
Bioassay: Protein A-IgG
In order to demonstrate selective detection with the proposed device, a bioassay characterizing the affinity of human, sheep and chicken IgG for protein A was performed. Protein A (Pierce Biotechnology) was prepared at a concentration of 0.5 mg / ml using 0.01 M PBS. The buffer was filtered before use using a 0.22 μm syringe filter (Nalgene). Human, sheep and chicken immunoglobulin-G (IgG) serum (Sigma-Aldrich) was diluted to a concentration of 0.5 mg / ml in 0.01 M PBS. The PBS solution was first pipetted into the sensor wells at 30 minutes between each step and 1 minute sampling. Next, this buffer solution was replaced with a protein A solution. The wells were then rinsed 3 times and filled with buffer. After signal stabilization, the buffer in the three wells was replaced with human, sheep or chicken IgG, but the fourth was left as a reference containing buffer only. Finally, the IgG was removed and the wells were rinsed again and filled with PBS, and data was acquired during the last 30 minutes.

  Protein A was introduced into 15 porous glasses and 16 polymer sensor wells. The PWV shift obtained after the cleaning process was about 4 times greater for porous glass devices. FIG. 13 illustrates the measured binding kinetics of human, sheep and chicken IgG with protein A for the porous glass sensor, while FIG. 14 illustrates the endpoint PWV between the two devices for each antibody. A shift comparison (compared to a reference well without IgG) is shown. Protein A surface adsorption saturated quite rapidly on the polymer sensor surface, similar to that observed in the PPL biomaterial adhesion test. Porous glass devices show greater sensitivity for antibodies with higher affinity for protein A. Human IgG binding was detected with a 2-fold sensitivity, but lacks specificity for protein A (see, for example, Richman et al., The binding of Staphylococci protein A by the sera of diffractive animal specialties. 2 .128: 2300-2305) produces an equivalent response and gives a measure of non-specific binding.

  Photonic crystal biosensors are designed to couple electromagnetic energy with biological material from a liquid test sample deposited on its surface. The device itself is composed of a low refractive index surface structure and a high refractive index dielectric coating, but the liquid test sample that fills the surface structure is also the only part of the sensor that can induce changes in the resonant wavelength and resonance wavelength. Must be considered a dynamic component. The motivation for incorporating an excessively low refractive index material into the photonic crystal biosensor structure interacts more strongly with the liquid test sample and weakly interacts with the internal region of the photonic crystal that does not interact with the surface adsorbent It is to bias the electromagnetic field of the resonance wavelength.

  The use of spin-on low-k dielectric materials eliminates significant investment in the integrated circuit manufacturing industry that requires rapid processing, structural stability and elimination of liquid penetration. A unique aspect of this study is the use of imprinting methods to accurately give sub-μm surface structures to nanoporous glass films without the use of photolithography. The presence of an imprint tool on the surface of the low-k film during the initial stage of the curing process did not change the refractive index of the final cured structural film. This imprinting method incurs substantial costs only in the creation of “master” silicon wafers, which can then be used to create an almost unlimited number of “daughter” PDMS imprinting tools. Each PDMS tool can be used to create multiple sensor structures without damaging the tool. This is because little force is required to create a spun-on liquid low-k layer that fits this tool. After imprinting, the low-k dielectric film is rapidly cured on a hot plate using an easily automated method. The use of a plastic imprinting tool has been found to be advantageous over direct imprinting from a silicon master wafer. This is because PDMS molds are easy to peel off from a partially cured low-k film and allow penetration of volatile solvents released during the curing process. Although the present study presented here imprinted only a 1 × 3 inch microscope slide area, this imprinting method covers an entire 96 or 384 well standard microplate (approximately 3 × 5 inches). Can be scaled up to a larger surface area that allows the creation of a sensor area size sufficient.

  An interesting and useful result found during the comparison between porous glass sensor structures and polymer sensor structures is the sensitivity gain imbalance between bulk refractive index sensitivity and surface adsorbed layer sensitivity. The computer model accurately predicts an approximate 2-fold increase in sensitivity measured for the PWV shift induced by the bulk refractive index change of the solution over the porous glass sensor surface, but for a thin layer of adsorbed material, the PWV An approximately 4-fold increase in shift was consistently measured. By measuring the PWV shift as a function of thickness using a polymer multilayer experiment (FIG. 11), we characterize the strength of the coupled electromagnetic field interaction as a function of distance from the sensor surface. Is possible. For porous glass sensors, the interaction is particularly strong for the first few monolayers of adsorbed polymer, but the relationship between polymer thickness and PWV is different for each of the polymer sensor structures. It is very linear for the adsorbed monolayer. The interaction between the test sample and the resonant electromagnetic field distribution is extremely complex. This is because the substance to be detected can be adsorbed to the horizontal and vertical planes of this structure, where a characteristic electromagnetic field profile extends from each surface into the sample. Surface-based detection sensitivity is enhanced beyond improving the bulk sensitivity of porous glass biosensors. Since most biomolecule interactions are expected to occur within the first 2-3 nanometers from the sensor surface, surface sensitivity is most important for increasing sensitivity in the context of surface-based biochemical assays. It is a thing.

FIG. 1 shows a diagram of a nanoreplicated nanoporous photonic crystal biosensor. FIG. 2 shows (A) bulk and (B) surface shift inferred by GSolver simulation. FIG. 3A shows an SEM image of an imprinted and cured periodic NANOGLASS® structure. FIG. 3B shows an SEM image of the imprinted and cured periodic NANOGLASS® structure. FIG. 4 shows the experimental response of the nanoporous sensor when immersed in deionized water (DI) and isopropyl alcohol (IPA). FIG. 5A shows a process path for manufacturing a nanoporous sensor. FIG. 5B shows a process path for manufacturing a nanoporous sensor. FIG. 5C shows a process path for manufacturing a nanoporous sensor. FIG. 5D shows a process path for manufacturing a nanoporous sensor. FIG. 5E shows a process path for manufacturing a nanoporous sensor. FIG. 5F shows a process path for manufacturing a nanoporous sensor. FIG. 6 shows a diagram of a high dielectric constant nanoporous photonic crystal sensor. FIG. 7 shows a cross-sectional view of a porous glass sensor. FIG. 8 shows the resonance peak of a porous glass sensor exposed to deionized water, as estimated by RCWA simulation. FIG. 9 shows the experimentally measured resonance peak of a nanoporous glass sensor immersed in deionized water. FIG. 10 shows a kinetic plot comparing PWV shifts for PPL deposited on both porous glass and polymer sensor designs. FIG. 11 shows a partial profile of PWV shift versus polymer thickness where alternating layers of PSS and PAH contribute to the total measured shift. FIG. 12 shows a kinetic plot comparing PWV shifts for protein A deposited on both porous glass and polymer sensor designs. FIG. 13 shows the binding kinetics of three animal IgGs to protein A measured with a nanoporous glass sensor. FIG. 14 shows a sensor comparison of PWV shift for each of the various IgG-Protein A interactions.

Claims (31)

A sensor comprising a nanoporous material having a low refractive index, supported on a bottom surface by a substrate, and coated on a top surface with a high dielectric constant dielectric coating;
The high dielectric constant dielectric coating or the high dielectric constant dielectric coating combined with the nanoporous material forms a sub-wavelength grating structure;
When the sensor is illuminated, a resonant grating effect occurs in the reflected radiation spectrum; and the depth and spacing of the sub-wavelength spacing grating structure is smaller than the wavelength of the resonant grating effect.   The sensor according to claim 1, wherein when the sensor is irradiated with a wide band of optical wavelengths, the narrow band of optical wavelengths is reflected from the sensor.   The sensor of claim 1, wherein the refractive index of the nanoporous material is from about 1.1 to about 2.2.   The sensor of claim 1, wherein the refractive index of the porous material is from about 1.1 to about 1.5.   The sensor of claim 1, wherein the spacing of the sub-wavelength spacing grating structure is from about 50 nm to about 1,500 nm, and the depth of the sub-wavelength spacing grating structure is from about 50 nm to about 900 nm.   The nanoporous material is porous silica xerogel, porous airgel, porous hydrogen silsesquioxane, B-stage polymer, porous methylsilsesquioxane, porous poly (arylene ether) or a combination thereof. The sensor according to claim 1.   The sensor of claim 1, wherein the substrate comprises glass, plastic, or epoxy.   The sensor of claim 1, wherein the refractive index of the dielectric coating is about 1.8 to about 3.0.   The sensor of claim 1, wherein the dielectric coating comprises tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or combinations thereof.   The sensor of claim 1, wherein the refractive index of the substrate is from about 1.4 to about 1.6.   The sensor of claim 1, wherein the dielectric coating has a thickness of about 30 nm to about 700 nm, and the nanoporous material has a thickness of about 10 nm to about 5,000 nm.   The sensor of claim 1, wherein the dielectric coating has a cover layer on its top surface.   The sensor of claim 1, wherein the sensor further comprises one or more specific binding substances immobilized on the high dielectric constant dielectric coating.   The sensor of claim 12, wherein the sensor further comprises one or more specific binding substances immobilized on the cover layer.   14. A sensor according to claim 13, wherein the one or more specific binding substances do not comprise a detectable label.   14. A sensor according to claim 13, wherein the one or more specific binding substances bind to their binding partners.   17. A sensor according to claim 16, wherein the one or more specific binding substances and the binding partner do not comprise a detectable label.   14. The sensor of claim 13, wherein the one or more specific binding substances are aligned in an array on the high dielectric constant dielectric coating.   The sensor of claim 14, wherein the one or more specific binding substances are aligned in an array on the cover layer.   A sensor comprising a waveguide structure formed by a waveguide film covering a substrate, wherein the waveguide film has a refractive index higher than that of the substrate, and a diffraction grating is in the waveguide structure. And the diffraction grating is composed of a nanoporous material having a low dielectric constant.   21. The sensor of claim 20, wherein the nanoporous material has a refractive index of about 1.1 to about 1.5.   21. The sensor of claim 20, wherein the nanoporous material has a refractive index of about 1.1 to about 2.2.   The nanoporous material is porous silica xerogel, porous airgel, porous hydrogen silsesquioxane, B-stage polymer, porous methylsilsesquioxane, porous poly (arylene ether) or a combination thereof. The sensor according to claim 20.   The sensor of claim 20, wherein the substrate comprises glass, epoxy, or plastic.   21. The sensor of claim 20, wherein the waveguide film comprises tin oxide, tantalum pentoxide, zinc sulfide, titanium dioxide, silicon nitride, or combinations thereof.   21. The sensor of claim 20, wherein the waveguide film comprises a polymer.   21. The sensor of claim 20, wherein the sensor further comprises one or more specific binding substances immobilized on the waveguide film.   28. The sensor of claim 27, wherein the one or more specific binding substances do not comprise a detectable label.   30. The sensor of claim 28, wherein the one or more specific binding substances bind to their binding partners.   30. The sensor of claim 29, wherein the one or more specific binding substances and the binding partner do not comprise a detectable label. 28. The sensor of claim 27, wherein the one or more specific binding substances are arrayed on the high refractive index dielectric coating.